U.S. patent application number 17/543773 was filed with the patent office on 2022-03-24 for non-uniformly stiff polymeric scaffolds and methods for producing thereof.
The applicant listed for this patent is Regenosca SA. Invention is credited to Eva-Maria Balet, Peter Frey, Jeffrey Alan Hubbell, Hans Mattias Larsson, Kalitha Pinnagoda, Elif Vardar, Ganesh Vythilingam.
Application Number | 20220088270 17/543773 |
Document ID | / |
Family ID | |
Filed Date | 2022-03-24 |
United States Patent
Application |
20220088270 |
Kind Code |
A1 |
Frey; Peter ; et
al. |
March 24, 2022 |
Non-Uniformly Stiff Polymeric Scaffolds and Methods for Producing
Thereof
Abstract
The invention relates to methods for producing a polymeric
scaffold for use in tissue engineering applications or soft tissue
surgery, as well as to the produced scaffolds and an associated
kit. The method features a first fast drying step of applying a
mechanical compression on a polymeric gel layer and a second slow
drying step of the gel up to reach a polymer mass fraction of at
least 60% w/w in the final scaffold. The method allows the
production of scaffolds with high regeneration and healing
properties of a grafted tissue via host cell invasion and
colonization, and a good suturability. These goals are achieved
through the formation within the scaffold of a non-uniform
architecture creating softer and stiffer areas, which is maintained
even upon re-swelling of the scaffold upon hydration of the final
dried product.
Inventors: |
Frey; Peter; (Epalinges,
CH) ; Hubbell; Jeffrey Alan; (Chicago, IL) ;
Larsson; Hans Mattias; (Lausanne, CH) ; Vardar;
Elif; (Lausanne, CH) ; Balet; Eva-Maria;
(Cressier, CH) ; Vythilingam; Ganesh; (Lausanne,
CH) ; Pinnagoda; Kalitha; (St-Legier, CH) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Regenosca SA |
Lausanne |
|
CH |
|
|
Appl. No.: |
17/543773 |
Filed: |
December 7, 2021 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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16092953 |
Oct 11, 2018 |
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PCT/IB2017/051796 |
Mar 29, 2017 |
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17543773 |
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International
Class: |
A61L 27/36 20060101
A61L027/36; A61L 27/20 20060101 A61L027/20; A61L 27/22 20060101
A61L027/22; A61L 27/24 20060101 A61L027/24; A61L 27/52 20060101
A61L027/52; A61L 27/54 20060101 A61L027/54; C08J 3/075 20060101
C08J003/075; C08L 89/00 20060101 C08L089/00 |
Foreign Application Data
Date |
Code |
Application Number |
Apr 13, 2016 |
IB |
PCT/IB2016/052096 |
Claims
1-25. (canceled)
26. A polymeric scaffold having a non-uniform density for use in
tissue engineering, diagnostics, or surgical procedures, wherein
the polymeric scaffold has a polymer mass fraction of at least 60%
w/w of a total scaffold weight, and wherein the polymeric scaffold
has a density at a surface that is denser as compared to a density
at a core.
27. The polymeric scaffold of claim 26, wherein a pore diameter at
the core is between 0.2 .mu.m and 200 .mu.m.
28. The polymeric scaffold of claim 26, wherein an average pore
diameter at the surface is less than 0.2 .mu.m.
29. The polymeric scaffold of claim 26, wherein a Young's modulus
at the core is between 0.1 kPa and about 20 kPa.
30. The polymeric scaffold of claim 26, wherein a Young's modulus
at the surface is between 5 kPa and 1500 kPa.
31. The polymeric scaffold of claim 26, wherein the polymeric
scaffold is suturable.
32. The polymeric scaffold of claim 26, wherein the polymeric
scaffold is at least 1 cm long.
33. The polymeric scaffold of claim 26, wherein the polymeric
scaffold includes a bioactive agent.
34. The polymeric scaffold of claim 26, wherein the polymeric
scaffold is cell-free.
35. The polymeric scaffold of claim 26, wherein the polymeric
scaffold includes cells.
36. The polymeric scaffold of claim 26, wherein the polymeric
scaffold is immersed in an aqueous solution.
37. The polymeric scaffold of claim 26, wherein the polymeric
scaffold does not include a crosslinking agent.
38. The polymeric scaffold of claim 26, wherein the polymeric
scaffold is configured to reswell in an aqueous solution by more
than 4% w/w to reach a steady-state water content.
39. The polymeric scaffold of claim 38, wherein the polymeric
scaffold is configured to reswell in a range of 5% to 48% w/w to
reach a steady-state water content.
40. The polymeric scaffold of claim 26, comprising one or more
polymeric material layers.
41. The polymeric scaffold of claim 26, comprising one or more
polymeric tubes.
42. The polymeric scaffold of claim 26, wherein a polymeric
material of the polymeric scaffold includes gelatin, elastin,
collagen, agar/agarose, chitosan, fibrin, methyl cellulose,
hyaluronic acid, proteoglycans, polyamino-acids or derivative
thereof, polysaccharides or derivatives thereof, glycosaminoglycans
or derivatives thereof, or any combination of the foregoing.
43. An implant kit comprising: a container; an aqueous solution
located inside the container; a polymeric scaffold immersed by the
aqueous solution inside the container, wherein the polymeric
scaffold has a polymer mass fraction of at least 60% w/w of a total
scaffold weight, and the polymeric scaffold has a density at a
surface that is denser as compared to a density at a core.
44. A graft made of a polymeric scaffold having a non-uniform
density for surgical applications, wherein the polymeric scaffold
has a polymer mass fraction of at least 60% w/w of a total scaffold
weight, and wherein the polymeric scaffold has a density at a
surface that is denser as compared to a density at a core.
45. The polymeric scaffold of claim 44, wherein the polymeric
scaffold is configured to reswell in an aqueous solution by more
than 4% w/w to reach a steady-state water content.
Description
TECHNICAL FIELD
[0001] The invention relates to tissue engineering, diagnostics and
surgery, particularly to non-uniformly stiff polymeric scaffolds
and methods for producing such objects.
BACKGROUND ART
[0002] Collagen is one of the first natural polymers used as a
biomaterial. Being the most abundant structural extracellular
matrix protein in the human body, it is therefore widely used as a
scaffold component for tissue engineering applications, and
considered among the first choices for natural scaffold
materials.
[0003] Collagen gels are formed by a process called
fibrillogenesis. Fibrillogenesis starts when a diluted acidic
collagen solution is adjusted in vitro to physiological pH and
temperature (Gross and Kirk, J. Bio. Chem 233 (1958): 355-360).
Collagen gels are interesting as scaffolds due to their good
biocompatibility, their low immunogenicity and their easy of
remodeling by cells. However, the large excess of fluid entrapped
between the collagen fibers render them inherently weak and thus
difficult to manipulate. Several strategies were proposed to obtain
collagen scaffolds having improved mechanical properties. The
simple culturing of cells within collagen gels results in the
expulsion of some of the interstitial fluid by gel contraction, but
cell compaction is a not an easily controllable process and the
increase in collagen density is linked to a reduction of the
scaffold surface (i.e. scaffold shrinkage). Cross-linking of the
collagen fibers enhances the mechanical properties of collagen
scaffolds without changing the scaffold surface. Physical, chemical
and biological collagen cross-linking methods are well reported in
literature (Paul and Bailey, The Scientific World Journal 3 (2003):
138-155). One option is the lyophilization (freeze-drying) of an
aqueous collagen solution followed by chemical cross-linking
(EP1561480). Another possibility is freeze-drying of a collagen gel
followed by thermal dehydration cross-linking (EP1098024).
[0004] The most common aldehyde cross-linking agent is
glutaraldehyde. However its use for medical applications is
questioned due to its cytotoxicity. Several chemical cross-linking
agents have been identified that achieve comparable cross-linking
levels to that of glutaraldehyde while improving the
cytocompatibility (i.e. diphenylphosphoryl azide,
1-ethyl-3-(3-dimethylaminopropyl)carbodiimide, genipin) (Petite H.
et al., Biomaterials 16 (1995): 1003-1008).
[0005] Another method to produce collagen gels with enhanced
mechanical properties consists of applying plastic compression.
During the process of plastic compression, the interstitial liquid
is expelled from a collagen gel and does not return on removal of
the load. This simple, fast, controllable, and cell-friendly
process to engineer collagen matrices with suitable mechanical
properties for soft tissue engineering was described in detail by
Brown and colleagues (WO 2006/003442; WO 2012/004564; Abou Neel E.
A. et al., Soft Matter 2 (2006): 986-992; Brown R. A. et al., Adv
Funct Mater 15 (2005): 1762-1770). Furthermore, collagen hybrid
structures have been reported where a synthetic polymer component
provided the mechanical support (Kawazoe N. et al., Biotechnol Prog
26 (2010): 819-826; Lu H. et al., Sci Technol Adv Mater 13 (2012)).
However, this approach decreases the biocompatibility of the
matrix.
[0006] Most scaffold design strategies aim to mimic the mechanical
properties of the native tissue as much as possible while also
providing a scaffold that is easy to handle and suturable. The
following suturable collagen matrices were reported in literature:
collagen sponge or collagen/glycosaminoglycan sponge crosslinked
with collagen gel (U.S. Pat. No. 5,567,806), double layered
collagen matrix resulting from a casting, freeze-drying, and
lyophilisation procedure where the collagen layers differ in
density (WO 2007/082295), and a high density freeze-dried collagen
matrix (WO 2000/029484).
[0007] Furthermore it is well known that the mechanical properties
of a scaffold influences cell behaviour. Myogenesis is best
achieved on collagen scaffolds that are moderately stiff (Engler A.
J. et al., J Cell Biology 166 (2004): 877-887) while osteogenesis
is achieved on stiff collagen scaffolds (Huebsch N. et al., Nat
Matter 9 (2010): 518-526). Researchers can change mechanical
properties of collagen hydrogels by varying collagen concentration
and using different cross-linking methods. However, these scaffolds
are uniform matrices and a soft collagen scaffold showing excellent
cell behaviour in term of proliferation and differentiation might
not be suitable for a clinical application (i.e. not suturable,
difficult to handle).
SUMMARY OF INVENTION
[0008] In order to address and overcome the drawbacks associated
with the prior art solutions concerning polymeric products for
tissue engineering and surgery, the inventors developed a brand new
method for producing polymeric scaffolds displaying several
advantages when used in an in vivo approach. Particularly, the
scaffold of the present invention features superior properties in
terms of resistance upon suture when grafted in vivo in/on soft
tissues, while exhibiting at the same time excellent mechanical and
functional properties in terms of facilitation of host cell
infiltration upon grafting, reduced inflammatory
reactions/rejection and improved tissue healing, with associated
faster recovery of a subject undergone to a surgical operation. The
scaffold of the invention can be stored off-the-shelf, thus
reducing the costs associated to the surgery, and providing
successful grafting results even in a cell-free scenario. In a
particular aspect, the present invention proposes an easy and
controllable process to provide non-uniformly stiff polymeric
scaffolds that consists of denser areas ensuring suturability and
easy handling for the surgeons, as well as softer zones to induce
the appropriate cell behaviour and fate in term of cell
proliferation and differentiation.
[0009] Accordingly, it is an object of the present invention to
provide a method for producing a non-uniformly stiff polymeric
scaffold for use in tissue engineering, diagnostics or surgical
procedures, characterized in that it comprises the following steps:
[0010] a) providing at least one layer of a polymeric material gel;
[0011] b) performing a first fast drying step by applying a
mechanical compression on said polymeric material gel layer; and
[0012] c) performing a second slow drying step of the gel up to
reach a polymeric material mass fraction of at least 60% w/w.
[0013] In one embodiment, the method is followed by a hydration
step where the scaffold re-swells with more than 4% w/w.
[0014] In one embodiment, the method is characterized in that step
c) is performed by water filtration means, air-drying or any
combination thereof.
[0015] In a particular embodiment, the method is characterized in
that water filtration is performed by dialysis.
[0016] In a particular embodiment, the method is characterized in
that air-drying is performed under gravity for at least 1 hour.
[0017] In one embodiment, the method further comprises the step of
circumferentially wrapping the polymeric scaffold around a support
in order to have a tubular polymeric scaffold comprising a
lumen.
[0018] In one embodiment, the method comprises step a') instead of
step a) of pouring a polymeric gel precursor solution into a
substantially cylindrical mould comprising therein a coaxial
elongated support for creating a tubular single layer of polymeric
material gel.
[0019] In one embodiment, the method is characterized in that the
polymeric scaffold is cell-free.
[0020] In one embodiment, the method is characterized in that the
polymeric material is natural polymeric material or an
extra-cellular matrix (ECM)-derived polymeric material.
[0021] In one embodiment, the method is characterized in that the
natural polymeric material or ECM-derived polymeric material is
gelatin, elastin, collagen, agar/agarose, chitosan, fibrin,
proteoglycans, a polyamino-acid or its derivatives, preferably
polylysin or gelatin methyl cellulose, carbomethyl cellulose,
polysaccharides and their derivatives, preferably
glycosaminoglycanes such as hyaluronic acid, chondroitinsulfate,
dermatansulfate, heparansulfate, heparine, keratansulfate or
alginate, or any combination of the foregoing.
[0022] In one embodiment, the method is characterized in that no
crosslinkers are added to the polymeric material gel.
[0023] In a particular embodiment, the method is characterized in
that the polymeric gel of step a) is obtained from a collagen gel
precursor solution having a collagen concentration of at least 2.1
mg/mL, preferably between 2.5 and 40 mg/mL.
[0024] A further aspect of the invention relates to a method for
producing a non-uniformly stiff polymeric scaffold for use in
tissue engineering, diagnostics or surgical procedures,
characterized in that it comprises the following steps: [0025] a)
providing a single layer of a mechanically compressed polymeric
material gel wrapped around a support; [0026] b) dipping a portion
of the support-wrapped polymeric gel layer into an aqueous solution
while leaving another portion air-drying under gravity for a
certain amount of time; [0027] c) flipping the polymeric gel layer
portion previously dipped into the aqueous solution onto the
air-dried portion; and [0028] d) optionally air-drying under
gravity both layers for a certain amount of time wherein at least
one layer of the so obtained polymeric scaffold has a polymer mass
fraction of at least 60% w/w.
[0029] In one embodiment, said method comprises step a') instead of
step a) of pouring a polymeric gel precursor solution into a
substantially cylindrical mould comprising therein a coaxial
elongated support for creating a tubular single layer of polymeric
gel, gelling it and mechanically compress it.
[0030] In one embodiment, the method is characterized in that the
polymeric scaffold is cell-free.
[0031] In one embodiment, the method is characterized in that the
polymeric material is natural polymeric material or an
extra-cellular matrix (ECM)-derived polymeric material.
[0032] In one embodiment, the method is characterized in that the
natural polymeric material or ECM-derived polymeric material is
gelatin, elastin, collagen, agar/agarose, chitosan, fibrin,
proteoglycans, a polyamino-acid or its derivatives, preferably
polylysin or gelatin methyl cellulose, carbomethyl cellulose,
polysaccharides and their derivatives, preferably
glycosaminoglycanes such as hyaluronic acid, chondroitinsulfate,
dermatansulfate, heparansulfate, heparine, keratansulfate or
alginate, or any combination of the foregoing.
[0033] In one embodiment, the method is characterized in that no
crosslinkers are added to the polymeric material gel.
[0034] In a particular embodiment, the method is characterized in
that the polymeric gel of step a) is obtained from a collagen gel
precursor solution having a collagen concentration of at least 2.1
mg/mL, preferably between 2.5 and 40 mg/mL.
[0035] A further aspect of the present invention relates to a
non-uniformly stiff polymeric scaffold for use in tissue
engineering, diagnostics or surgical procedures produced according
to the previously-described methods, said scaffold being
characterized in that it has at least a polymeric material layer
having a polymer mass fraction of at least 52% w/w at steady state
upon re-swelling, with a denser polymeric structure at the surface
and less dense in the core of the scaffold.
[0036] In one embodiment, the non-uniformly stiff polymeric
scaffold is characterized in that it is suturable.
[0037] In one embodiment, the non-uniformly stiff polymeric
scaffold is characterized in that it is at least 1 cm long.
[0038] In one embodiment, the non-uniformly stiff polymeric
scaffold is characterized in that it comprises a bioactive
agent.
[0039] Another object of the present invention relates to a kit
comprising: [0040] a) a container; [0041] b) an aqueous solution;
and [0042] c) the non-uniformly stiff polymeric scaffold as
previously described.
[0043] In one embodiment, the kit is characterized in that is
sterilized through electrode beam, gamma-rays radiation or X-rays
radiation.
BRIEF DESCRIPTION OF DRAWINGS
[0044] FIG. 1 depicts a sketch of the method of the invention for
the production of sheets (1A) or tubes (1B) of a non-uniformly
stiff polymeric scaffold;
[0045] FIG. 2 shows one embodiment of a mold used to produce a
polymeric tubular scaffold with an inner lumen;
[0046] FIG. 3 shows a schematic representation of the "sock
technique" according to the invention for producing non-uniformly
stiff polymeric scaffolds;
[0047] FIG. 4 shows images of the "sock technique" and the final
structure of a double-layered and a single-layered collagen tube,
as well as a collagen sheet obtained by opening of a double-layered
tube. (A) The upper half of a collagen tube is hydrated. (B-E) The
hydrated collagen half is pulled over the dry half. (F) A single
layered suture compatible collagen tube on a forceps. (G) A
double-layered suture compatible tube on a forceps. (H) The opened
up double-layered tube forming a double layered sheet;
[0048] FIG. 5 shows the influence of the drying method and the
final drying state on the re-swelling behaviour of collagen
scaffolds. (A) Mechanical compressed followed by air-drying is
compared to only air-drying of collagen, with monitoring of the
water content and time of air-drying. There is a clear trend in the
time of air-drying and water content (B) No drying of produced
collagen scaffolds results in a 100% water content in the collagen
scaffold (white bars after drying), and the water content does not
change upon incubation in a hydrating environment (black bars after
re-swelling). Utilizing the method of freeze-drying to reduce water
content in the collagen scaffold results in a dried scaffold with
an average of 1.3% water content, and reaches an average of around
20% water content after re-swelling. Utilizing only air-drying
allows obtaining an average dry level of 1.5% water content that
increases to 9.5% after re-swelling. Utilizing mechanical
compression with air-drying allows obtaining an average of 0.8%
water content that increases to 5.2% after re-swelling. Utilizing
only mechanical compression allows obtaining an average dry level
of 47.6% water content that increases to 48.3% after re-swelling.
(C) The air-drying method can be used to manipulate the re-swelling
behaviour of collagen scaffolds by changing the final dry-state (%
water); A linear trend between the final dry-state of the collagen
scaffold and the degree of re-swelling is observed; At least 1 h of
air drying after mechanical compression yields a different material
with different water content compared to only a mechanical
compressed scaffold;
[0049] FIG. 6 shows the re-swelling steady state of collagen
scaffolds. Freeze dried, air-dried, and mechanically compressed
with air-dried collagen scaffolds rapidly re-swell (within 30 min)
in a hydrating environment. A steady-state re-swelling is occurring
within 72 h in hydrating conditions;
[0050] FIG. 7 shows the relation between the mechanical properties
and percentage of water present within surgery-compatible collagen
scaffolds. (A and B) Air-dried collagen scaffolds show that water
percentage is a crucial parameter for the mechanical properties
(Young modulus and Ultimate tensile strength) of the scaffold. (C
and D) An air-dried (1.5% water content) and re-swelled collagen
scaffold (9.5% water content) is anastomosed to a goat urethra
using Vicryl 6.0 sutures. (E) Even a mechanical compressed plus
air-dried for 1 h (38% water content) and re-swelled collagen
scaffold (45% water content) is anastomosed to a dog urethra using
Vicryl 6.0 suture. (F and G) Hematoxylin & eosin staining of
tissue samples harvested 1 month and 3 months after surgery in
rabbits. White arrows show the sutures and black arrows point to
the remaining collagen of the implanted scaffold that is slowly
being replaced by host tissue and finally completely disappeared 3
months post-surgery;
[0051] FIG. 8 shows the microscopic collagen structure at the
micro-nano scale after different drying methods. A and D show a
collagen tube produced by mechanical compression (plastic
compression) followed by air drying (semi-elastic compression)
yielding a heterogeneous or "non-uniform" collagen structure. The
non-uniform structure is the result of the application of two
different drying methods during the fabrication procedure. B and E
shows a collagen tube that has only been air-dried yielding a
homogenous or "uniform" collagen structure. C and F show a collagen
tube that has been freeze-dried yielding a random, non-uniform
collagen structure, with pronounced holes distributed randomly in
the collagen network;
[0052] FIG. 9 visualizes the different collagen areas present
within a tubular scaffold that was produced by applying both
compressions, the plastic (mechanical load) and the semi-elastic
(air-drying) one. (A) Schematic description of the manipulations to
obtain a non-uniformly stiff tubular collagen scaffold. (B)
Schematic representation of the collagen structure of the final
tubular scaffold. (C and D) Immunohistochemistry using
anti-collagen type 1 antibodies on samples of a tubular bovine
collagen scaffold to visualize the collagen type 1 fiber density.
Images taken under a LSM700 confocal microscope (Zeiss) visualize
the surface of the inner lumen (D) and the outer side (C) of a
cross-section of a collagen tube;
[0053] FIG. 10 shows the morphology of urothelial cells cultured on
collagen scaffolds with different densities. GFP-lentivirial
transduced urothelial cells were used to easily visualize them
under a fluorescence stereomicroscope (Leica). 24 h after cell
seeding, urothelial cells growing on (A) less dense collagen
scaffolds form more single units and less connected colonies as
compared to (B) urothelial cells growing on higher density collagen
gels.
DESCRIPTION OF EMBODIMENTS
[0054] The present disclosure may be more readily understood by
reference to the following detailed description presented in
connection with the accompanying drawing figures, which form a part
of this disclosure. It is to be understood that this disclosure is
not limited to the specific conditions or parameters described
and/or shown herein, and that the terminology used herein is for
the purpose of describing particular embodiments by way of example
only and is not intended to be limiting of the claimed
disclosure.
[0055] As used herein and in the appended claims, the singular
forms "a", "an" and "the" include plural referents unless the
context clearly dictates otherwise. Thus, for example, reference to
"a bioactive agent" includes a plurality of such agents and
reference to "a layer" includes reference to one or more layers,
and so forth.
[0056] Also, the use of "or" means "and/or" unless otherwise
stated. Similarly, "comprise", "comprises", "comprising",
"include", "includes" and "including" are interchangeable and not
intended to be limiting. It is to be further understood that where
descriptions of various embodiments use the term "comprising",
those skilled in the art would understand that in some specific
instances, an embodiment can be alternatively described using
language "consisting essentially of" or "consisting of".
[0057] The main object of the present invention relies at least in
part on a novel method for producing a polymeric material scaffold
for use in tissue engineering or surgical procedures, as well as
for diagnostics and basic research, having very peculiar mechanical
and functional characteristics. In fact, once produced with the
methods according to the invention, the scaffolds of the invention
surprisingly show a non-uniform architecture, wherein some zones
within the polymeric scaffold display differences in polymer
density and therefore also in stiffness. Since there is a close
relation between the polymer concentration and the mechanical
properties of the scaffold such as stiffness and elasticity, this
feature advantageously guide cell migration, proliferation and
differentiation in vivo thanks to the presence of one or more less
dense polymeric area that can be easily remodeled by invading host
cells, thus enhancing the tissue regeneration capacity of the
scaffold, while keeping at the same time a portion of the scaffold
with a sufficient mechanical strength for being sutured during
surgery.
[0058] The method of the invention covers two main approaches for
the manufacturing of the non-uniformly stiff polymeric scaffolds of
the invention. Contrary to some prior art approaches involving a
process of plastic compression for creating uniform polymer (e.g.
collagen) matrices, the present method exploits a new method
renamed by the inventors "semi-elastic compression" of polymeric
gels. In contrast to an elastic compression, where the polymeric
gel swells back to their initial state after removal of the
compressive force applied therein, the gels processed according to
the present method swell back only to an intermediate state, which
is dependent on the initial polymer concentration and the final dry
state. Surprisingly, these semi-elastic compressed and rehydrated
polymeric structures are stable enough to be easily handled and to
be sutured in an in vivo surgical approach. This is realized by
applying a first fast mechanical compression step using mechanical
load followed by a slow drying phase, preferably followed by a
final rehydration.
[0059] The invention will be better understood with the help of the
following definitions.
[0060] As used herein, a "polymeric material" is any material
comprising polymers, large molecules (also known as macromolecules)
composed of many repeated smaller units, or subunits, called
monomers, tightly bonded together preferably by covalent bonds.
Polymer architecture at the molecular scale can be rather diverse.
A linear polymer consists of a long linear chain of monomers. A
branched polymer comprises a long backbone chain with several short
side-chain branches covalently attached. Cross-linked polymers have
monomers of one long or short chain covalently bonded with monomers
of another short or long chain. Cross-linking results in a
three-dimensional molecular network; the whole polymer is a giant
macromolecule. Another useful classification of polymers is based
on the chemical type of the monomers: homopolymers consist of
monomers of the same type, copolymers have different repeating
units. Furthermore, depending on the arrangement of the types of
monomers in the polymer chain, there are the following
classification: the different repeating units are distributed
randomly (random copolymer) or there are alternating sequences of
the different monomers (alternating copolymers) in block copolymers
long sequences of one monomer type are followed by long sequences
of another type; and graft copolymers consist of a chain made from
one type of monomer with branches of another type. A sufficiently
dense polymer solution can be crosslinked to form a polymer gel,
including a hydrogel or a cryogel, which is a soft solid.
[0061] Polymer materials may also be formed by blending two or more
polymers into physical mixtures. For example, the poor impact
strength of a polymer can be greatly improved by incorporating
small particles of an elastomer. Many properties of polymeric
materials depend on the microscopic arrangement of their molecules.
Polymers can have an amorphous (disordered) or semicrystalline
(partially crystalline, partially ordered) structure. Polymers can
be mixed with inorganic particles (usually in the form of
continuous fibres, such as glass or particulates such as mica, talc
and clay) in order to modify and improve (mainly but not
exclusively) their mechanical properties.
[0062] Suitable polymeric materials include, but are not limited
to, synthetic polymers such as polyurethanes, poly-olefins,
polyethylene glycol (PEG), poly(glycolide) (PGA), poly(L-lactide)
(PLA) or poly(lactide-co-glycolide) (PLGA); natural polymeric
materials (i.e., non-synthetic polymers, polymers that can be found
in nature) and/or polymers derived from the extracellular matrix
(ECM) as gelatin, laminin, elastin, collagen, agar/agarose,
chitosan, fibrin, proteoglycans, a polyamino-acid or its
derivatives, preferably polylysin or gelatin methyl cellulose,
carboxymethylcellulose (CMC), polysaccharides and their
derivatives, preferably glycosaminoglycanes such as hyaluronic
acid, chondroitinsulfate, dermatansulfate, heparansulfate,
heparine, keratansulfate or alginate, nucleotides, polylipides,
fatty acids, as well as any derivative thereof, fragment thereof
and any combination of the foregoing. Natural and ECM-derived
polymers are a first choice biomaterial for tissue engineering
applications envisaged by the present disclosure, due to their
biological and chemical similarities to natural tissues and the
presence of biologically active sites in their structures. For
example, collagen and hyaluronic acid can be considered as the most
preferred embodiments for constituting the polymeric scaffolds of
the invention.
[0063] As used herein, the term "gel" refers to a non-fluid
colloidal network or polymer network that is expanded throughout
its whole volume by a fluid. A gel is a solid three-dimensional
network that spans the volume of a liquid medium and ensnares it
through surface tension effects. The internal network structure may
result from physical bonds (physical gels) or chemical bonds
(chemical gels).
[0064] As used herein, the term "hydrogel" refers to a gel in which
the swelling agent is water. A hydrogel is a macromolecular polymer
gel constructed of a network of crosslinked polymer chains. It is
synthesized from hydrophilic monomers, sometimes found as a
colloidal gel in which water is the dispersion medium. Hydrogels
are highly absorbent (they can contain over 90% water) natural or
synthetic polymeric networks. As a result of their characteristics,
hydrogels develop typical firm yet elastic mechanical
properties.
[0065] Several physical properties of the (hydro)gels are dependent
upon concentration. Increase in (hydro)gel concentration may change
its pore radius, morphology, or its permeability to different
molecular weight proteins. One skilled in the art will appreciate
that the volume or dimensions (length, width, and thickness) of a
(hydro)gel can be selected based on instant needs, such as for
instance the region or environment into which a scaffold
substantially made of an (hydro)gel is to be implanted.
[0066] For "collagen" is herein meant any kind of natural or
recombinant collagen of animal origin, preferably mammals, such as
bovine, porcine, rat, human and the like. Accordingly, collagen
type 1 to 28 can be used in the frame of the present invention, as
well as fragments thereof or any combinations of the foregoing.
Collagen can also possibly derive from any body location of the
said animals such as skin, intestine, cartilage and so forth, as
long it fits with biocompatibility requirements or with translation
into an in vivo approach. In a preferred embodiment, type 1
collagen of bovine origin is used.
[0067] The polymeric gel of the invention can be
manufactured/processed through any suitable manufacturing method
known in the art allowing to create a highly interconnected pore
network in the material, such as e.g. (photo)lithography including
two-photon lithography, casting and molding, 3D printing, inkjet
printing, porogen leaching, emulsion freezing/freeze drying
technique, inverse opal hydrogelation, cryogelation,
electrospinning or fiber extrusion and bonding, gas foaming,
thermally induced phase separation and so forth.
[0068] In one embodiment, the polymeric gel can be fabricated by
top-down processes such as 3D printing. This technique allows full
control over the distribution and size of pores and wall material
in nearly arbitrary geometries, and therefore holds great
possibilities in terms of organ-specific 3D design and
fabrication.
[0069] In the frame of the present disclosure, a "scaffold" is any
three dimensional material having a framework architecture, i.e. a
support structure usually comprising hollow spaces within it.
Generally speaking, a scaffold material is an artificial, usually
temporary template structure capable of supporting
three-dimensional body tissue/organ formation in vivo, ex vivo or
in vitro. In this context, a scaffold material is also referred
herewith as a "biomaterial" or "bioscaffold". A bioscaffold, inter
alia, allows cell attachment and migration, delivers and retains
cells and biochemical factors, enables diffusion of vital cell
nutrients and expressed products, exerts certain mechanical and
biological influences to modify the behaviour of the cells and so
forth.
[0070] The scaffold material of the invention has been conceived
and manufactured in order to act as a biocompatible, non-migratory
and non-carcinogenic scaffolding agent. The general purpose was the
development of a material having improved long-term efficacy, which
can stimulate host cell infiltration and integrates with the
surrounding tissue once implanted in a host, thus triggering
neo-tissue formation via the promotion of a bioactive environment
within the application site and giving rise to long-term functional
repair. The scaffold material herein disclosed addresses and
solves, among others, this problem.
[0071] The scaffold material is moreover, at least on some portions
thereof, highly porous. In a preferred embodiment of the invention,
the pores are all interconnected in order to create a continuous
net of material that can act as a plausible physical support for
other elements such as cells or bioactive agents, while providing
at the same time additional key features to the scaffold such as
its softness/stiffness, easiness of cell/tissue invasion and so
forth. Accordingly, where present, pore diameters preferably
comprised between about 0.2 and 200 .mu.m in the less dense areas,
with more preferred pore diameters comprised between about 0.5 and
10 .mu.m. As will be evident to a person skilled in the art, these
less dense areas within the scaffold are present, due to the
non-uniform stiffness (and therefore density) of the scaffold
itself, in the softer (and therefore, less dense) portion, while
the stiffer (and therefore, denser) portion thereof is
characterized by a negligible porosity, with pores, where present,
with an average diameter of less than 0.2 .mu.m. Moreover, it is
preferred that the stiffer portion of the polymeric scaffold has a
Young's modulus comprised between about 5 and 1500 kPa, whereas
preferred Young's modulus values for the softer portion span
between about 0.1 and about 20 kPa. In any embodiment, however, at
least a 2.5 kPa difference in terms of Young's modulus between the
softest and stiffest part of the scaffold is present.
[0072] In order to optimize the mechanical properties of the
scaffold material of the invention and, in some aspects, its
resorption/biodegradation rate, in preferred embodiments it is
contemplated that the scaffold comprises a final polymer weight in
a dried form of at least the 60% of the total scaffold weight (i.e.
60% w/w, also called mass fraction, wherein the remaining 40% or
less is substantially composed of water or of an aqueous solution).
A suitable polymer mass fraction depends on e.g. the molecular
weight of the monomer, the nature of the monomer, the crosslinking
strategy and the ratio of polymers. However, this value is tightly
associated to the manufacturing steps of the method for producing
said polymeric scaffold; in fact, as it will be more deeply
detailed later on, the final polymer mass fraction in the scaffold
can even be of at least 52% w/w upon re-swelling (re-hydration) of
the scaffold in an aqueous solution, at the end of the production
procedure.
[0073] Concerning the possible degradation/resorption rate of the
scaffold upon in vivo application/implantation in a subject, this
is mainly dependent on physico-chemical properties of the polymeric
material of which it is composed of, as well as further factors
such as crosslinking of the polymers, the polymer concentration,
the site of implantation into a host and the like. The
degradation/resorption rate can be calibrated by adjusting said
physico-chemical parameters, such as for instance by polymer
crosslinking, the use of inhibitor molecules, by changing the
polymer density, crystallinity and/or its molecular weight
distribution, changing the materials' porosity and so forth.
Generally speaking, the scaffold material may be, at least in part
and at least in some portion thereof, intrinsically biodegradable
in vivo.
[0074] The scaffolds and the gels of the invention are preferably
cell-free, but they can be seeded in vitro or in vivo with cells
before, during and/or after their production. The scaffold, if to
be used to transplant cells in a tissue engineering approach,
comprises pores to permit the structure to be seeded with cells and
e.g. to allow the cells to proliferate. For example, scaffolds are
seeded by incubating the structure in a solution containing the
cells for a suitable period of time. Alternatively, a
multi-component scaffold is built in stages with each layer being
seeded prior to laying down of another layer or before adherences
of another pre-formed component. Alternatively, each compartment of
the multi-compartment scaffold can be seeded separately. Different
cell types, e.g., stem vs. differentiated, and/or with various
phenotypes in terms of differentiation, activation, metabolic or
functional state, can be seeded in the scaffold material. The
scaffolds are suitable for use with any cell type, preferably
forming soft tissues, which one may want to transplant. Such cells
include but are not limited to, various stem cell populations
(embryonic stem cells differentiated into various cell types), bone
marrow or adipose tissue derived adult stem cells, mesenchymal stem
cells, cardiac stem cells, pancreatic stem cells, neuronal cells,
glial cells, spermatozoids and oocytes, endothelial progenitor
cells, outgrowth endothelial cells, dendritic cells, hematopoietic
stem cells, neural stem cells, satellite cells, side population
cells. Such cells may further include but are not limited to,
differentiated cell populations including osteoprogenitors and
osteoblasts, chondrocytes, keratinocytes for skin, intestinal
epithelial cells, smooth muscle cells, cardiac muscle cells,
epithelial cells, endothelial cells, urothelial cells, fibroblasts,
myoblasts, chondroblasts, osteoclasts, hepatocytes, bile duct
cells, pancreatic islet cells, thyroid, parathyroid, adrenal,
hypothalamic, pituitary, ovarian, testicular, salivary gland cells,
adipocytes and combinations thereof. For example, smooth muscle
cells and endothelial cells may be employed for muscular or tubular
scaffolds, e.g., scaffolds intended as vascular, oesophageal,
intestinal, rectal, or ureteral scaffolds; chondrocytes may be
employed in cartilaginous scaffolds; cardiac muscle cells may be
employed in heart scaffolds; hepatocytes and bile duct cells may be
employed in liver scaffolds; myoblasts may be used in muscle
regeneration; epithelial, endothelial, fibroblast, and nerve cells
may be employed in scaffolds intended to function as replacements
or enhancements for any of the wide variety of tissue types that
contain these cells. In general, scaffolds of the invention may
comprise any cell population competent to participate in
regeneration, replacement or repair of a target tissue or organ,
particularly (semi)soft or barely-reachable ones such as kidneys,
brain, meninges, lungs, liver, stomach, sphincters, gut, bladder,
ear tissues, cartilages, skin, pharynx, trachea, blood vessels,
urethra, ureter, vagina, pelvic floor or pancreas.
[0075] Some mechanical/functional properties of the scaffold can be
tailored according to the needs by changing the physical or
chemical properties thereof (such as e.g. the collagen molecular
chain length). Concerning collagen, in order to optimize these
parameters and, in some aspects, the resorption/biodegradation
rate, in preferred embodiments it is contemplated an average
molecular weight for the collagen molecules substantially composing
the scaffold comprised between about 30 and about 250 kDa.
Concerning hyaluronic acid, an average molecular weight comprised
between about 50 kDa and about 2 MDa is particularly preferred.
[0076] In most preferred embodiments, the final scaffold is not
crosslinked or minimally crosslinked, in order to eliminating or
reducing as much as possible e.g. the potential in vivo toxicity
associated with some crosslinking agents such as glutaraldehyde.
When a crosslinking process is used, crosslinking agents and their
amount can be chosen at the operator's discretion, and a skilled in
the art would easily envisage such parameters based on common
practice. By practicing the production method of the invention that
will be detailed later on, it is possible to produce scaffolds with
such mechanical and functional properties to allow at the same time
successful grafts in a host with ease suturability, without the
impellent need of using crosslinking agents. For being considered
successful, a graft should show no complications at the grafting
site, such as no presence of opened structures as fistulae,
strictures or stenoses, and/or the treated tissue/organ area should
look histologically normal.
[0077] In some embodiments, the scaffold can be differentially
permeable, allowing cell displacement only in certain physical
areas thereof. The permeability of the scaffold composition is
regulated, for example, by engineering the polymeric material for
greater or smaller pore size, porosity, density, polymer
cross-linking and/or viscoelasticity.
[0078] The scaffold material of the invention can be organized,
during or after the production method, in a variety of geometric
shapes such as beads, pellets, strip, block, toroid, capillary,
patches, tubes, planar layers (e.g., thin sheets) and so forth, the
final shape depending on the particular context it is to be used,
and can comprise in some embodiments at least two areas or
compartments having different structural/functional properties. For
example, multicomponent scaffolds are constructed in e.g.
concentric layers, each of which is characterized by different
physico-chemical properties (% polymer, % crosslinking of polymer,
chemical composition of scaffold, pore size and/or architecture,
stiffness, porosity, presence or different concentration of
bioactive agents and so on). Each niche can host one or more cell
populations and/or have a specific effect on them, e.g., promoting
or inhibiting a specific cellular function, proliferation,
differentiation, migration and so forth. Such a configuration is
particularly useful in maintaining for long time periods the
"sternness" of a population of cells, while simultaneously pushing
daughter cells to multiply rapidly and differentiate appropriately
for participation in tissue regeneration.
[0079] A compartmentalization of the device can be obtained by
rational design and fabrication processes using different
compositions or concentrations of compositions for each
compartment. For example, a stem cell population can be
encapsulated within hydrogels, using standard encapsulation
techniques. Alternatively or additionally, two different areas or
compartments can be formed within the bioscaffold that contains
distinct factors (e.g., morphogens, growth factors, adhesion
ligands), or the material in a distinct form (e.g., varying
mechanical properties or porosity). In some embodiments, the
compartments can be fabricated individually, and then operably
coupled to each other with any suitable method known in the art
such as use of an adhesive or exploitation of the intrinsic
adhesiveness of each compartment, polymerization or cross-linking
of one compartment to another and so forth. The scaffold can be
designed to have a number of compartments or areas in which cells
enter in parallel and distribute according to their characteristics
(e.g., size or mobility), serially pass through all or some of the
compartments, or a combination of both. The different compartments
can even be construed to induce distinct fates for the contained
cells during the passage therein.
[0080] During or after the manufacturing process, the scaffold
material can be functionalized with additional elements such as for
instance bioactive molecules. Said elements can be coated on or
embedded within the bioscaffold with any suitable means known in
the art, either in a homogeneous or non homogenous manner and can
provide additional functional properties to the material such as
enhanced/reduced biodegradation, physical stabilization, biological
activity and the like. As used herein, a "bioactive molecule", as
well as "(bio)active compound" or "therapeutic agent", is any
active agent having an effect upon a living organism, tissue, or
cell. The expression is used herein to refer to any compound or
entity that alters, inhibits, activates, or otherwise affects
biological or chemical events.
[0081] One skilled in the art will appreciate that a variety of
bioactive compounds can be used depending upon the needs. Exemplary
therapeutic agents include, but are not limited to, a small
molecule, a growth factor, a protein, a peptide, an enzyme, an
antibody or any derivative thereof (such as e.g. multivalent
antibodies, multispecific antibodies, scFvs, bivalent or trivalent
scFvs, triabodies, minibodies, nanobodies, diabodies etc.), an
antigen, a nucleic acid sequence (e.g., DNA or RNA), a hormone, an
anti-inflammatory agent, an anti-viral agent, an anti-bacterial
agent, a cytokine, an oncogene, a tumor suppressor, a transmembrane
receptor, a protein receptor, a serum protein, an adhesion
molecule, a lypidic molecule, a neurotransmitter, a morphogenetic
protein, a differentiation factor, an analgesic, organic molecules,
metal particles, radioactive agents, polysaccharides, a matrix
protein, and any functional fragment or (chimeric) derivative of
the above, as well as any combinations thereof. For "functional
fragment" is herein meant any portion of an active agent able to
exert its physiological/pharmacological activity. For example, a
functional fragment of an antibody could be an Fc region, an Fv
region, a Fab/F(ab')/F(ab').sub.2 region and so forth.
[0082] The bioactive compounds can be added to the scaffold
material or to the polymeric gels by using any suitable method
known in the art, such as surface absorption, physical
immobilization, e.g., using a phase change to entrap the substance
in the scaffold material, and the like. For example, a growth
factor can be mixed with a polymeric (e.g. collagen) composition
while it is in an aqueous or liquid phase, and after a change in
environmental conditions (e.g., pH, temperature, ion
concentration), the liquid gels or solidifies, thereby entrapping
the bioactive substance. Alternatively, covalent coupling, e.g.
using alkylating or acylating agents, is used to provide a stable,
long-term presentation of a bioactive substance on the scaffold in
a defined conformation. Alternatively, non-covalent adsorption can
be used, for example electrostatic, hydrophobic, dipole-dipole,
hydrogen bonding, physisorption and the like. In an additional or
alternative embodiment, bioactive molecules can be encapsulated
within nano/micro spheres or beads included within the scaffold
and/or a polymeric gel during or after a manufacturing step.
Particularly suitable nano//micro beads are those described in
European Patent Application EP15193284.5, incorporated herein in
its entirety by reference.
[0083] The scaffold material described herein is useful in the
treatment of a variety of diseases, disorders, and defects where
surgery or a tissue engineering approach can be a suitable
therapeutic solution. This is particularly true for soft tissues,
and the scaffold material may therefore be utilized in a variety of
surgical procedures as well as for cosmetic purposes, and for the
treatment or prevention of a plethora of pathological conditions.
The scaffold material of the invention results particularly
convenient for treating tissues or organs like a genitourinary
tract components, including kidney, bladder, urethra, ureter,
vaginal and pelvic floor tissue, a blood vessel (including big
veins and arteries), a muscle, a cartilage, skin, liver, an eye
portion such as cornea, conjunctiva or sclera, a Central Nervous
System (CNS) or Peripheral Nervous System (PNS) component including
cerebral or meningeal tissues, trachea, oesophagus, lungs, stomach,
heart, sphincters, pancreas, gut, pharynx or inner ear tissue.
[0084] The scaffold is usually transplanted on or close to a target
tissue, introduced into or onto a bodily tissue/organ of a subject
or a host using a variety of methods and tools known in the art,
preferably via minimally invasive surgical devices and
procedures.
[0085] The term "subject" or "host" as used herein refers to
animals, particularly to mammals. For example, mammals contemplated
by the present invention include humans, primates, pets,
domesticated animals such as cattle, sheep, pigs, horses, rabbits,
rodents and the like.
[0086] As used herein, "treatment", "treating" and the like
generally mean obtaining a desired pharmacological and/or
physiological effect. The effect may be prophylactic in terms of
preventing or partially preventing a disease, symptom or condition
thereof and/or may be therapeutic in terms of a partial or complete
cure of a disease, condition, symptom or adverse effect attributed
to the disease. The term "treatment" or "treating" as used herein
covers any treatment of a disease in an animal, particularly in a
mammal, more particularly a human, and includes: (a) inhibiting the
disease, i.e., arresting its development; or (b) relieving the
disease, i.e., causing regression of the disease and/or its
symptoms or conditions such as improvement or remediation of
damage. The term "prevention" or "preventing" relates to hampering,
blocking or avoid a disease from occurring in a subject which may
be, for any reason, predisposed to the disease but has not yet been
diagnosed as having it for example based on familial history,
health status or age.
[0087] Initially, the method for producing the scaffold of the
invention foresees a first fast drying step via a mechanical
compression of a polymeric gel, preferably hydrogel, layer (FIG.
1A). It has been assessed that this first step of mechanical
compression allows to achieve a non-uniformly stiff micro/nano
structure of the gel and, consequently, of the final scaffold. When
working with polymeric gel layers in form of sheets, the mechanical
compression is generally done by applying a rolling or linear
compressive force, which must be orthogonal to the planar surface
of the sheet. This latter is e.g. placed on a nylon mesh positioned
on tissue paper to augment the water absorption upon leakage
following the compression. When working with polymeric gel layers
already in form of tubes (FIG. 1B), these latter can be rolled
against an adsorbent tissue, with or without e.g. a nylon mesh
positioned in between, via e.g. a glass rod or otherwise a
cylindrical mandrel placed within the lumen of the tube, until no
more water is seen on the tissue paper, for instance by performing
10 revolutions of about 30 seconds each on the tissue paper. In
some embodiments, an additional filtration step to mechanically
suck water from a filter in contact with the lumen can be
performed. Tubular polymeric single layers can be obtained by
pouring a polymeric gel precursor solution into a substantially
cylindrical mould comprising therein a removable, coaxial elongated
support (e.g. a glass rod) of any suitable diameter so to create a
void space, the tube's lumen, upon formation of the polymeric gel
(through e.g. gelation of the precursor solution; FIG. 2).
[0088] A key feature of the method according to the invention
relies in a second, slow drying step of the polymeric gel in order
to obtain the final single layer polymeric scaffold. This step
envisages the use of techniques such as water filtration or
preferably air drying, as well as combinations thereof, for
gradually and gently reducing the water content of a polymeric gel
up to a polymer mass fraction in the final scaffold of at least 60%
w/w. Preferred water filtration means can be for instance dialysis
means. In the most preferred embodiment, the method is
characterized in that the second slow drying step is performed by
air-drying under gravity for at least 1 hour, preferably for at
least 4 hours. The air drying process is preferably carried out at
room temperature (i.e., between about 20.degree. and about
30.degree. C.), and in some embodiments in a controlled atmosphere
to keep sterility of the final scaffold. Once obtained a single
layer polymeric scaffold, this can be possibly circumferentially
wrapped around a suitable support of any shape, such as for
instance a removable glass rod, in order to realize a tubular
Collagen scaffold comprising a lumen therein.
[0089] The so-obtained slowly-dried scaffolds feature peculiar
characteristics both in terms of mechanical values when they are
sutured in a surgical approach, as well as for what concern tissue
regeneration and healing in a host, compared to solutions known in
the art. In fact, the scaffolds of the invention offer an excellent
compromise between these two issues, that must contemporary be
considered when it comes to the clinics. Current clinically used
natural materials sold as off-the shelf solutions include for
instance freeze-dried collagen scaffolds, cross-linked collagen
scaffolds and decellularized human and animal tissue-derived
scaffolds. These products are designed with a surgical mind set of
being easy to surgically work with. The scaffolds of the invention,
on the other hand, are strong enough to be sutured while providing
in parallel a more "cellular friendly" matrix, compared to the
commercially available solutions, i.e. surgeon-friendly and stiff
scaffolds with poor cellular responses. As a way of example,
commercially available tubular acellular materials do not work in
urethral replacement of length defects longer than 1 cm; if cells
are placed within these commercial available materials they can
achieve successful grafts for patients. On the other hand, the
present collagen-made scaffolds produced according to the methods
of the invention show successful results with 1 cm long or even
bigger (e.g., 2 cm long) tubular grafts without the addition of any
cells. In the same way, planar scaffolds such as sheets or patches
sized at least 1 cm.sup.2 can be successfully used in vivo.
[0090] In this context, without being bound to this theory, the
present inventors observed that very mechanically stable donor
tissues, which has been decellularized or devitalized to not elicit
immune response in the graft host patients, are too stiff, or
otherwise not enough soft, to give a functional repopulation of the
grafted scaffold by host cells. The extreme internal complexity
and/or density creates a sort of "extracellular matrix
(ECM)-mismatch" between the grafted matrix and the injured/tissue
regeneration site. This mismatch is plausibly the cause for an
ill-favoured cell ingrowth, leading to failure of the therapy for
the patients. The scaffold of the invention targets instead a novel
"Mechanical and ECM-complexity niche": it is structurally less
complex, less stiff, less dense and therefore easier to be invaded
by endogenous host cells without giving strong inflammatory
response, thus favouring a better and accelerated regeneration
profile than other materials, but still easily suturable. The
method of the invention allows to produce scaffolds targeting this
specific "Mechanical and ECM-complexity niche".
[0091] For instance, in the case of the urinary tract, the
scaffold's surface stiffness favours a better and faster urothelial
coverage, which is probably linked to a better protective
environment for sub-urothelial tissue regeneration (like muscle and
vascular regeneration). Without being necessarily bound to this
theory, at least one layer in the scaffold shall lay close to
native adult tissue in terms of mechanical properties for having a
better surgical experience, but a major component should be
protected to be soft. This could explain why decellularized native
tissues, having mechanic properties very close to those of native
adult tissues, fail or do not have perfect results; mimicking the
mechanics of early development tissues, at least on portions of a
scaffold material, could be physiologically more fitting and
"better sensed" by a receiving host. In this context, in preferred
embodiments of the invention, the stiffer portion of the polymeric
scaffold should have a retention strength of at least 0.2 N upon
suture, preferably of at least 0.4 N, even more preferably with
values comprised between about 1.5 and 2 N.
[0092] In terms of swelling behaviour, which then also reflects the
mechanical properties of the scaffold of the invention (water
content in polymeric scaffolds are known to be related to the
mechanical properties, i.e. the more water content the lower
mechanical strength), it has been surprisingly assessed by the
inventors that the scaffolds produced with the present methods,
when immersed in an aqueous solution such as Phosphate buffered
saline (PBS), re-swell and increase their water content only to up
of a maximum of 12% w/w at steady state. Compared to other
manufacturing techniques, this has important advantages in terms of
functional and mechanical properties of the final scaffold: by
applying a fast mechanical compression to a first polymeric gel
followed by a slow drying step, a scaffold ready to be sutured can
be easily achieved. In comparison, a freeze-dried collagen gel
swells and increase its water content to more than 20% w/w (FIG.
5B). Mechanically compressed collagen gel swells and increases its
water content only by 0.7%. Substantially, the developed method
permits to better control the re-swelling behaviour of the obtained
scaffold, and then ultimately the mechanical and functional
properties ideal for surgery.
[0093] This property represents moreover a huge advantage for what
concerns the stocking and the packaging of the scaffold of the
invention, since it could be provided to a surgeon as an
off-the-shelf, ready-to-use solution. In fact, it can be imagined a
polymeric scaffold of the invention prepared and packaged in a
liquid/aqueous solution already adapted for use in an in vivo
approach, without any change in its properties at steady state even
when hydrated. The "re-swelling steady state" is defined as the
state in terms of water content of the polymeric scaffold produced
with the methods of the invention in which the mechanical
properties and the water content of the said scaffold do not change
anymore.
[0094] With the aim of accentuating the mechanical differences
between two different areas of a polymeric scaffold of the
invention, the inventors developed a very ingenious and elegant
method for producing non-uniformly stiff scaffolds, said method
being characterized in that it comprises the following steps:
[0095] a) providing a single layer of a mechanically compressed
polymeric material gel wrapped around a support;
[0096] b) dipping a portion of the support-wrapped polymeric gel
layer into an aqueous solution while leaving another portion
air-drying under gravity for a certain amount of time;
[0097] c) flipping the polymeric gel layer portion previously
dipped into the aqueous solution onto the air-dried portion;
and
[0098] d) optionally air-drying under gravity both layers for a
certain amount of time
[0099] wherein at least one layer of the so obtained polymeric
scaffold has a polymer mass fraction of at least 60% w/w. This
method, renamed by the inventors "the sock-technique" and shown in
FIGS. 4 and 5, can be applied, mutatis mutandis, to polymeric gels
formed by e.g. pouring a polymeric gel precursor solution into a
substantially cylindrical mould comprising therein a coaxial
elongated support for creating a tubular single layer of polymeric
gel, gelling it and mechanically compress it.
[0100] An additional object of the invention relates to a kit
comprising a container, an aqueous solution and the polymeric
scaffold of the invention. The three components can be provided
separated within the kit, or the container can already include the
liquid solution and also possibly the polymeric scaffold in a
re-swelled, steady-state form. Preferably, the entire kit is
sterilized or sterilisable by electrode beam, gamma rays or X-rays
radiations.
EXAMPLES
[0101] The present, non-limiting examples present one embodiment of
the fabrication method of a non-uniformly stiff tubular collagen
scaffold according to the invention. The method can be easily
adapted to make sheet and dome-shaped structures, and dimensions
and shape of the collagen scaffold depend only of the used mold.
For the fabrication of a tubular non-uniformly stiff collagen
scaffold with an inner lumen of 3 mm, 8.5 mL of liquid
acido-soluble bovine collagen solution (5 mg/mL, Symatese, France)
are mixed with 0.8 mL 10.times. MEM (Invitrogen) and neutralized by
the addition of approximately 1.85 mL of 0.1M NaOH. The exact
volume of NaOH to be added is determined by observing a colour
change from yellow to pink. The neutralized collagen solution is
poured into a tubular shaped, stainless steel mould (length 70 mm,
inner diameter 12 mm). The lumen of the tubular collagen scaffold
is formed by a glass stick (diameter 3 mm) placed inside the
tubular mould. Collagen gelation takes place at room temperature
for 15 min. After gelation, the collagen tube is removed from the
mould but kept on the glass stick. This gelated collagen tube is
the starting matrix to produce non-uniform (denser external walls
and a less dense internal side) tubular collagen scaffolds.
[0102] The tubular collagen gel is mechanically compressed by
rolling on around its own axis, to ensure unequal compression and
drainage of water within the final scaffold. The collagen gel is
rolled over a nylon mesh placed on five layers of double-layer
tissue paper (Weita, Arlesheim, Switzerland) until no more water is
coming out. Subsequently, a further drying step is performed by
exposing the collagen tube to air at room temperature for 1 to 24
hours. The desired dryness corresponds to 0.5-40% of the initial
water content of the gel, preferably 18-22% of water. The weight of
the collagen gel after gelation represents 100% of water. The
percentage of water in each dried collagen scaffold is calculated
by taking the difference of the collagen gel weight before and
after drying:
% of water=1-[(weight before drying-weight after drying)/weight
before drying]
[0103] Once the collagen tube has reached the desired dryness, the
tube is placed in PBS (pH 7.4) for 3 days in order to re-swell it
to a steady-state water content. The water content of the final
collagen scaffold is in the range of 5 to 48%, preferably 24-30% of
water content.
[0104] For the fabrication of a multi-layered collagen tube through
the "sock technique", half of a tube dried with a first rolling
step as previously described is placed into a rehydration bath
(PBS, pH 7.4) for 1 to 60 minutes, preferably for 30 minutes. The
rehydrated half tube is then pulled over the dried tube portion or
the dry portion is pushed under the rehydrated part for creating a
multi-layered tube. Multi-layered tubes (with more than two layers)
can be obtained by iterating the above described steps.
[0105] Ultimate tensile strength (UTS) and Young's modulus were
determined under stress strain conditions using an Instron tensile
machine (Norwood, Mass., USA), applying a 250 N load with a strain
rate of 1 mm/min until the break point. Tubular samples of 5 mm
length was pulled from the lumen side in one direction using
L-shaped hookes placed in the lumen. Young's modulus values were
obtained from the slope of the initial linear region of the
stress-strain curves, while the UTS were obtained from the maximum
tensile stress recorded.
[0106] To determine the collagen density and porosity in the
scaffolds, scanning electron microscopy (SEM) was performed. The
samples were fixed with 1% tannic acid and 1.25% glutaraldehyde,
then washed with 0.1 M cacodylate, and dehydrated in increasing
ethanol concentrations prior to critical point drying (CPD). They
were coated with gold/palladium and imaged at a voltage of 10 kV
using a scanning electron microscope (SEM, XLF30, Philips). To
further visualize the collagen density and porosity within the
scaffolds, formalin fixed (4% PFA) and paraffin embedded cut
sections (8 .mu.m thick) were stained with anti-collagen typel
antibodies (1:250 dilution, rabbit anti-collagen type 1, Abcam,
Switzerland). Corresponding secondary AlexaFluor antibodies (Abcam,
Switzerland) were used to visualize the type 1 collagen
distribution under a LSM700 microscope (Zeiss, Germany).
[0107] To determine the in vitro suture capacity, anastomosis of
two produced collagen tubes was performed using Vicryl 6.0.RTM.
sutures. Furthermore, to determine the in vivo suture capacity of
the produced tubes, at least 9 rabbits had a urethral replacement
surgery, where a double end-to-end anastomoses to the native
urethra were performed with interrupted sutures (Vicryl 6.0.RTM.).
Rabbits were sacrificed after 1 and 3 months. Thereafter, rabbit
penises were harvested and fixed in 4% formalin, embedded in
paraffin, and sectioned in 8 .mu.m thick for Hematoxylin-eosin
staining. Images were taken with a Leica DM5500 microscope
(Germany).
[0108] To understand different cell type behaviour (e.g. urothelial
and smooth muscle cells) on various collagen-based scaffold with
different stiffness and collagen fiber densities, human urothelial
cells that were lentiviral-transduced with GFP were seeded onto
these collagen scaffolds and cell morphology was observed under a
fluorescence stereomicroscope (Leica). The human cells were placed
in a fibrin drop to control the spatial distribution of cells on
the collagen, and the cells at various time points were then imaged
to verify the cells fate choices.
[0109] Results
[0110] First of all, mechanical compression with air-drying and
only air-drying yields different materials, as seen in FIG. 5A it
also takes longer time to reach different dry states for only
air-dried material compared to mechanical compression with
air-drying. A fast compression yields less re-swelling compared to
the other method of slow drying. As shown in FIG. 5B, the collagen
scaffolds dried with either freeze-drying (FD) air-drying (AD) or
mechanical compression with air drying (M-AD) re-swell when they
are placed in hydrating conditions (FD dry: 1.3% to FD re-swell:
20%, AD dry: 1.5 to AD re-swell: 9.5%, M-AD dry: 0.8% to M-AD
re-swell 5.2% water content). However, mechanically compressed (M)
have a very low swelling capacity when placed in hydrating
conditions (M dry: 47.6% to M re-swell: 48.3% water content).
Furthermore, the mechanical compressed plus air-dried collagen
scaffold's re-swelling can be controlled by its dry state, as shown
in FIG. 5C. There is a linear relationship between the dry state
and the re-swelling state of the produced collagen scaffolds. The
re-swelling in hydration (in PBS-bath) is a rapid process where
most of the swelling has occurred within the first 30 minutes (FIG.
6). A more stable re-swelling steady-state is seen after 72 h in a
hydrating environment. The mechanical properties, Young's modulus
and UTS of the collagen scaffolds with different water contents
follow an exponential curve (FIGS. 7A and 7B). As shown in FIGS. 7C
and 7D, the collagen tube dried and re-swelled to a 9.5% water
content is sutured with ease into rabbits and goat urethra's (goat
urethra is shown in FIG. 7D). Also the re-swelled collagen scaffold
of 45% water content is sutured to rabbit and dog urethra (dog
urethra surgery is seen in FIG. 7E). Classifying what is suture
compatible is difficult since it is highly dependent on the
surgeon, and can only be done on an empirical basis. However,
starting from the mechanical properties recorded and surgeons
opinion when handling the collagen grafts of different water
content, a material that has a dry state of at least one collagen
layer of no more than 40% w/w water content (i.e., at least 60% w/w
collagen mass fraction) was classified as suture compatible
directly after it has been re-swelled to steady-state in hydrating
condition, when produced from a mammalian collagen precursor liquid
solution with range of collagen concentration comprised between 2.1
and 40 mg/mL.
[0111] The high functional regeneration potential of the scaffold
of the invention is highlighted in a surgical setting where
collagen tubes produced according to the methods of the invention
were implanted and sutured in a rabbit urethra. The collagen
scaffolds stay in the sutured area as shown in a histology section
of 1 month after surgery (FIG. 7E), black arrows show non-degraded
collagen from remaining parts of the collagen scaffolds). After 3
months, no more of the implanted collagen can be visualized by
histology, since the rabbits' own cells have remodelled the
collagen and populated the area (FIG. 7F).
[0112] A crucial observation done by the inventors is that a
slow-drying procedure creates an uniform and homogenous collagen
architecture at micro- and nanoscale (FIGS. 8B and 8D). While a
freeze-dried collagen structure yields an unorganized non-uniformed
architecture with random distribution of holes (diameter of 50-150
.mu.m) (FIGS. 8C and 8F) at micro and nanoscale. In comparison, if
one combine a first mechanical compression followed by a
slow-drying step, this yields to a non-uniform but organized
heterogeneous collagen architecture at micro- and nanoscale, where
the surface that is in close contact with the mechanical force gets
a more dense structure as compared to the internal side (FIGS. 8A
and 8C). In FIG. 9, the same feature is shown by utilizing
immunohistochemistry to visualize collagen type 1.
[0113] Many studies report that material stiffness, fiber density
and porosity influence cell fate choices, and that different cell
types behave differently by changing the material properties. FIG.
10 depicts the behaviour of urothelial cells in response to
collagen matrices that differ in stiffness. In a less dense
collagen matrix, the urothelial cells are less in contact with
neighbours and migrate away from their starting position. On the
other hand, urothelial cells seeded on a denser scaffold stay more
in contact with neighbour cells and migrate less.
[0114] To further increase mechanical stability and the potential
for creating non-uniformly stiff materials, with an option of
spatially distribute growth factors and reagents to modify the
collagen at various position in a scaffold, a further method, shown
in FIG. 3, was developed. It describes a "sock-technique", taking
advantage of the drying and swelling behaviour of the collagen
described before. Various manipulations can be performed in order
to get a structure with single layers or multiple layers, by
pulling or pushing over a wetted structure on top of the other. In
FIG. 4 it is shown the technique and the final structure of a
double-layered and single-layered tube, wherein the former is
clearly more stable since it can hold its own weight without having
the lumen collapsing. Furthermore, tubes produced in this way can
also be opened up by e.g. blade cut, resulting in single- or
multi-layered sheets, which can be used as patches for surgical
use. A double-layered sheet, originating from an opened
double-layered tube using a scissor, is shown in FIG. 4H.
[0115] Additionally, the bioactivity of the non-uniform collagen
scaffolds could be enhanced by the addition of bioactive molecules
to the rehydration bath using one of the known platforms to
efficiently bind bioactive molecules to collagen resulting in a
controlled release from the collagen scaffold. Collagen scaffolds
prepared according to the described method are highly useful for
soft tissue engineering or surgical applications of organs/tissues
such as genitourinary system (i.e. for bladder augmentation,
vaginal and pelvic floor reconstruction, ureter and/or urethra
reconstruction), for vascular tissue engineering, for the
reparation of the oesophagus or as patches for hernia repair.
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