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Journal of Neural Engineering
PAPER
Decellularized porcine peripheral nerve based
injectable hydrogels as a Schwann cell carrier for
injured spinal cord regeneration
To cite this article: Gopal Agarwal
et al
2024
J. Neural Eng.
21 046002
View the article online for updates and enhancements.
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J. Neural Eng. 21 (2024) 046002 https://doi.org/10.1088/1741-2552/ad5939
Journal of Neural Engineering
RECEIVED
29 January 2024
REVISED
20 May 2024
ACC EPT ED FOR PUB LICATI ON
17 June 2024
PUBLISHED
1 July 2024
PAPER
Decellularized porcine peripheral nerve based injectable hydrogels
as a Schwann cell carrier for injured spinal cord regeneration
Gopal Agarwal, Samantha Shumard, Michaela W McCrary, Olivia Osborne, Jorge Mojica Santiago,
Breanna Ausec and Christine E Schmidt∗
J. Crayton Pruitt Family Department of Biomedical Engineering, Herbert Wertheim College of Engineering, University of Florida,
Gainesville, FL 32610, United States of America
∗Author to whom any correspondence should be addressed.
E-mail: schmidt@bme.ufl.edu
Keywords: Schwann cells, injectable hydrogel, porcine peripheral nerves, decellularization, delipidation, cell delivery, spinal cord injury
Supplementary material for this article is available online
Abstract
Objective. To develop a clinically relevant injectable hydrogel derived from decellularized porcine
peripheral nerves and with mechanical properties comparable to native central nervous system
(CNS) tissue to be used as a delivery vehicle for Schwann cell transplantation to treat spinal cord
injury (SCI). Approach. Porcine peripheral nerves (sciatic and peroneal) were decellularized by
chemical decellularization using a sodium deoxycholate and DNase (SDD) method previously
developed by our group. The decellularized nerves were delipidated using dichloromethane and
ethanol solvent and then digested using pepsin enzyme to form injectable hydrogel formulations.
Genipin was used as a crosslinker to enhance mechanical properties. The injectability, mechanical
properties, and gelation kinetics of the hydrogels were further analyzed using rheology. Schwann
cells encapsulated within the injectable hydrogel formulations were passed through a 25-gauge
needle and cell viability was assessed using live/dead staining. The ability of the hydrogel to
maintain Schwann cell viability against an inflammatory milieu was assessed in vitro using
inflamed astrocytes co-cultured with Schwann cells. Main results. The SDD method effectively
removes cells and retains extracellular matrix in decellularized tissues. Using rheological studies,
we found that delipidation of decellularized porcine peripheral nerves using dichloromethane and
ethanol solvent improves gelation kinetics and mechanical strength of hydrogels. The delipidated
and decellularized hydrogels crosslinked using genipin mimicked the mechanical strength of CNS
tissue. The hydrogels were found to have shear thinning properties desirable for injectable
formulations and they also maintained higher Schwann cell viability during injection compared to
saline controls. Using in vitro co-culture experiments, we found that the genipin-crosslinked
hydrogels also protected Schwann cells from astrocyte-mediated inflammation. Significance.
Injectable hydrogels developed using delipidated and decellularized porcine peripheral nerves are a
potential clinically relevant solution to deliver Schwann cells, and possibly other therapeutic cells,
at the SCI site by maintaining higher cellular viability and increasing therapeutic efficacy for SCI
treatment.
1. Introduction
In the United States alone, ∼300 000 patients live
with traumatic spinal cord injury (SCI) [1]. SCI can
result from contusion, compression, laceration, or
stretch of spinal cord tissue due to traumatic insult,
which leads to permanent loss of function and sen-
sation throughout the body [2]. Presently, there is no
effective treatment that results in functional recovery
after SCI. Surgical decompression and methylpred-
nisolone administration are the clinical standard-of-
care for early acute SCI management, but they show
no effect on functional recovery in SCI patients [3].
Because of the lack of effective therapies to treat acute
or subacute SCI, more than 95% of patients with SCI
are in the chronic phase of injury, when glial scar is
© 2024 IOP Publishing Ltd
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
fully developed, and spontaneous behavioral recovery
has reached its plateau [4]. In human clinical trials,
cell-based treatments for SCI are yet to provide repro-
ducible evidence for clinical efficacy [5]. As noted
from preclinical studies, the primary reason for lim-
ited cell efficacy is due to poor cell viability and short
residence at the SCI site [6,7]. To address these issues,
in this study, we developed an injectable hydrogel
using porcine peripheral nerves to increase Schwann
cell efficacy by maintaining higher cellular viability
during injection and increasing residence time post
injection at the target site.
Autologous human Schwann cell transplantation
is currently being investigated in U.S. Food and
Drug Administration clinical trials for SCI patients
with thoracic and cervical SCI [8]. Phase I human
clinical trials have demonstrated the feasibility and
safety of transplantation of human Schwann cells
with reduction in cyst volume [9]. In rodent SCI
models, transplanted Schwann cells secrete growth
factors and extracellular matrix components such
as collagen and laminin, ultimately providing a
physical and bioactive substrate to promote axonal
growth [10,11]. Transplanted Schwann cells have
also been shown to support new growth of sensory
and propriospinal axons, myelinate regenerating, or
demyelinated axons, and restore axonal conduction
[10]. In addition, Schwann cell transplantation has
been shown to improve functional recovery in pre-
clinical subacute and chronic SCI models [11].
However, direct local injection of Schwann cells into
SCI results in significant transplanted cell loss and
death [12], which limits the extent of the cells’ thera-
peutic action. Previous studies have shown that after
a contusion injury in rats, only ∼20% of trans-
planted cells survive at 1 week and fewer than ∼5%
of Schwann cells survive 1 month after transplant-
ation. The harsh SCI microenvironment with high
levels of oxidative stress, inflammation, and immune
response results in a low survival rate. In addition,
typically up to ∼60% of cells fail to even reach the
target site, which is attributed to cell membrane dam-
age during injection and cell reflux out of the spinal
cord [6,7,13]. Because Schwann cell retention has
been correlated with decreased SCI cavitation and
increased symptomatic relief, increasing the number
of viable transplanted cells early after delivery will
likely improve therapeutic efficacy [7,13]. To address
this issue, we developed a clinically relevant inject-
able hydrogel formulation using decellularized por-
cine peripheral nerves.
Injectable hydrogels have the capacity to undergo
sol-to-gel transition after in vivo injections at desired
sites, resulting in an in situ solid hydrogel that con-
forms to the injury site [14]. Injectable hydrogels are
in a flowable state before injection, allowing various
water-soluble substances like drugs, proteins, cells,
and genes to be easily loaded within the pre-gel inject-
able solution; thus, injectable hydrogels can be used as
scaffolds or carriers for release of various therapeutic
substances [15,16]. Recently, injectable hydrogels
have gained wide attraction, especially in biomed-
ical applications, since injectability provides a way to
deliver hydrogels with minimally invasive procedures
[17]. Chemical and physical crosslinking strategies
have been used widely for the fabrication of inject-
able hydrogels for SCI treatment [18]. Unfortunately,
chemical crosslinking of injectable hydrogels typic-
ally requires use of toxic crosslinking agents, which
limits clinical translational potential [15]. Physical
gelation of hydrogels, on the other hand, is usually
caused by changing intermolecular forces (e.g. hydro-
gen bonding, hydrophobic interactions, electrostatic
ionic forces) between different polymer chains, which
can be achieved by changes in physical stimuli such
as temperature, pH, ionic strength, or pressure [15,
19]. However, physically crosslinked hydrogels typ-
ically suffer from weak mechanical strength. To
address this, in this study, we developed chemically
crosslinked injectable hydrogels that mimic central
nervous system (CNS) tissue mechanical strength by
using decellularized porcine peripheral nerves and
natural non-toxic crosslinking agents.
Decellularized extracellular matrix (dECM) has
been widely used for over 20 years as regenerative
biomaterial scaffold material because of its unique
biological activity and excellent biocompatibility [20,
21]. Decellularized scaffolds are immunologically tol-
erated matrices that retain the native structure, tis-
sue cell adhesion proteins, growth factors, and glyc-
osaminoglycans to provide a substrate for appropri-
ate tissue remodeling in the host [22,23]. Hydrogels
derived from decellularized tissue matrix have gained
wide attention because of their higher tunability,
injectability, range of mechanical properties, and
ease for modification with bioactive molecules [24].
Decellularized porcine-based hydrogels have been
shown to promote peripheral nerve regeneration
[24–26]. However, in most studies, porcine-based
dECM has been used as a filler or for coating of
electrospun nanofibrous scaffolds [27–29]. In addi-
tion, the potential of decellularized porcine peri-
pheral nerve-based hydrogels as injectable delivery
vehicles for cells or drugs is still unexplored.
Various chemical and enzyme-based decellular-
ization methods have been used for porcine peri-
pheral nerves [24,30]. Also, other porcine tissues have
been decellularized, but since each tissue has a dif-
ferent structure and composition, the decellulariza-
tion methods are tissue specific [31]. To keep tissue
as intact as possible after decellularization, chemical
methods need to be chosen carefully to minimize the
damage caused to structural macromolecules of the
ECM [31].
Previously, our group developed injectable hydro-
gels derived from decellularized rat peripheral nerves
using a prior decellularization method (the ‘Hudson
method’) and have evaluated the potential for the
2
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
Figure 1. Schematic summary of the project: Porcine peripheral nerves (sciatic and peroneal) were decellularized using the SDD
method [37] and then subsequently delipidated using dichloromethane and ethanol solvent and digested using pepsin enzyme.
The digested ECM was neutralized to undergo gelation and then in some studies, genipin was added as a chemical crosslinker to
modify mechanical stiffness. The genipin-crosslinked hydrogels were then analyzed for injectability and Schwann cell delivery.
(Image credit: BioRender.com).
resulting hydrogels to improve functional out-
comes as cell carriers in SCI [20,32–34]. Our lab
also modified our prior Hudson decellularization
method and developed a novel sodium deoxycholate-
deoxyribonuclease (SDD) method for rat peripheral
nerves and evaluated the potential of resulting
hydrogels as injectable delivery vehicles for neural
applications [35,36]. However, the SDD decellular-
ization method has not yet been applied to porcine
peripheral nerves. In this study, we have evaluated
the use of the SDD method for decellularization of
porcine peripheral nerve branches (peroneal and sci-
atic) and fabrication of injectable hydrogels. Also, we
studied the effect of remaining lipids in decellularized
porcine peripheral nerve branches on gelation kinet-
ics of injectable hydrogels. Moreover, we have studied
the effect of genipin-mediated chemical crosslinking
of injectable porcine peripheral nerve formulations to
tune the mechanical strength of hydrogels. The effic-
acy of developed injectable hydrogel formulations
to promote Schwann cell viability during injection
and to protect the injected Schwann cells from the
local inflammatory microenvironment of the SCI was
also explored (figure 1). Overall, this study provides
insight into the development of an injectable hydrogel
formulation using decellularized porcine peripheral
nerves, which can be used for cell delivery in SCI
tissue regeneration.
2. Materials and methods
2.1. Decellularization of porcine peripheral nerves
Peripheral nerve branches (peroneal and sciatic) were
isolated aseptically from mini pigs, supplied by the
University of Florida (UF) animal care facilities.
Nerves were not isolated from live pigs; post-mortem
animals were used for nerve isolation. After isola-
tion, the nerves were kept at −30 ◦C until further
use. Decellularization of peripheral nerves was adap-
ted from methods previously used for rodent sci-
atic nerves [35]. All the chemicals were of standard
quality unless specified. Briefly, nerves were subjec-
ted to 7 h of washing in water, followed by 18 h
of washing in 125 mM sulfobetaine (SB-10) (Sigma
Aldrich, D4266) buffer. Thereafter, the nerves were
washed with 100 mM sodium/10 mM phosphate buf-
fer for 15 min, and then with 3% sodium deoxy-
cholate (SD) (Sigma Aldrich, D6750) /SB-16 (Sigma
Aldrich, H6883) solution for 1.5 h, with 50 mM
Na/10 mM phosphate buffer thrice, for 15 min each.
To remove DNA content, the nerves were subjected
to 75 U ml−1of DNAase solution (Sigma Aldrich,
D4527) for 3 h. Subsequently, the nerves were washed
with 50 mM Na/10 mM phosphate buffer thrice, for
1 h each. The nerves were then washed once with
0.2 U ml−1of Chondronitase ABC enzyme (Sigma
Aldrich, C3667) solution. Finally, the nerves were
washed with 1X PBS for 3 h and stored at −30 ◦C until
further use. To understand the effect of remaining lip-
ids after decellularization on injectable hydrogel for-
mulation and gelation kinetics, a part of decellular-
ized nerve piece was subjected to ethanol (Fischer
Scientific, NC1150367) and dichloromethane (Sigma
Aldrich, 270997-2l) in a 1:2 ratio for 24 h to remove
residual lipids, and then washed with distilled water
thrice, 30 min each. The decellularized and delip-
idated nerves were then lyophilized and stored at
−80 ◦C until further use. These nerves are referred
to as ‘decell +delipid’ in the rest of the manuscript.
Decellularized nerves without delipidation were used
3
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
as controls and referred to as ‘decell only’ in the
manuscript.
2.2. dsDNA quantification assessment
Fresh and decell +delipid nerves were lyophilized for
72 h and tissue was weighed, and DNA was isolated
using a Quick-DNA Purification Kit (Zymo Research;
D4068). Briefly, the tissue samples were incubated in
proteinase K solution overnight at 55 ◦C, after which
the solution was centrifuged at 12 000 g for 1 min
and the supernatant was collected and used for fur-
ther steps. DNA binding solution was added to the
supernatant and the combined solution was added to
Zymo spin columns provided by the manufacturer.
The DNA was then washed using DNA washing buffer
provided by the manufacturer and finally, DNA was
eluted into 50 µl of DNA elution buffer provided by
the manufacturer. The isolated DNA was then quanti-
fied using Quant-iT PicoGreen™ dsDNA reagent kit
(Invitrogen; P11496) as per the manufacturer’s pro-
tocol. The isolated DNA amount was calculated by
extrapolating from a standard curve using lambda
DNA.
2.3. Immunohistochemistry analysis
For immunohistochemistry, after decellularization,
nerves were fixed using 4% paraformaldehyde and
stored in 30% sucrose for 1 week. Nerves were
then trimmed, soaked in optimal cutting temper-
ature (OCT) compound at 4 ◦C overnight, and
flash frozen at −80 ◦C for at least 5 min. OCT
blocks were then stored at −20 ◦C until section-
ing. Frozen sections were obtained using a Leica
CM1950 cryostat (Leica Biosystems, Germany)
with a tissue thickness of 10 µm. Sections were
stained using antibodies against collagen-I (Col-I)
(Millipore Sigma; C2456, 1:500), collagen-IV (Col-
IV) (Abcam; ab6586, 1:500), laminin (Millipore
Sigma; L9393, 1:500), β-III tubulin for axons
(Abcam; ab7751, 1:1000), S-100 for Schwann cells
(Millipore Sigma; S2644, 1:500), myelin basic pro-
tein (MBP) for myelin (Millipore Sigma; M3821,
1:150), and 4′,6-diamidino-2-phenylindole (DAPI)
for nuclei (Thermo Fisher Scientific; D1306, 1:1000).
Negative controls (no primary antibody) were pre-
pared to ensure autofluorescence or nonspecific
staining was minimal.
2.4. Decellularized nerve digestion
Decellularized peripheral nerves were solubilized
using the SDD method published earlier by our
lab [36]. Immediately after the delipidation treat-
ment, processed nerves were subject to three 30-
minute washes in sterile ddH2O on an orbital rotator,
flash-frozen in liquid nitrogen, and then lyophilized
(Labconco) for 3 d. Lyophilized nerves were weighed
and minced into small pieces in a scintillation vial
(Millipore Sigma). Pepsin at 1 mg ml−1(Sigma
Aldrich P7012) in 0.01 M hydrochloric acid (Sigma
Aldrich 258 148) was added to achieve tissue con-
centrations of 13.4 mg of tissue per 1 ml of pepsin
solution. Vials were capped and incubated for 72 h at
room temperature under constant agitation through
use of a magnetic stir bar. Nerve digests, which at
this point were homogenous solutions, were trans-
ferred from scintillation vials into microcentrifuge
tubes and placed on ice. Then, 0.1 M NaOH was
added to the digests until the pH reached 7.4 to neut-
ralize the solution, and the solution was stabilized by
adding 10X PBS at an amount 1/9th the volume of the
digest. The neutralized solution is referred to as a ‘pre-
gel’ solution and was used for lipid estimation and
rheometry as explained below. Decellularized nerves
without delipidation treatment were digested simil-
arly and used as controls.
2.5. Lipid estimation in decellularized nerves
To understand the effect of remaining lipids on
gelation kinetics of injectable hydrogels, we initially
estimated lipid content in decellularized nerves using
the Lipid Assay Kit (unsaturated fatty acids) supplied
by Abcam (catalog number: ab242305). The Lipid
Assay Kit allows measurement of the lipid content
(unsaturated fatty acids only) of samples using sulfo-
phospho vanillin reagent by colorimetric absorbance.
Briefly, nearly 15 µl of the pre-gel solution both decel-
lularized with and without delipidation treatment
were placed into a 96-well plate, and then incubated
at 90 ◦C for 30 min. The samples were transferred to
4◦C for 5 min, following which 150 µl of 18 M sul-
furic acid was added to each well. The samples were
then incubated at 90 ◦C for 10 min, then at 4 ◦C for
5 min. 100 µl of each sample was then transferred
into a clean 96-well plate and background absorb-
ances were read at 540 nm. Following this, 100 µl
of sulfo-phospho vanillin reagent was added to each
well and mixed well with a pipette. The samples were
then incubated at 37 ◦C for 15 min, and absorb-
ance was read at 540 nm. The background absorbance
was subtracted from each well. The concentrations of
lipids in samples were extrapolated from a standard
curve plotted using the lipid standard supplied by the
manufacturer.
2.6. Crosslinking of decellularized nerve pre-gel
solution
Pregel solutions were crosslinked individually using
different concentrations of genipin (Sigma Aldrich
G4796). Different concentrations of genipin (dis-
solved in DMSO, stock 89 mM) were added to
10 mg ml−1neutralized pre-gel solution to achieve
final concentrations of 1, 5, 10 mM genipin and
incubated at 37 ◦C to undergo gelation.
2.7. Rheological analysis
Frequency sweep and shear thinning experiments
were carried out using an Anton Paar MCR 302 in
parallel plate geometry with a 1 mm gap and 8 mm
4
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
sandblasted parallel plate. For frequency sweep stud-
ies, the hydrogels were cast in 8 mm silicone molds
by pouring 125 µl of pre-gel neutralized solution
with/without genipin crosslinker. Frequency varied
from 0.01 to 100 Hz at a constant 1% strain at 25 ◦C,
and the storage (G′) and loss (G′′) moduli were
recorded. Shear thinning experiments were conduc-
ted by placing 125 µl of neutralized pre-gel solution
with/without crosslinker and the shear rate was var-
ied from 0–150 Pa and temperature was constant at
25 ◦C and viscosity was recorded.
2.8. In vitro astrocyte-Schwann cell co-culture
analysis
Rat astrocytes (P 2–3) and rat Schwann cells (P 2–4)
were cultured in astrocytes/Schwann media provided
by Sciencell manufacturer, with addition of 10%
fetal bovine serum and 1% penicillin-streptomycin.
The cells were cultured in 5% CO2at 37 ◦C, and
media was changed every 2–3 d. Initially, to determ-
ine efficient lipopolysaccharide (LPS) concentration
for inflamed astrocytes, different concentrations of
LPS (0, 1, 5, and 10 µg ml−1) were added to the
astrocyte culture media and exposed to ∼1×105
rat astrocytes cultured in 6-well plates for 24 h.
Thereafter, the rat astrocytes were fixed with 4% par-
aformaldehyde (PFA) for 15 min and immunostained
using glial fibrillar acidic protein (GFAP) marker.
The GFAP intensity was measured using Image J
software to determine accurate LPS concentration
for inflammatory astrocytes. For co-culture experi-
ments, ∼0.05 ×106rat astrocytes were cultured on
top of 24 well culture inserts, while ∼0.05 ×106
rat Schwann cells within 10 mg ml−1hydrogel or
alone were cultured on bottom of 24-well cell cul-
ture plate. 5 µg ml−1LPS in 1:1 Schwann: Astrocyte
complete media was added to the well and incub-
ated for 48 h. For rat Schwann cells within hydrogels,
15 mg ml−1decellularized peripheral nerves were
minced into small pieces using sterilized micro scis-
sors and digested in pepsin solution (1 mg ml−1)
for 48 h. After complete digestion, the solution was
neutralized using appropriate volume of 0.2 µm syr-
inge filtered 1 N NaOH and 10x PBS. An appro-
priate volume of rat Schwann cell suspension hav-
ing ∼0.05 ×106rat Schwann cells in each hydrogel
was added to the neutralized pre-gel solution and a
hydrogel concentration of 10 mg ml−1was achieved
by adding cell culture media. The pre-gel solutions
with cells were cast into 8 mm silicone molds and
incubated at 37 ◦C, 5% CO2for 20 min. For genipin-
crosslinked hydrogel groups, the required volume
of genipin was added to neutralized pre-gel solu-
tion containing cell suspension. After 48 h of incub-
ation, the cells within the hydrogel or alone were
stained using live/dead stain as per Thermo Fisher
manufacturer recommendations. The hydrogel/cells
were then imaged using a Leica 880 confocal, and
the live/dead cell ratios were analyzed using Image J
software.
2.9. Injectability analysis
To determine the injectability of the hydrogel solu-
tion, 10 mg ml−1neutralized pre-gel solution with
or without genipin solution was passed through a 25-
G 5/8-inch syringe at room temperature in a 96-well
plate.
2.10. In vitro Schwann cell post-injection and rat
spinal cord neuron biocompatibility analysis
Decell +delipid peroneal and sciatic nerve hydro-
gels were digested and neutralized to pH 7.4, as pre-
viously explained above. Rat Schwann cells (P 4–5,
1×105) were added to neutralized pre-gel solutions
and final concentrations of hydrogels were adjus-
ted to 10 mg ml−1using Schwann cell culture com-
plete media. 1 mM genipin was added to the hydro-
gel cell solution and then the solution was passed
through a 25-G needle and the injected solution was
cast into 8 mm silicone mold, incubated at 37 ◦C,
5% CO2for 30 min. The hydrogels were cultured up
to 7 d and cell viability was assessed using live/dead
stain (LIVE/DEAD™ Viability/Cytotoxicity Kit, for
mammalian cells, L3224, Invitrogen) at different time
points post injection (30 min, days 4 and 7). Rat
Schwann cells in saline were also processed sim-
ilarly and used as controls. Additionally, we ana-
lyzed the rat spinal cord neuron viability. Briefly,
∼5×103rat spinal cord neurons (procured from
Sciencell) /hydrogel were added to a neutralized
hydrogel formulation and then the solution was cast
into 8 mm silicone molds. The hydrogels were then
incubated at 37 ◦C, 5% CO2for 30 min to undergo
complete hydrogel gelation. The hydrogels were then
transferred into 48 well plates containing 1 ml of
neuronal culture media (procured from Sciencell)
with 1% Pen-Strep and 1x nerve growth solution.
After 48 h of incubation, live/dead staining was per-
formed as explained above. The live/dead cell ratio
was then analyzed using Image J software. Briefly,
the images were exported to Image J software, and
converted to 8 bits and the same threshold values
were applied to all images. The number of cells were
then counted using Image J software Analyze tool-
bar. Following that, the ratio of live/dead cells were
calculated.
3. Statistical analysis
All experiments were performed with n=3
in each group, and data were analyzed using
GraphPad Prism. Wherever applicable, the One-
Way ANOVA Tukey ttest was applied, or
unpaired ttest was applied to analyze signific-
ant differences between the groups. p<0.05
was considered as significant difference between
the groups.
5
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
Figure 2. (A) Immunostaining images representing efficient cell removal using SDD method: DAPI (blue), β-tubulin III (yellow),
S100 (orange) and MBP (green) staining to analyze the removal of various cellular contents in decellularized nerves. (Scale
bar =50 µm) (B) DNA quantification of fresh and decell +delipid nerves using Quant-it Pico Green kit. Data were analyzed
using Two Way ANOVA, Tukey ttest, n=2.
4. Results
4.1. SDD method efficiently removes cells in
decellularized nerves
Decellularization is the process of efficiently remov-
ing cells and cellular membrane remnants while
retaining ECM components [22,36]. We initially ana-
lyzed the proficiency of Schwann cell, axon, myelin,
and DNA removal using immunostaining for vari-
ous markers, S100, β-tubulin III, myelin basic pro-
tein (MBP) and DAPI, respectively. As shown in
figure 2(A), DAPI staining revealed that both sci-
atic and peroneal nerves decellularized using the SDD
method had no staining for nuclei, as compared to
fresh nerves. Similarly, the decellularized nerves had
significantly lower staining for β-tubulin III, S100,
and MBP, demonstrating efficient removal of axons,
Schwann cells, and myelin, respectively. Overall, these
results demonstrate that the SDD method can be
used to successfully decellularize porcine peripheral
(sciatic and peroneal) nerves. Similar results for the
SDD method to efficiently remove cells were seen for
rat peripheral nerves [35]. Furthermore, as shown in
figure 2(B), both decell +delipid sciatic and peroneal
nerves have ∼23 ng dsDNA/mg of tissue, compatible
with current decellularization guidelines of <50 ng
dsDNA/mg of tissue, proposed by Crapo et al [38].
4.2. SDD method efficiently retains ECM in
decellularized nerves
The presence of ECM proteins like collagen I, collagen
IV, fibronectin, and laminin in decellularized nerves
plays an important role in cellular adhesion, prolif-
eration and provides pro-regenerative cues for neur-
onal regeneration [39]. We analyzed the presence of
key ECM Col-I, Col-IV, and laminin in porcine sciatic
and peroneal decellularized nerves. ECM preserva-
tion was also analyzed to ensure that the SDD method
retains important ECM constituents in decellularized
nerves. As shown in figure 3, the SDD method effect-
ively retained and preserved ECM structure in decel-
lularized nerves.
4.3. Effects of residual lipids on gelation kinetics
and mechanical strength of hydrogels
We further analyzed the ability of pepsin-digested
solutions of decellularized nerves to undergo gela-
tion at body temperature, which is essential for
an in-situ gelling, injectable formulation. The SDD
method, when previously used to decellularize rat
peripheral nerves, resulted in forming an injectable
formulation undergoing sol-to-gel transition at body
temperature [36]. However, decellularized porcine
peripheral nerves failed to undergo efficient sol-to-gel
transition, resulting in weak and unstable hydrogels.
6
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
Figure 3. Immunostaining images representing efficient ECM retention in decellularized nerves: Col-IV, Col-I and laminin
staining to analyze the ECM structure maintenance in decellularized nerves (Scale bar =50 µm).
We studied the effect of lipid remnants in decellular-
ized porcine nerves on the gelation kinetics of hydro-
gels. To effectively reduce lipid remnants, the decel-
lularized porcine nerves (using SDD method) were
subjected to dichloromethane and ethanol solvent
for 24 h and are referred to as ‘decell +delipid’,
whereas only SDD decellularized nerves are addressed
as ‘decell only’ in the manuscript.
As shown in figures 4(A) and (B), dichloro-
methane and ethanol solvent treatment effectively
reduced the lipid remnants in decellularized nerve.
These decell +delipid nerves were digested using
pepsin solution and neutralized to form a pre-gel
solution. Following that, the pre-gel solution was
loaded on the rheometer stage at 37 ◦C and using
frequency sweep, G′and G′′ were recorded. G′
values of decell +delipid hydrogels for both sci-
atic and peroneal samples were higher than values
for decell only hydrogels (figures 4(C) and (D)).
Furthermore, decell +delipid sciatic and peroneal
hydrogels had no increase in G′after approximately
13.5 min, indicating complete gelation at body tem-
perature. Furthermore, at frequency 1 rad s−1, G′of
decell +delipid sciatic and peroneal hydrogels were
significantly higher than decell sciatic and peroneal
hydrogels only (figures 4(E) and (F)). Therefore, the
presence of lipid remnants in decellularized nerves
hinders the gelation kinetics of hydrogels and results
in lower mechanical strength (lower storage modu-
lus) of hydrogels.
On analyzing the G′with respect to temperature
sweep, we found that there was no significant differ-
ence between decell +delipid and decell only sciatic
and peroneal hydrogels. The decell +delipid sciatic
hydrogel started to gel at 30 ◦C, as evidenced by a sud-
den increase in G′and became stable at 37 ◦C, shown
as steady G′(figure 4(G)). Similarly, for peroneal
decell +delipid hydrogels, G′started to increase at
26 ◦C while it had stable G′at 37 ◦C (figure 4(H)).
Overall, this study indicates that delipidation of
porcine peripheral nerves is crucial for forming a
stable injectable hydrogel formulation, which is con-
sistent with prior studies demonstrating that delip-
idation of other fatty tissues improves mechanical
strength and gelation kinetics of hydrogels [38].
However, our prior research shows that hydrogels
with a mechanical strength of 2000 Pa would best
mimic CNS tissue mechanical strength [40]. Here,
we found that even 10 mg ml−1of decell +delipid
sciatic and peroneal hydrogels had lower mechan-
ical strength than desired and increasing the concen-
tration of hydrogel solution more than 10 mg ml−1
would further increase the viscosity of the hydro-
gel solution, making it difficult to pass through a
syringe. Therefore, to achieve a desired mechan-
ical strength of around 2000 Pa, we crosslinked the
decell +delipid hydrogels using the natural crosslink-
ing agent genipin.
4.4. Genipin crosslinking tunes mechanical
strength of hydrogels and affects injectability and
gelation kinetics
Genipin is a natural and low toxicity crosslinking
agent, derived from gardenia fruit, that can crosslink
free amino groups of lysine or hydroxylysine residues
on different polypeptides chains [41]. We initially
analyzed different genipin concentrations (0, 1, 5,
10 mM) on the mechanical strengths of peroneal and
7
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
Figure 4. Effect of lipids and delipidation on gelation kinetics and mechanical strength of hydrogels: Sulfo vanillin reagent-based
lipid estimation (A) sciatic decell +delipid and decell only (B) peroneal decell +delipid and decell only. Frequency sweep studies
for (C) sciatic decell +delipid and decell only (D) peroneal decell +delipid and decell only. G′at 1 rad s−1frequency of (E)
sciatic decell +delipid and decell only (F) peroneal decell +delipid and decell only. Data are analyzed using ttest, n=3.
Temperature sweep studies of (G) sciatic decell +delipid and decell only (H) peroneal decell +delipid and decell only.
sciatic decell +delipid hydrogels. To easily differenti-
ate between crosslinked and uncrosslinked hydrogels,
we have now referred to ‘decell +delipid’ hydrogels as
‘uncrosslinked’. As shown in figures 5(A) and (B), an
increase in genipin concentration was shown to stat-
istically increase the mechanical strength of hydro-
gels, as evident by an increase in G′. Uncrosslinked
hydrogels that are decell +delipid hydrogels had
lower mechanical strength compared to genipin-
crosslinked gels, as shown in figure 5(A). However, as
hydrogels of around 2000 Pa are desired for CNS tis-
sue regeneration, sciatic and peroneal decell +delipid
hydrogels crosslinked with 1 mM genipin concentra-
tion with mechanical strengths of around 2000 Pa
were selected for further studies.
We further analyzed the gelation kinetics of both
uncrosslinked and 1 mM genipin-crosslinked hydro-
gel solutions using rheometer time sweep analysis. As
shown in figure 5(C), uncrosslinked decell +delipid
sciatic hydrogel started gelling within 2.5 min of
incubating at 37 ◦C as evidenced by an increase
in G′, while the hydrogel was fully gelled within
15 min of incubation at 37 ◦C, with steady G′. The
1 mM genipin crosslinking of decell +delipid sciatic
8
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
Figure 5. Genipin crosslinking tunes mechanical strength of hydrogels and affects injectability and gelation kinetics: Effect of
genipin-mediated crosslinking on decell +delipid hydrogels, (A) sciatic (B) peroneal. Data were analyzed using One Way ANOVA
Tukey ttest, n=3. Time sweep experiment to analyze the effect of genipin mediated crosslinking on decell +delipid hydrogel,
(C) sciatic (D) peroneal. Shear thinning experiment to demonstrate injectability of decell +delipid, (E) sciatic (F) peroneal.
hydrogels had faster gelation kinetics as evident by
an increase by G′and after around 10 min of incub-
ation at 37 ◦C, the hydrogels were fully gelled as
evident by stable G′(figure 5(C)). Similarly, uncross-
linked decell +delipid peroneal hydrogels started
gelling within 7.5 min of incubation at 37 ◦C, and
fully gelled within 15 min. 1 mM genipin crosslink-
ing of decell +delipid peroneal hydrogels speeds
the gelation kinetics and G′starts to increase within
3 min and becomes stable within 15 min of incub-
ation at 37 ◦C (figure 5(D)). This study shows that
the addition of genipin enhances the gelation kinetics
of decell +delipid hydrogels and both uncrosslinked
and 1 mM genipin-crosslinked sciatic and peroneal
hydrogels fully gel within 15 min of incubation at
37 ◦C.
On analyzing the injectability of the solution
using a rheometer, there was a decrease in viscosity
of the solution for both peroneal and sciatic uncross-
linked and 1 mM genipin decell +delipid hydrogels
with an increase in shear strain (figures 5(E) and
(F)). This demonstrates injectability of the hydrogel
as evidenced in previous literature [42]. As seen in
figure 5(F), peroneal uncrosslinked injectable hydro-
gels at lower shear stress showed an increase in the
viscosity of the solution, demonstrating shear thick-
ening. At lower shear stresses, shear thickening is
observed indicating that the rate of bond formation
is slower than the chain relaxation time, however, at
higher shear stresses, the solution exhibits shear thin-
ning property. Similar behaviors have been observed
by other injectable hydrogels [43]. Furthermore,
we tested the injectability of peroneal and sciatic
decell +delipid uncrosslinked and 1 mM genipin-
crosslinked hydrogels using 25-G needles, the most
widely used needle for cell delivery [9,10,44]. As
9
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
Figure 6. Injectable hydrogel maintains Schwann cell viability postinjection: (A) Schematic representative of overall experiment
performed: Schwann cells encapsulated within injectable hydrogel formulations were passed through 25-G needles fitted into 1 ml
syringes. The cells were stained with live/dead stain and imaged using confocal microscopy immediately after 30 min and at days 4
and 7. (B) Representative confocal images of live (green) and dead (red) stained cells after 30 min and days 4 after passing through
the 25-G needle. Scale bar =200 µm. The red arrows denote dead cells; green arrows denote live cells. (C) Bar graph showing
live/dead ratio of Schwann cells, analyzed using Image J software. Data were analyzed using Two Way ANOVA, and Bonferroni
test, compared vs saline control, n=3. ∗∗∗p<0.0001.
shown in supplementary video S1–4, both uncross-
linked and 1 mM genipin-crosslinked peroneal and
sciatic hydrogel solutions were able to pass through
syringes easily, demonstrating the suitability of these
hydrogel solutions for cell delivery for spinal cord
regeneration [44,45].
4.5. Injectable hydrogels maintain Schwann cell
viability post injection
We initially analyzed rat spinal neuron viability
post 48 h of culturing within the hydrogels using
live/dead stain. As shown in figure S3, both peroneal
and sciatic decell +delipid uncrosslinked and
genipin-crosslinked hydrogels were able to sup-
port rat spinal neuron viability with no significant
difference between any groups. Thereafter, we ana-
lyzed the ability of hydrogels to efficiently deliver
cells. Cell injection procedures typically result in
high cell death, with viabilities as low as 1%–32%
post transplantation [46]. Therefore, we initially
analyzed Schwann cell viability post injection using
our injectable hydrogel formulations. Schwann cells
were mixed with both uncrosslinked and 1 mM
genipin peroneal and sciatic hydrogel formulations
and passed through 25-G needles. Schwann cell viab-
ility was analyzed post injection at 30 min, 4 d,
and 7 d using live/dead staining (figure 6(A)). As
shown in figures 6(B) and (C), Schwann cells had
a statistically higher live/dead cell ratio in hydro-
gels as compared to saline controls at 30 min and
10
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
4 d post injection. There was no significant differ-
ence between any of the hydrogel formulations, and
both uncrosslinked and 1 mM genipin peroneal and
sciatic hydrogel formulations maintained similar cel-
lular viability post injection. Additionally, Schwann
cells within uncrosslinked and genipin-crosslinked
hydrogels stained at 30 min of incubation post injec-
tion have a rounded morphology, because the cells
were non adherent, whereas on day 4 the cells star-
ted attaining elongated morphologies. On day 4,
based on qualitative assessment, Schwann cells within
genipin-crosslinked hydrogels were less elongated,
whereas cells within uncrosslinked hydrogels were
highly elongated. The difference in substrate stiff-
ness may result in such a difference in Schwann cell
morphology [47]. However, on day 7 (figure S1) there
was no difference in live/dead ratios between hydrogel
formulations, as compared to saline controls.
4.6. Injectable hydrogel maintains Schwann cell
viability against inflammation
After SCI, astrocytes, resident glial cells in the
spinal cord, secrete pro-inflammatory cytokines and
chemokines, and are proposed as mediators of SCI
inflammation and scarring [44]. Therefore, we ana-
lyzed the effects of astrocyte-mediated inflammation
on Schwann cell viability. Initially, we determined
the adequate LPS concentration to inflamed astro-
cytes cells in vitro. Different concentrations of LPS
(0, 1, 5, 10 mM) were added to cultured astrocytes
and incubated for 24 h. Following that, GFAP intens-
ity was measured to determine astrocyte reactivity
using immunochemistry. Of all LPS concentrations
tested, we found that compared to the no LPS con-
trol group, astrocytes treated with 5 mM LPS had
higher GFAP intensity, indicating astrocyte reactivity
[48–50] (figures 7(A) and S2). Thereafter, in co-
culture experiments the effect of inflamed astrocytes
on Schwann cell viability was assessed. As shown in
figure 7(C), Schwann cells cultured on tissue culture
plastic (TCP) in the presence of inflamed astrocytes
exhibited higher cell death and a lower live/dead cell
ratio as compared to Schwann cells cultured on TCP
in the absence of inflamed astrocytes. Schwann cells
encapsulated within decell +delipid peroneal and
sciatic 1 mM genipin-crosslinked hydrogels exhibited
higher viability and fewer dead cells in the presence
of inflamed astrocytes as compared to Schwann cells
cultured on TCP in the absence of inflamed astro-
cytes (figure 7(D)). Therefore, this study indicates
that peroneal and sciatic 1 mM genipin-crosslinked
hydrogels protect Schwann cells from inflammation-
mediated cell death from activated astrocytes.
5. Discussion
Various methods have previously been stud-
ied to decellularize porcine sciatic nerves
[24,26,30,51,52]. However, developing injectable
hydrogel formulations using a decellularized porcine
matrix has remained challenging. Furthermore, the
use of decellularized porcine peripheral nerve-based
matrices for cell delivery in SCI treatment has not
been explored previously. In this study, we developed
an injectable hydrogel formulation using decellu-
larized porcine peripheral nerves for increasing the
efficiency of Schwann cell transplantation in SCI
treatment by maintaining cellular viability during
injection and mitigating inflammatory conditions in
SCI.
We initially evaluated the potential of our pre-
viously developed SDD chemical decellularization
method [35] for efficient decellularization of por-
cine peripheral nerve, specifically the sciatic and
peroneal branches. The SDD method had previ-
ously been successful [35] in the efficient decellu-
larization of rat peripheral nerves and fabrication
of injectable hydrogel formulations for neural tis-
sue regeneration [36]. After decellularization of por-
cine peripheral nerves (sciatic and peroneal) using the
SDD method, DAPI staining revealed the absence of
nuclei, while βIII tubulin, S100, and MBP staining
revealed the absence of any cellular content includ-
ing axons, Schwann cells, and myelin debris (figure 2).
The SDD method also resulted in efficient retention
of key ECM components and structure, as evidenced
by the presence of Col-I, Col-IV, and laminin staining
(figure 3). These results indicate the applicability of
the SDD method to decellularize porcine peripheral
nerves effectively. In addition, these results are con-
sistent with our previous literature, where the SDD
method was effective in decellularizing rat peripheral
nerves [35].
Upon using these decellularized matrices for
developing injectable hydrogels, we found that decel-
lularized porcine peripheral nerves formed a weak
and unstable injectable hydrogel formulation, unlike
our prior experience with rat nerves. Previous literat-
ure supports that the presence of lipids in decellular-
ized tissue hinders the gelation kinetics of hydrogels,
resulting in hydrogels with weak mechanical strength
[53,54]. Besides this, the presence of residual lip-
ids has been shown to inhibit target tissue regen-
eration by resulting in lower cellular removal effi-
ciency and residual cell debris that can cause immune
responses [55]. Similar results were seen in our study:
the presence of lipids in decellularized tissues resul-
ted in the formation of hydrogels with weak mechan-
ical strength, insufficient for mechanical handling. To
reduce lipid remnants in decellularized porcine peri-
pheral nerves, we delipidated the decellularized tissue
using dichloromethane and ethanol solvent. Various
other solvents like ethanol [56], 2-propanol, meth-
anol, and acetones [38,53] have been used previ-
ously to reduce lipid remnants for various decellular-
ized tissues. However, dichloromethane and ethanol
solvent treatment is most frequently used for redu-
cing lipid remnants in the decellularized spinal cord
11
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
Figure 7. Genipin-crosslinked hydrogel maintains Schwann cell viability in an in vitro inflammatory model: (A) Schematic
showing astrocyte-Schwann cell co-culture experiment with LPS as an in vitro inflammatory model (B) representative image
showing GFAP intensity in LPS and non-treated astrocytes (C) live/dead staining images of Schwann cells after 48-h incubation in
the presence of LPS activated astrocytes; live cells are represented by green staining and dead cells are represented by red staining.
TCP: tissue culture plastic seeded Schwann cells (D) live/dead cell ratio as analyzed by Image J software. Data are analyzed by One
Way ANOVA, Tukey t-test, n=3.
and sciatic nerve tissues [24,26,28,57]. In our study,
we also found that treatment of decellularized tissue
using dichloromethane and ethanol solvent signific-
antly reduced lipid remnants and aided in appropriate
hydrogel gelation (figure 4). Therefore, the delipida-
tion of decellularized porcine peripheral nerve tissue
is necessary for developing injectable hydrogel formu-
lations with sufficient mechanical strength suitable
for injectability.
However, even after delipidation of decellular-
ized porcine peripheral nerves, a 10 mg ml−1
hydrogel concentration resulted in a mechanical
strength less than 2000 Pa. But, for CNS tissue
regeneration, it is desired to have hydrogels of
around 2000 Pa to mimic CNS tissue mechan-
ical strength [40]. To achieve desired mechanical
strength, we crosslinked the decell +delipid hydro-
gels using a chemical crosslinker, genipin. Genipin
12
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
is a natural crosslinker that has been widely used by
others in the field to crosslink decellularized ECM
and genipin-crosslinked decellularized hydrogels are
biocompatible [58–60]. In our study, we found that
with an increase in genipin concentration, there was
an increase in mechanical strength of hydrogels as
evidenced by increases in G′(figure 5(A) and (B)).
The increase in mechanical strength can be attributed
to genipin-mediated crosslinking. Genipin crosslinks
with the free amino groups of lysine or hydroxylysine
residues present on the polypeptides in ECM [41]. As
1 mM genipin crosslinking of decell +delipid hydro-
gels provided hydrogels of around 2000 Pa mech-
anical strength, we proceeded with 1 mM genipin-
crosslinked hydrogels for our remaining studies.
Additionally, the decell +delipid hydrogels and both
uncrosslinked and 1 mM genipin-crosslinked hydro-
gels were able to undergo complete gelation within
15 min of incubation at 37 ◦C (figures 5(C) and (D)).
Furthermore, crosslinking with 1 mM genipin has
been shown to initially reduce the time for gelation
of the hydrogels (figures 5(C) and (D)). This gelation
time has previously been adequate for biomolecule
delivery in SCI animal models [61].
On analyzing the injectability of the hydro-
gels using a flow curve rheology study, we found
that both uncrosslinked and genipin-crosslinked
decell +delipid hydrogels exhibited a decrease in vis-
cosity with an increase in shear stress. Such shear thin-
ning property of the hydrogels demonstrates that the
injectable hydrogel formulations can easily be injec-
ted at the target sites without clogging the needles
(figures 5(E) and (F)). This was further demonstrated
by passing the decell +delipid uncrosslinked and
1 mM genipin-based sciatic and peroneal hydrogel
formulations through 25-G needle syringes. This size
syringe needle has been used for various cell delivery
applications in SCI treatment [44,45], and hence,
this study demonstrates that the decell +delipid
hydrogels are potentially viable options for
various clinical cell and biomolecule delivery
applications.
In previous pre-clinical animal studies, direct
local injection of Schwann cells in saline into the SCI
lesion cavity results in significant transplanted cell
death [12]. Besides this, regardless of the detrimental
host microenvironment, initial cell damage occurs
during the actual injection procedure. Cell damage
may occur during ejection because of the mech-
anical disruption of cells. When flowing through a
needle, cells experience various types of mechan-
ical forces, including extensional forces and shear
forces [62]. Hence, we initially assessed Schwann cell
viability when passed through a 25-G needle. We
found that as compared to saline controls, Schwann
cells had higher cellular viability in both uncross-
linked and 1 mM genipin sciatic and peroneal inject-
able hydrogel formulations (figures 6(B) and (C)).
Similar results were observed with other injectable
hydrogel formulations studied earlier with other cel-
lular types [63,64]. The increase in cellular viabil-
ity using injectable hydrogel formulations is a res-
ult of the shear protective effect of injectable hydro-
gel formulations. Next, we studied the effect of the
local inflammatory microenvironment on Schwann
cell viability when encapsulated within our hydrogels
using an in vitro co-culture experiment of Schwann
cells and inflamed astrocytes. Astrocytes are resident
glial cells in the CNS that secrete pro-inflammatory
cytokines and chemokines, mediating SCI inflamma-
tion and scarring. Reactive astrocytes have upregu-
lated GFAP expression [48,65]. We initially optim-
ized the LPS concentration required to activate astro-
cytes (figure S2). In our study, we found that 5 µg
of LPS induced astrocytes to significantly upregu-
late GFAP expression (figure 7(B)), consistent with
previous literature [66,67]. LPS-mediated astrocyte
reactivity is mediated by upregulation of Notch sig-
naling, leading to NFkB activation [68]. As genipin-
crosslinked hydrogel matched the CNS tissue mech-
anical strength, and there was no significant differ-
ence between uncrosslinked and genipin-crosslinked
hydrogels in maintaining cellular viability post injec-
tion, we studied only decell +delipid sciatic and
peroneal genipin-crosslinked hydrogel formulations
to protect Schwann cells against inflamed astrocytes.
In our in vitro inflammatory model, we found that
Schwann cells encapsulated within 1 mM genipin-
crosslinked decell +delipid hydrogels (both sci-
atic and peroneal nerve) had higher cellular viabil-
ity as compared to Schwann cells cultured on TCP
(figures 7(C) and (D)). This study shows that our fab-
ricated hydrogels protect the Schwann cells from the
inflammatory microenvironment to maintain cellu-
lar viability and ultimately increase cellular efficacy
for potential axonal regeneration in SCI treatment.
Additionally, we studied the effect of decell +delipid
hydrogels on rat spinal neuron cells. We found no
significant difference in the live/dead ratio of rat
spinal neuron cells when cultured within peroneal
or sciatic uncrosslinked/1 mM genipin crosslinked
hydrogels (figure S3). Overall, this study suggests
that the developed injectable hydrogel formulations
can support Schwann cell viability post-injection and
additionally will support neuronal cell growth and
proliferation.
6. Conclusion
In this study, we developed a clinically relevant inject-
able hydrogel formulation using decellularized por-
cine peripheral nerves (sciatic and peroneal). An
SDD-based decellularization method was used to effi-
ciently remove cell nuclear debris and retain import-
ant ECM in both sciatic and peroneal decellular-
ized nerves. The presence of lipids in the decel-
lularized porcine nerves was found to hinder the
gelation kinetics of the hydrogels, in contrast to
13
J. Neural Eng. 21 (2024) 046002 G Agarwal et al
prior studies using rat nerve. Thus, a dichloro-
methane and ethanol solvent treatment was used
to efficiently reduce lipid remnants in decellularized
porcine nerves, allowing for the fabrication of mech-
anically stable injectable hydrogels. Delipidation of
both sciatic and peroneal decellularized porcine
nerves improved gelation kinetics of hydrogels, which
would be beneficial for future clinical translation.
Furthermore, genipin-mediated chemical crosslink-
ing of both decell +delipid hydrogels (for both sci-
atic and peroneal) increased the mechanical strength
of the hydrogels, and a 1 mM genipin concentration
of sciatic and peroneal hydrogel achieved a mechan-
ical strength native to CNS tissue, around 2000 Pa.
The decell +delipid hydrogels and 1 mM genipin-
crosslinked hydrogels (sciatic and peroneal) both
demonstrated shear thinning properties and were
able to pass through 25-G needles, which are required
for clinical translation. Schwann cells within uncross-
linked and 1 mM genipin decell +delipid hydro-
gels (sciatic and peroneal), when passed through a
25-G needle, had higher cellular viability as com-
pared to saline controls. In addition, 1 mM genipin-
crosslinked hydrogels protected Schwann cells from
an astrocyte-mediated inflamed microenvironment.
In conclusion, the developed injectable hydrogel for-
mulation holds great potential to efficiently deliver
Schwann cells (or other cells) and maintain higher
cellular viability. In the future, we will study the effi-
ciency of developed injectable hydrogel formulations
to deliver Schwann cells in a contused SCI model in
rodents.
Data availability statement
The data in this manuscript are part of provisional
patent and will be available once patent is granted.
The data that support the findings of this study are
available upon reasonable request from the authors.
Acknowledgments
The authors would like to acknowledge the UF animal
care service (ACS) for tissue sharing of porcine peri-
pheral nerves.
Conflict of interest
The authors declare no conflict of interest.
Funding
The authors would like to acknowledge the funding
support received by NIH (P0225954) and Pruitt chair
funding awarded to PI Christine E Schmidt.
ORCID iDs
Gopal Agarwal https://orcid.org/0000-0002-7906-
9611
Jorge Mojica Santiago https://orcid.org/0000-
0001-6350-1759
Christine E Schmidt https://orcid.org/0000-0002-
5865-6016
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