Bioprinted Anisotropic Scaffolds with Fast Stress Relaxation Bioink for Engineering 3D Skeletal Muscle and Repairing Volumetric Muscle Loss

Abstract

Viscoelastic hydrogels can enhance 3D cell migration and proliferation due to the faster stress relaxation promoting the arrangement of the cellular microenvironment. However, most synthetic photocurable hydrogels used as bioink materials for 3D bioprinting are typically elastic. Developing a photocurable hydrogel bioink with fast stress relaxation would be beneficial for 3D bioprinting engineered 3D skeletal muscles in vitro and repairing volumetric muscle loss (VML) in vivo; however, this remains an ongoing challenge. This study aims to develop an interpenetrating network (IPN) hydrogel with tunable stress relaxation using a combination of gelatin methacryloyl (GelMA) and fibrinogen. These IPN hydrogels with faster stress relaxation showed higher 3D cellular proliferation and better differentiation. A 3D anisotropic biomimetic scaffold was further developed via a printing gel-in-gel strategy, where the extrusion printing of cell-laden viscoelastic FG hydrogel within Carbopol supported gel. The 3D engineered skeletal muscle tissue was further developed via 3D aligned myotube formation and contraction. Furthermore, the cell-free 3D printed scaffold was implanted into a rat VML model, and both the short and long-term repair results demonstrated its ability to enhance functional skeletal muscle tissue regeneration. These data suggest that such viscoelastic hydrogel provided a suitable 3D microenvironment for enhancing 3D myogenic differentiation, and the 3D bioprinted anisotropic structure provided a 3D macroenvironment for myotube organization, which indicated the potential in skeletal muscle engineering and VML regeneration.

Statement of Significance
The development of a viscoelastic 3D aligned biomimetic skeletal muscle scaffold has been focused on skeletal muscle regeneration. However, a credible technique combining viscoelastic hydrogel and printing gel-in-gel strategy for fabricating skeletal muscle tissue was rarely reported. Therefore, in this study, we present an interpenetrating network (IPN) hydrogel with fast stress relaxation for 3D bioprinting engineered skeletal muscle via a printing gel-in-gel strategy. Such IPN hydrogels with tunable fast stress relaxation resulted in high 3D cellular proliferation and adequate differentiation in vitro. Besides, the 3D hydrogel-based scaffolds also enhance functional skeletal muscle regeneration in situ. We believe that this study provides several notable advances in tissue engineering that can be potentially used for skeletal muscle injury treatment in clinical.

Full text

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Acta Biomaterialia xxx (xxxx) xxx
Contents lists available at ScienceDirect
Acta Biomaterialia
journal homepage: www.elsevier.com/locate/actbio
Full length article
Bioprinted anisotropic scaffolds with fast stress relaxation bioink for
engineering 3D skeletal muscle and repairing volumetric muscle loss
✩
Ting Li
a , 1
, Juedong Hou
a , 1
, Ling Wang
c , 1
, Guanjie Zeng
a
, Zihan Wang
a
, Liu Yu
a
, Qiao Yang
c
,
Junfeiyang Yin
a
, Meng Long
a
, Lizhi Chen
a
, Siyuan Chen
d
, Hongwu Zhang
a
, Yanbing Li
a
,
Yaobin Wu
a , ∗, Wenhua Huang
a , b , ∗
a
Guangdong Engineering Research Center for Translation of Medical 3D Printing Application, Guangdong Provincial Key Laboratory of Medical
Biomechanics, Department of Human Anatomy, School of Basic Medical Sciences, Southern Medical University, Guangzhou, China
b
Guangdong Medical Innovation Platform for Translation of 3D Printing Application, Southern Medical University, The Third Affiliated Hospital of Southern
Medical University, Southern Medical University, Guangzhou, China
c
Biomaterials Research Center, School of Biomedical Engineering, Southern Medical University, Guangzhou Guangdong 510515, China
d
The First School of Clinical Medicine, Southern Medical University, Guangzhou, 510515, China
a r t i c l e i n f o
Article history:
Received 5 May 2022
Revised 12 August 2022
Accepted 16 August 2022
Available online xxx
Keywo rds:
Viscoelasticity
IPN hy drogel
3D printing gel-in-gel
3D aligned biomimetic scaffold
Volumetric muscle loss
a b s t r a c t
Viscoelastic hydrogels can enhance 3D cell migration and proliferation due to the faster stress relaxation
promoting the arrangement of the cellular microenvironment. However, most synthetic photocurable hy-
drogels used as bioink materials for 3D bioprinting are typically elastic. Developing a photocurable hy-
drogel bioink with fast stress relaxation would be beneficial for 3D bioprinting engineered 3D skeletal
muscles in vitro and repairing volumetric muscle loss (VML) in vivo ; however, this remains an ongoing
challenge. This study aims to develop an interpenetrating network (IPN) hydrogel with tunable stress re-
laxation using a combination of gelatin methacryloyl (GelMA) and fibrinogen. These IPN hydrogels with
faster stress relaxation showed higher 3D cellular proliferation and better differentiation. A 3D anisotropic
biomimetic scaffold was further developed via a printing gel-in-gel strategy, where the extrusion printing
of cell-laden viscoelastic FG hydrogel within Carbopol supported gel. The 3D engineered skeletal muscle
tissue was further developed via 3D aligned myotube formation and contraction. Furthermore, the cell-
free 3D printed scaffold was implanted into a rat VML model, and both the short and long-term repair
results demonstrated its ability to enhance functional skeletal muscle tissue regeneration. These data sug-
gest that such viscoelastic hydrogel provided a suitable 3D microenvironment for enhancing 3D myogenic
differentiation, and the 3D bioprinted anisotropic structure provided a 3D macroenvironment for myotube
organization, which indicated the potential in skeletal muscle engineering and VML regeneration.
Statement of significance
The development of a viscoelastic 3D aligned biomimetic skeletal muscle scaffold has been focused
on skeletal muscle regeneration. However, a credible technique combining viscoelastic hydrogel and print-
ing gel-in-gel strategy for fabricating skeletal muscle tissue was rarely reported. Therefore, in this study,
we present an interpenetrating network (IPN) hydrogel with fast stress relaxation for 3D bioprinting en-
gineered skeletal muscle via a printing gel-in-gel strategy. Such IPN hydrogels with tunable fast stress
relaxation resulted in high 3D cellular proliferation and adequate differentiation in vitro . Besides, the 3D
hydrogel-based scaffolds also enhance functional skeletal muscle regeneration in situ . We believe that
this study provides several notable advances in tissue engineering that can be potentially used for skele-
tal muscle injury treatment in clinical.
©2022 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
✩ "Part of the Special Issue on Biofabrication for Orthopedic, Maxillofacial, and
Dental Applications, guest-edited by Professors Hala Zreiqat, Khoon Lim, and Debby
Gawlitta."
∗Corresponding authors.
E-mail addresses: wuyaobin2018@smu.edu.cn (Y. Wu), huangwenhua2009
@139.com
(W. Huang) .
1 These authors contributed equally to this work.
https://doi.org/10.1016/j.actbio.2022.08.037
1742-7061/© 2022 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
Please cite this article as: T. Li, J. Hou, L. Wang et al., Bioprinted anisotropic scaffolds with fast stress relaxation bioink for engineering
3D skeletal muscle and repairing volumetric muscle loss, Acta Biomaterialia, https://doi.org/10.1016/j.actbio.2022.08.037

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1. Introduction
Volumetric muscle loss (VML), which may be caused by severe
trauma, tumor excision, or muscle wasting diseases, reduces the
endogenous self-healing ability and causes loss of muscle function-
ality [ 1 , 2 ]. The clinical treatment of VML is limited owing to the
limitations of the donor site and functional deficiency [3–5] . Tis-
sue engineering strategies have recently been regarded as poten-
tial treatment options for VML repair, and various skeletal mus-
cle tissue engineering scaffolds have been developed in the past
decade [6–8] . Furthermore, the fabrication of engineered skeletal
muscle tissues and muscle organ-on-chips is attracting increasing
interest because of their potential as micro-physiological ex vivo
models for investigating skeletal muscle diseases [ 1 , 9-13 ]. Natural
skeletal muscles with a highly oriented structure are formed by
aligned muscle fiber bundle fusion, which plays a key role in force
generation and anisotropic movement [ 14 , 15 ]. Hence, biofabricat-
ing biomimetic anisotropic scaffolds to mimic the aligned structure
of the native skeletal muscle, which can induce cellular elongation
in a three-dimension (3D) environment, is crucial for engineering
skeletal muscle models in vitro and for volumetric muscle loss re-
pair in vivo .
Different advanced approaches, such as micropatterning, elec-
trospinning, and self-assembly, have been applied to fabricate
anisotropic scaffolds to induce 3D cell alignment [ 16 , 17 , 8 , 12 ]. For
instance, our previous studies reported several aligned nanofiber
scaffolds f or developing 3D anisotropic engineered muscle tissues
[ 18 , 19 ], and also cardiac or nerve tissue engineering applications
[ 18 , 20-23 ]. However, these previous approaches are limited in their
practical application for fabricating 3D anisotropic scaffolds, ow-
ing to the complicated or uncontrollable fabrication processes in-
volved. In contrast, 3D bioprinting has received more attention for
the development of scaffolds with 3D complex structures in re-
cent years [ 16 , 24-26 ]. In particular, hydrogel materials are used as
bioinks for 3D bioprinting of engineered tissues [ 4 , 27 ]. However,
the poor mechanical properties of hydrogels are a major limiting
factor for developing bioprinting scaffolds with multi-layer struc-
tures. To overcome this challenge, the printing gel-in-gel strategy,
which prints hydrogel bioinks within a supported gel, has been de-
veloped to fabricate 3D multi-layer scaffolds with complex struc-
tures [28–30] . Recently, 3D bioprinting has been paid more at-
tention to developing engineered skeletal muscle with highly or-
ganized structures [ 12 , 31 ]. Especially, recent advances have led to
improvements in bioprinting strategies to construct large muscle
structures using various hydrogel materials as bioink. However, the
poor mechanical properties of hydrogels are a major limiting fac-
tor for developing bioprinting scaffolds with multi-layer structures.
In contrast, the gel-in-gel strategy could improve the controllabil-
ity and complexity of biomimetic biostructures due to the support
gel. This strategy has been applied to various engineered tissues,
including 3D skeletal muscle tissue [ 17 , 32 ]. However, it remains
challenging to develop a suitable bioink for 3D bioprinting of engi-
neered skeletal muscle tissue using this printing gel-in-gel strategy.
Photo-crosslinking offers improved gelation controllability, and
various photocurable hydrogels, such as gelatin methacryloyl
(GelMA), have been developed as bioinks for 3D bioprinting. How-
ever, most of these photocurable hydrogels are elastic and show
slow stress relaxation properties owing to their covalent crosslink-
ing, as an increasing number of studies have indicated that the
slow stress relaxation of hydrogels limits cell proliferation or even
triggers cellular apoptosis, viscoelastic hydrogels are more desir-
able. Viscoelastic hydrogels with fast stress relaxation properties
can enhance 3D cell spreading and proliferation [33–36] . Natural
hydrogels, such as collagen, alginate, and decellularized extracellu-
lar matrix (dECM), have been widely used as bioinks for 3D bio-
printing of engineered skeletal muscle tissues, not only because of
their good biocompatibility but also due to their viscoelastic prop-
erties [ 11 , 37 , 38 ]. However, the inconvenient gelation processes as-
sociated with these natural hydrogels restrict their practical ap-
plication using the printing gel-in-gel strategy. Therefore, design-
ing a hydrogel that has fast stress relaxation ability and concur-
rently shows a convenient gelation process is an urgent require-
ment. Furthermore, the combination of viscoelastic hydrogel bioink
and printing gel-in-gel strategy for engineering skeletal muscle tis-
sue development has rarely been researched [ 39 , 40 ].
This study aimed to develop an interpenetrating network (IPN)
hydrogel with fast stress relaxation as a bioink for fabricating 3D
bioprinted aligned skeletal muscle biomimetic scaffolds via the
printing gel-in-gel strategy ( Fig. 1 ). This IPN hydrogel was pre-
pared by the combination of GelMA and fibrinogen (FG) follow-
ing the photo-crosslinking and enzyme-crosslinking, which showed
the tunable viscoelasticity by adjusting the concentration ratio of
GelMA and fibrinogen. C2C12 myoblast was encapsulated within
these IPN hydrogels, and the IPN hydrogels with fast stress relax-
ation showed higher 3D cellular proliferation and better differenti-
ation. 3D aligned biomimetic scaffold was further fabricated based
on this IPN hydrogel by using a printing gel-in-gel strategy, and the
3D engineered skeletal muscle tissues in vitro were formed after
cultivation. Moreover, these cell-free 3D printed INP hydrogel scaf-
folds were able to enhance VML model repair in vivo due to their
ability to recruitment of native muscle cells and promote revascu-
larization in situ . Our study would open new avenues for research
on skeletal muscle biomimetic scaffolds and would aid in develop-
ing new treatment strategies for VML.
2. Materials and methods
2.1. Stress relaxation FG hydrogel preparation and characterization
Gelatin methacryloyl (GelMA) was synthesized via the reaction
of gelatin with methacrylic anhydride as following our previous
study [ 19 , 41 ]. The details are available in SI Materials and Methods.
The fibrinogen-GelMA interpenetrating (FG IPN) network hydrogels
were prepared as the following processes. Fibrinogen and GelMA
were both dissolved in DPBS at a different concentration to make
the FG hydrogel solution. The FG hydrogel solution was photopoly-
merized at 365 nm UV light at ∼12mW/cm
2
for 15s with the pres-
ence of Lithium phenyl-2, 4, 6-trimethyl benzoyl phosphinate (LAP)
(0.2 wt%) as the photoinitiator. The photo-crosslinked FG hydrogel
was then immersed into a thrombin-CaCl
2
solution (thrombin: 25
U/ml, CaCl
2
: 10mM) at 37 °C for 30 min to make the fibrinogen
form the enzyme-crosslinking fibrin network. In this study, a list
of FG hydrogels with a different concentration ratio of Fibrinogen
and GelMA were prepared, which were named as GelMA, F2.5G5,
and F5G5, respectively (where the final concentration ratio of fib-
rinogen and GelMA were set as 0 wt%: 5 wt%, 2.5 wt %: 5 wt%,
and 5 wt%: 5 wt%). The FT-IR spectra of GelMA, fibrinogen, and
the freeze-dried FG IPN hydrogel were recorded in the range of
40 0–40 0 0 cm
−1
over 30 scans (NICOLET 6700, Thermo Fisher). The
micro-morphology of the freeze-dried FG hydrogel was visualized
under a scanning electron microscope (SEM) (N70 0 0, Hitachi) af-
ter being coated with gold for 120 s, all the images were mea-
sured and analyzed by Image J software. The rheological analysis
of hydrogel samples was conducted with the advanced rheometer
(HAAKE MARS 40, Thermo Fisher) using a 20 mm-diameter paral-
lel plate. For stress relaxation tests, strains were applied with 1%,
5%, and 10%, except for experiments with a constant strain rate
for 0-400s. The viscosity of the FG physical-crosslinking hydrogel
was tested via oscillatory frequency sweeps at 15 °C, and oscilla-
tion was applied in 0.01 s
−1 to 100 s
−1 with 1 % strain. The sweep
tests of the FG interpenetrating network hydrogel were evaluated
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Fig. 1. Schematic of design and fabrication of a stress-relaxed FG IPN hydrogel-based 3D bioprinted aligned biomimetic scaffold that mimics the native structure of skeletal
muscle tissues. (a) Design of the viscoelastic FG IPN hydrogel for 3D bioprinted aligned biomimetic scaffolds with oriented structure prepared via the printing gel-in-gel
strategy. (b) Scheme showing the use of stress-relaxed FG IPN hydrogel-based 3D bioprinted aligned skeletal muscle scaffold for inducing 3D aligned myoblast growth in
vitro and repairing skeletal muscle injuries in vivo .
with frequencies ranging from 0.1 to 10 rad s
−1 at a rotation am-
plitude of 1 % at 25 °C.
2.2. C2C12 cells culture within FG hydrogel and cell characterization
C2C12 myoblasts were cultured in the FG hydrogels as the fol-
lowing procedure. First, C2C12 myoblasts were digested by trypsin
and resuspended (2 ×10
5 cells/mL) in the FG hydrogel solution.
The cell-laden FG hydrogel solution was added into a PDMS model
with a certain shape, and then photo-crosslinked by UV light (365
nm) for 15 s. Subsequently, the photo-crosslinked FG hydrogel was
immersed in the thrombin-CaCl2 solution (thrombin: 25 U/mL,
CaCl2: 10 mM) for 30 min at 37 °C. Finally, these cell-laden hy-
drogels were transferred from the PDMS model into the culturing
petri dish, and the growth medium (Contained Dulbecco’s mod-
ified eagle’s medium-high glucose supplemented with 10 % fetal
bovine serum and 1 % P/S). The growth medium was changed ev-
ery 2 days. Besides, after 7 days of cultivation, the differentiation
medium (Contained Dulbecco’s modified eagle’s medium-high glu-
cose supplemented with 2 % horse serum and 1 % P/S) was added
for 7 days, and it was changed every 2 days.
The live/dead staining was performed as described in the fol-
lowing steps. In brief, cells were washed with DPBS three times for
10 min to remove the cell culture medium. Then, cells were treated
with ethidium homodimer-1 (0.5 μM) and Calcein-AM (0.25 μM)
for 30 min at 37 °C. The results of green color indicate the liv-
ing cells, while the red color indicates the dead cells. The F-actin
staining of cells within the FG hydrogel was applied as the fol-
lowing process. After the cells-laden FG hydrogels were cultured in
a growth medium for 7 days, the FG hydrogels were rinsed three
times gently with DPBS and then fixed with 4 % paraformaldehyde
for 25 min at room temperature. The fixed cells were rinsed three
times gently with DPBS and then treated with 0.2 % Triton X-100
for 30 min. After being blocked in 1% BSA in DPBS for 1.5 h, the
cells were incubated with phalloidin conjugated with FITC at room
temperature for 1.5 h. Cell nuclei were counterstained with DAPI
before observation under a confocal laser microscope (LSM 880,
Carl Zeiss). The mysin heavy chain (MHC) staining was performed
as being similar to the above steps. The cell-laden FG hydrogels
were rinsed using DPBS and then fixed with 4 % paraformaldehyde.
Then they were treated with 0.2 % Triton X-100. And after being
blocked in 1 % BSA in DPBS for 1.5 h, they were incubated in the
mouse anti-MHC (myosin heavy chain) antibody at 4 °C overnight.
After being washed with DPBS, Alexa Fluor 488 conjugated sec-
ondary antibody was added and incubated for 1. 5 h at 24 °C. All
fluorescence staining images were observed under a confocal laser
microscope (LSM 880, Carl Zeiss) and analyzed using the ImageJ
software.
2.3. Carbopol Supported gel preparation and characterization
The Carbopol gel was prepared as the detail in SI Materials and
Methods. Then, three different concentrations of the Carbopol sup-
ported gels were set as 0.2 wt%, 0.8 wt%, and 1.4 wt%, respectively.
Then the Carbopol supported gels were added into the transpar-
ent boxes to explore their transparence properties. The rheologi-
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cal properties of the Carbopol supported gels were tested by an
AR20 0 0 rheometer (HAAKE MARS 40, Thermo Fisher) with a 20
mm diameter parallel plate at 4 °C. The viscosity was measured by
varying the shear rate from 0.01 s
−1 to 10 s
−1 by using a rotary
testing setup. The FG hydrogel solution was transferred into the
printing syringe with a 22 G nozzle at 15 °C and then was printed
(Pressure: 2 bar, feed rate: 5 mm/s) in the 0.2 wt%, 0.8 wt%, and
1.4 wt% Carbopol supported gel to explore the swelling degree.
2.4. 3D printing hydrogel-based scaffolds in the Carbopol supported
gel
3D printing hydrogel-based scaffolds in the Carbopol supported
gel were applied as following process. The FG hydrogel solution
was added to the printing syringe with a 21 G nozzle and main-
tained at 15 °C to form the physical-crosslinking hydrogel before
3D printing. The 0.8 wt% Carbopol supported gel bath was pre-
cooled for 4 °C and put in a transparent dish, and the printing noz-
zle was inserted in the Carbopol supported gel for printing scaf-
folds. All the 3D structures were designed by Autodesk Fusion 360
software and then printed using a 3D bioprinter (Envision, Ger-
many). A list of FG hydrogel-aligned filament structures with dif-
ferent diameters was fabricated by setting the printing pressures
as 1, 1.5 , 2, 2.5, and 3 bar, respectively, and the feed rates were set
as 3, 5, 7, 9, and 11 mm/s, respectively. The diameters of printed
filaments in the Carbopol supported gel were observed by an opti-
cal microscope (Olympus) and measured by Image J software. Be-
sides, the stability was explored by printing some structures such
as "SMU", hollow hexagonal prisms, and trigeminal vessels in the
0.8 wt% Carbopol supported gel. The printing parameters were set
as 1.5 bar and 5 mm/s. Besides, a 3D bioprinted aligned biomimetic
scaffold was printed as the described process, and it was taken
out after being cross-linked by UV light (365 nm) for 15 s and
thrombin-CaCl
2
(thrombin: 25 U/ml, CaCl
2
: 10 mM) at 37 °C for 30
min, and the Carbopol supported gel was cleaned by rinsed with
200 mM NaCl solution.
2.5. 3D bioprinting cell-laden aligned biomimetic scaffolds
The 3D printed aligned biomimetic skeletal muscle scaffold
was designed as the size of 15 mm ×6 mm ×4 mm
(length ×width ×height) ( Fig. 6 a). C2C12 cells were encapsu-
lated and cultured in 3D bioprinted aligned biomimetic scaffolds
as described in the following process. C2C12 cells were digested
and encapsulated in the F5G5 bioink at a concentration of 2 ×10
5
cells/mL. The cell-laden FG bioink was gently mixed to avoid cre-
ating bubbles and was then transferred to a printing syringe with
a commercially 21 G nozzle. Meanwhile, the printing syringe was
maintained at 15 °C stably and made the FG bioink as a physical-
crosslinking gel for printing. 3D printing was performed by using a
3D bioprinter (Envision, Germany) with a syringe pump extruder,
and the Carbopol supported gel was loaded into holders. The print-
ing speed was set as 5 mm s
−1
, and the printing pressure was set
as 1.5 bar. Once the scaffold was printed within the Carbopol sup-
ported gel, it was cured by UV light (365 nm) for 15 s to form
the photo-crosslinked hydrogel scaffold. Subsequently, NaCl solu-
tion (200 Mm) was added to dissolve the Carbopol supported gel,
and the scaffold was then cross-linked by thrombin and CaCl
2
so-
lution (thrombin: 25 U/ml, CaCl
2
: 10 mM) for 30 min at 37 °C. Fi-
nally, the growth medium (Contained Dulbecco’s modified Eagle’s
medium-high glucose supplemented with 10 % fetal bovine serum
and 1 % P/S) was added for cell cultivation. All these processes
were conducted in sterilized conditions.
The cell-laden 3D bioprinted aligned biomimetic scaffold was
cultured in the growth medium for 7 days at 37 °C in a humidi-
fied 5 % CO
2
atmosphere. The live/dead staining was applied af-
ter cell culturing within 3D aligned scaffolds f or 1, 3, and 5 days,
respectively. The F-actin and the MHC staining were applied after
culturing for 5 days and differentiation for 7 days. The 3D aligned
scaffold was washed with DPBS three times to remove the cell cul-
ture medium and treated with Ethidium homodimer-1 (0.5 μM)
and calcein-AM (0.25 μM) for 30 min. The staining scaffold was ob-
served under a confocal laser scanning microscope (LSM 880, Carl
Zeiss), and the green signal resulted from staining with calcein-
AM, indicating living, while the red signal resulted from staining
with Ethidium homodimer-1 indicating dead cell. The number of
live and dead cells was analyzed by calculating the ratio of viable
cells to the total cell number using the ImageJ software from at
least three different microscopic fields (n ≥3 in each experiment).
In addition, cells cultivation within the FG hydrogel was regarded
as the control group. For the F-actin staining of C2C12 cells within
the 3D aligned scaffolds, the 3D cell-laden aligned scaffolds were
fixed in 4 % paraformaldehyde for 15 min, washed with DPBS three
times, and then permeabilized with 0.2 % Triton-X 10 0 in DPBS
for 40 min at room temperature. The 3D aligned scaffolds were
blocked with 1 % bovine serum albumin for 1.5 h to reduce non-
specific binding. Then, the 3D aligned scaffolds were treated with
phalloidin-rhodamine (1: 800 dilution) for 90 min at 24 °C and
counterstained with DAPI. The staining images were observed by
a confocal laser scanning microscope (LSM 880, Carl Zeiss) and a
stereomicroscope (Carl Zeiss). The orientation of C2C12 cells and
the area coverage of the cytoskeleton in the cell-laden 3D aligned
scaffold were measured and analyzed by Image J software.
The primary skeletal muscle cells were isolated as described in
a previous study [42] and encapsulated in the 3D aligned scaf-
folds culturing for 9 days. The growth medium (Contained 1: 1
DMEM and Ham’s F-12 medium supplemented with 20 % fetal
bovine serum, 2.5 ng/ml bFGF, and 1 % P/S) was added for cell
cultivation and changed every 2 days. The MHC staining was per-
formed for observing the myotubes. After 9 days of culture, the
cell-laden 3D scaffolds were stained with 5 μm Fluo-4 AM to vi-
sualize the calcium transients according to the manufacturer’s in-
structions. Briefly, after removing the culture medium, the scaf-
folds were washed two times with Tyrode’s solution and treated
with 5 μM Fluo-4 AM (prepared in Tyrode solution) for 30 min
at 37 °C in 5% CO
2
. After staining, the scaffolds were washed with
Tyrode’s solution, and calcium transients were recorded under a
fluorescence microscope (Olympus, BX53). Data analysis was per-
formed using ImageJ software. The background fluorescence was
subtracted from the mean fluorescence in the ROI (Region of Inter-
est), giving a background-corrected normalized fluorescence value.
Changes in the [Ca
2 +
] are expressed as F/F0, where F is the fluores-
cence intensity at intermediate calcium levels, and F0 is the base-
line fluorescence intensity. The MHC and F-actin staining process
of the primary skeletal muscle cells cultured in 3D aligned scaf-
folds was similar to the steps as mentioned previously. The images
were measured and analyzed by Image J software.
2.6. Inflammation and microvascularization of 3D bioprinted aligned
biomimetic scaffolds in subcutaneous
The animal experiments were approved by the Institutional An-
imal Ethics Committee of the Southern Medical University, and
all institutional and national guidelines for the laboratory animals
were followed. The subcutaneous implantation in rats was pre-
pared to determine the angiogenesis of the 3D bioprinted aligned
biomimetic scaffolds in vivo . Briefly, Sprague Dawley (SD) rats (Fe-
male, 180 g - 200 g, n ≥3) were anesthetized with isoflurane
at the concentration of 3-4 % for induced euthanasia The 3D gels
were set as the control group. The 3D gel group and the 3D aligned
scaffold group were implanted in the notches, respectively. The
skin incision was then sutured by an absorbable surgical suture
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(Ethicon). After implanting for 14 days, the rats were sacrificed
and the implanted samples were carefully dissected for histological
staining to evaluate the vascularization of the 3D gel group and the
3D aligned scaffold group in rats. Tissues were embedded in paraf-
fin, sectioned (3 μm thick), and stained by hematoxylin and eosin
(H&E). The H&E and immunofluorescence staining results were ob-
served by an optical microscope and analyzed by the Image J soft-
ware.
2.7. The functional recovery tests after the regeneration following
VML
Animal experiments were performed according to the guide for
the care and use of laboratory animals, approved by the Institu-
tional Animal Ethics Committee of the Southern Medical Univer-
sity. The 3D gel and the 3D scaffold were implanted into the defect
(about 15 mm ×10 mm ×5 mm) of the rat tibialis anterior mus-
cle, the blank group which just surgery but no implant was set as
the control group. After being implanted for 2 weeks, rats were eu-
thanized with isoflurane at the concentration of 3-4%. The regen-
erated defects and surrounding tissues were carefully dissected for
histological staining. Tissues were embedded in paraffin, sectioned
(3 μm thick), and stained by H&E. On the other hand, after de-
paraffinization using xylene and ethanol, the samples were perme-
abilized with 0.5% Triton X-100 and blocked with 5% goat serum.
Samples were subsequently stained with the CD31 and MHC pri-
mary antibodies (red: myosin, green: CD31, blue: DAPI). The my-
ofiber diameter, the number of centronucleated myofibers, and the
capillary density in skeletal muscle defect were evaluated by Im-
ageJ software.
The rat treadmill experiments were taken after implantation for
6 months through 1 week of acclimatization. Rats were running on
a track (animal experiment treadmill, ZH-PT/5S) within increased
5m/min speed per 5 mins from 5 m/min to 35m/min until to frag-
ile moving. Each group contain at least 3 rats and each experi-
ment would rest for 30min. Ultrasound images of the tibialis an-
terior muscle were taken after implantation for 6 months in both
side-section and cross-section planes. After anesthesia, the ultra-
sonic gel was used to identify the muscle that was removed hairs
with a shaver. Images were recorded under the same acquisition
settings bilaterally in B-mode using the Vevo2100 imaging system
(FUJIFILM VisualSonics) equipped with a 30 MHz linear transducer
(MS400). Muscle contractility was tested on isolated strips from
animals at 6 months post-surgery. The sciatic nerve in the thigh
was exposed and two electrodes were attached to the nerve. Af-
ter anesthetized, the animals were grounded to reduce electrical
noise from the recordings. The tibialis anterior muscle was half
resected and attached at one end to a force transducer that was
integrated with BL-420N biological signal acquisition and analy-
sis system (TECHMAN, China). Krebs solution (contained Ca
2 + and
Mg
2 +
) was pipetted on the muscle strips all time. The maximum
isometric tetanic contractile force was determined at different fre-
quencies (20-180 Hz) with a pulse of 1ms delay, and contractile re-
sponses were expressed as isometric tension/cross-section area. All
the groups contained three strips and each strip was tested three
times.
2.8. Statistical analysis
Experiments were run at least triplicate for each sample (n ≥3
per sample). The results are presented as mean ±standard devia-
tion, and the error bars in all figures represent the standard devia-
tion from the mean. Quantitative data of the cell aspect ratio were
obtained using the ImageJ software. Two groups were compared
using the t-test analysis, and three or more groups were analyzed
using a one-way analysis of variance (ANOVA). A significance level
of 0.05 was applied to determine significant differences.
3. Results and discussion
3.1. Preparation and characterization of stress-relaxed FG IPN
hydrogel
As shown in the schematic for fibrinogen/GelMA interpenetrat-
ing network (FG IPN) hydrogel preparation ( Fig. 2 and Fig. S1a),
it was gelled after photo-crosslinking under UV irradiation and
enzyme-crosslinking of fibrinogen in thrombin-CaCl
2
solution. Sev-
eral FG IPN hydrogels were prepared by adjusting the concentra-
tion ratio of fibrinogen to GelMA as 0 wt%:5 wt%, 2.5 wt%:5 wt%,
and 5 wt%:5 wt% and were named as GelMA, F2.5G5, and F5G5,
respectively ( Fig. 2 b and SI Table 1). The transparency of FG INP
hydrogels decreased when the concentration of fibrinogen was in-
creased ( Fig. 2 c). This was mainly due to the increased number of
fibrin networks formed by cross-linking of thrombin and calcium
chloride to produce protein nanofibrils. Furthermore, the SEM im-
ages showed the microporous structure of FG hydrogels covered
with fibrin nanofibers, indicating that the fibrin formed a nanofiber
structure to form the second crosslinking network. However, the
GelMA hydrogel did not exhibit any nanofibers on its microporous
surface ( Fig. 2 d, Fig. S3a). We further analyzed the porosity rate
and the mesh density of the FG hydrogel systems according to the
SEM images. As shown in Fig. S4a, these three groups exhibited
no significant difference in the porosity rate. In contrast, the mesh
density of the F5G5 group (43 ±2%) and the F2.5G5 group (27
±2%) was significantly higher than that of the GelMA group (11
±2%) (Fig. S4b). Therefore, these results revealed that the fibrin
fiber networks distributed in the F5G5 and F2.5G5 hydrogel could
supply more fibrilla structures compared with the GelMA hydro-
gel. The combined properties of the FG IPN hydrogels were fur-
ther investigated using FT-IR spectroscopy. A decrease in the peak
around 3400 cm
−1 corresponding to OH groups and a decrease in
the peak around 1640 cm
−1
related to the C = C bond was observed
in the spectrum of the FG IPN hydrogel compared with those in the
spectrum of the GelMA polymer, which confirmed the photopoly-
merization of GelMA in the FG IPN hydrogel ( Fig. 2 e). However,
compared with that in the FT-IR spectrum of fibrinogen, the peak
ranged from 1600 cm
−1
to 170 0 cm
−1
, corresponding to the amide
group ( Fig. 2 e), indicating that enzyme-crosslinking resulted in de-
protonation and formation of new hydrogen bonds within the FG
IPN hydrogel [43] . These results indicate a good blending of GelMA
and fibrinogen in the FG IPN hydrogel.
Recent studies have demonstrated that stress-relaxation prop-
erties play a key role in inducing 3D cell spreading and prolifer-
ation. Therefore, rheological tests were performed to investigate
the stress-relaxation behaviors of the FG IPN hydrogels. The rate
of stress relaxation under constant strain (1%, 5%, and 10% ) was
quantified as the time for the initially measured stress to relax
to half of its original value. Fibrinogen formed a weak and vis-
coelastic fibrin hydrogel after enzyme-crosslinking, and the fibrin
gel exhibited fast stress relaxation ( Fig. 2 f and g), which was con-
sistent with the results of a previous study [44] . Accordingly, in
our study, the inclusion of fibrin was also expected to accelerate
the stress relaxation of these FG IPN hydrogels owing to the fast
stress relaxation of fibrin. After normalization to the initial stress,
the FG IPN hydrogels exhibited stress-relaxation behaviors, com-
pared with the elastic GelMA hydrogel, under constant 5% and 10%
strains, and the increased concentration of fibrinogen resulted in
a higher stress relaxation rate ( Fig. 2 f). The mechanical properties
of FG IPN hydrogels were examined using a frequency sweep test.
All these hydrogel samples showed a plateau storage modulus (G’)
and loss modulus (G’’), as well as G’/G’’ > 100, in the range of 0.1–
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Fig. 2. Stress relaxation is associated with the fibrin fibers connected to the networks in FG hydrogel. (a) Illustration of the mechanism responsible for viscosity, elasticity,
and viscoelasticity in the stress-relaxed FG IPN hydrogels. (b) Schematic showing that the FG IPN hydrogel methodology for tuning the FG stress-relaxation rate is dependent
on the concentration of fibrinogen. The gross image (c) and SEM image (d) of the FG hydrogel. Scale bar: 50
μm. (e) FT-IR spectra of GelMA, fibrinogen, and FG hydrogel s
after lyophilization. (f) Strains were normalized by the strain at the original storage modulus for stress relaxation at different strains (1%, 5%, and 10%) for GelMA, F2.5G5,
and F5G5 hydrogels. (g) Initial elastic modulus from unconfined compression tests. (h) The initial storage modulus (G’) and loss modulus (G’’) of GelMA, F2.5G5, F5G5, and
fibrinogen hydroge l in the frequency ranging from 0.1 to 10 Hz.
10 Hz, which suggested the formation of stable hydrogel networks.
Moreover, the G’ of F5G5 hydrogels was slightly increased com-
pared with that of F2.5G5 and GelMA samples, while there was
no significant difference in storage modulus between these groups,
which also suggested a synergistic effect on the mechanical prop-
erties of FG IPN hydrogels owing to the interaction between the
individual GelMA and fibrin networks ( Fig. 2 h). These rheological
results illustrate that the FG IPN hydrogel successfully formed a
stress-relaxed IPN hydrogel after photo-crosslinking and enzyme-
crosslinking, which could provide stress relaxation and suitable
mechanical properties for tissue engineering scaffolds.
3.2. 3D cellular spreading, proliferation, and differentiation within
the FG IPN hydrogel
To investigate the effects of viscoelastic fibrin in FG IPN gels on
the cellular response, F2.5G5 and F5G5 IPN hydrogels with differ-
ent fibrin concentrations were used in further studies, and stati-
cally crosslinked GelMA (without fibrin) was used as the control
group. The viability of C2C12 cells encapsulated within FG IPN hy-
drogels was analyzed using live/dead staining ( Fig. 3 a). The stain-
ing images showed that most of the cells within these hydrogel
samples were stained green after one, three, and five days of cul-
tivation, and cells exhibited significant growth with increasing cul-
ture time ( Fig. 3 a, Fig. S5). Approximately 90% of the cells re-
mained alive during five days of cultivation within these hydrogels,
and no significant difference in viability was observed when the
cells were cultured in GelMA, F2.5G5, and F5G5 hydrogels ( Fig. 3 b).
These data suggested that the stress-relaxed FG IPN hydrogels
showed good biocompatibility within the viscoelastic 3D environ-
ment. The effect of the viscoelastic hydrogel on cellular morphol-
ogy was further investigated using F-actin staining after culturing
for seven days, which showed that most of the cells within the
F2.5G5 and F5G5 hydrogels exhibited significant elongation com-
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Fig. 3. Morphologies and differentiation of C2C12 cells within the stress-relaxed FG IPN hydrogels. (a) Fluorescence images of C2C12 cells in the GelMA, F2.5G5, and F5G5
hydrogels were obtained with live/dead staining after five days of cultivation. Scale bar: 200
μm. (b) Cell viability percentages of C2C12 cells in the FG hydrogel after
culturing for one, three, and five days. (c) Fluorescence images and 3D views of C2C12 cells in the GelMA, F2.5G5, and F5G5 hydrogels after staining with F-actin and DAPI
after seven days of cultivation. Scale bar: 100
μm. Statistical analysis of the cellular aspect ratio (d) and the relative percentage of area coverage (e) of the C2C12 cells by
F-actin in the GelMA, F2.5G5, and F5G5 hydrogels . (f) The views of immunofluorescence images of myotub es in the GelMA, F2.5G5, and F5G5 hydrogels were obtained after
staining with MHC (green) and DAPI (blue) after culturing for seven days in a differentiation medium. Scale bar: 100 μm and 200 μm. Statistical analysis of myotube width
(g) and myotube length (h) of C2C12 cells in the GelMA, F2.5G5, and F5G5 hydrogels after culturing for seven days in the differentiation medium. (i) Illustration of the effect
of elastic GelMA and viscoelasticity FG gel on cell spreading in three dimensions.
∗p < 0 .05 and
∗∗p < 0.01, compared with the blank group.
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pared with those in the GelMA hydrogel group ( Fig. 3 c). Cellu-
lar elongation was also measured using the cellular aspect ratio,
which is defined as the ratio between the length of the longest line
and the shortest line across the nucleus. Statistical analysis showed
that the cellular aspect ratio of cells in the F5G5 group (35.2 ±9.1)
was significantly higher than that of cells in the F2.5G5 (12.7 ±
1.5 ) and GelMA (5 ±1.5) groups ( Fig. 3 d). Moreover, 3D view im-
ages showed that the cytoskeleton filled the FG IPN hydrogel after
seven days of cultivation ( Fig. 3 c); additionally, it showed that the
area coverage of cytoskeleton in the F5G5 group (57 ±7.5%) was
much more than that in GelMA (26.4 ±12.5%) and F2.5G5 (48.3 ±
6.3%; Fig. 3 e). These results demonstrated that the stress relaxation
rate plays a key role in regulating cellular spreading and elongation
in 3D viscoelastic hydrogels.
To further investigate the effect of stress relaxation of the
hydrogel on myoblast differentiation, myosin heavy chain (MHC)
staining was performed after cultivation in a differentiation
medium for seven days. The immunofluorescence staining images
indicated that the rate of myotube formation in C2C12 cells within
the F5G5 hydrogel was significantly higher than that in the other
groups; the 3D views further indicated that the formed myotubes
filled the stress-relaxed FG IPN hydrogel ( Fig. 3 f). In particular,
more myotubes were formed in the FG IPN hydrogel with a faster
stress relaxation ( Fig. 3 f). Furthermore, myotube width analysis
showed that myotubes formed within the F5G5 group (37 ±9 μm)
were significantly larger than those formed within the F2.5G5 (22.
±4 μm) and GelMA (16 ±3 μm) groups ( Fig. 3 g). In addition,
the myotubes in the F5G5 hydrogel (427 ±88 μm) were much
longer than those in the F2.5G5 (133 ±14 μm) and GelMA hy-
drogels (82 ±11 μm) ( Fig. 3 h). Furthermore, the orientation dis-
tribution and aspect ratio of the myotubes and nuclei were both
enhanced with the increase of fibrinogen contents in the FG hy-
drogel system, which might be due to the positive effect of the
viscoelasticity and fibrilla structures on myotubes formation and
maturation (Fig. S6). Besides, the maturation index is another dif-
ferentiation parameter to evaluate the maturation of myotubes by
measuring the numbers of myotubes containing more than 3 nu-
clei. The results showed that more mature myotubes were formed
in the F5G5 and F2.5G5 groups compared with the GelMA group
(Fig. S6). However, some limitations are still needed to overcome
in this present study. Especially, the measurement of myogenic
genes was lacking due to the difficulty of collecting DNA from the
cells encapsulated within such FG hydrogel. Therefore, we will fur-
ther investigate the suitable method to analyze the gene expres-
sion level of cells culturing within this FG hydrogel. Previous stud-
ies have demonstrated that the viscoelastic properties of hydrogels
have an important effect on cellular responses in 3D environments,
wherein viscoelastic hydrogels facilitate the release of energy to re-
duce stored energy through stress relaxation, which makes cells
spread more easily than elastic hydrogels [ 31 , 42 ]. In particular,
some researchers found that an IPN hydrogel with fast stress re-
laxation, which was prepared via the integration of collagen fibers
and a hyaluronic acid hydrogel, could enhance cell spreading be-
cause of the local increase in integrin density and collagen fiber
realignment through cellular focal adhesions [45] . Such cellular
focal adhesions are generally observed only in fibrillar environ-
ments, such as collagen, fibrin, and cell-derived matrices [46] . In
our study, based on the above-mentioned observations, we found
that FG hydrogels with faster stress relaxation showed the ability
to enhance cell spread, proliferation, elongation, and myogenic dif-
ferentiation in a 3D environment. These results suggest that cells
apply force through integrins on fibrin nanofibers, and fibrin re-
alignment enhances cell adhesion and permits cellular elongation
along the fibrin nanofibers ( Fig. 3 i). Therefore, we speculated that
such FG IPN hydrogels integrating both viscoelasticity and fibrillar
architecture may provide a suitable microenvironment for guiding
cellular responses in a 3D environment. In particular, the F5G5 hy-
drogel with higher fibrinogen concentration was more suitable for
enhancing cell growth and myogenic differentiation. To investigate
the potential of the stress-relaxed FG IPN hydrogel as a bioink for
3D bioprinting, the F5G5 hydrogel samples were selected for all
subsequent experiments (hereafter termed the FG IPN hydrogel).
3.3. 3D printing of stress-relaxed FG IPN hydrogel in the
Carbopol-supported gel
We further used the FG IPN hydrogel as a bioink for fabricat-
ing 3D bioprinted biomimetic scaffolds via a printing gel-in-gel
strategy ( Fig. 4 a). And the physical properties of the FG IPN gel
were investigated to confirm the feasibility of the printing process.
Rheological tests showed that the FG solution transformed into a
physically crosslinked gel below 15 °C ( Fig. 4 b and c), mainly due
to the formation of a hydrogen-bonding network in the gelatin
chain at low temperatures [47] . The viscosity of the precooled FG
gel significantly decreased with increasing shear rate. Such shear
thinning made the FG IPN bioink squeeze through the printing
needle easily ( Fig. 4 d). To support the weak gel scaffold during
printing, Carbopol gel was chosen as the supported gel because
of its good biocompatibility, thixotropy, and transparent properties
[48] . For preparing the optimum Carbopol supported gel, we ex-
plored the Carbopol supported gels with three concentrations, in-
cluding 0.2 wt%, 0.8 wt%, and 1.4 wt%, respectively. Fig. 4 e shows
that the transparency of the Carbopol gels decreased with in-
creasing gel concentration. Furthermore, the viscosity of these Car-
bopol gels decreased significantly with increasing shear rate, in-
dicating their shear-thinning and thixotropy properties. Therefore,
the printing nozzle could easily move without resistance in the
Carbopol-supported gels ( Fig. 4 f). Furthermore, to investigate the
ability of Carbopol gels to maintain the stability of the printed hy-
drogel structure, the swelling degrees of the printed FG physically
crosslinked hydrogel (before photo-crosslinking) in the Carbopol-
supported gels were analyzed after printing for 15 min and 9 h. As
shown in Fig. S7, the FG IPN hydrogel exhibited a high swelling de-
gree in the 0.2 wt% Carbopol-supported gel; however, there were
no obvious swelling changes in the 0.8 wt% and 1. 4 wt% gels. Based
on these results, we chose the 0.8 wt% Carbopol-supported gel,
which could maintain the stability of printing scaffolds and simul-
taneously showed good transparency for observation, as the can-
didate Carbopol-supported gel. Furthermore, to analyze the effect
of printing parameters on the line width of hydrogel fibers in the
Carbopol-supported gel, different moving speeds (3, 5, 7, 9, and
11 mm/s) and printing pressures (1.0, 1.5 , 2.0, 2.5, and 3.0 bar)
were applied ( Fig. 4 g. h). When the moving speed was increased
from 3 to 11 mm/s at a pressure of 2 bar, the line width of the
printed fibers decreased from 543 ±89 μm to 313 ±89 μm.
In addition, when the pressure was increased from 1.5 bar to 3.0
bar with 5 mm/s, the line width of the printed fibers increased
from 266 ±58 μm to 597 ±111 μm ( Fig. 4 i). These results il-
lustrated that the line width of the printed fibers increased with
increasing printing pressure, whereas it decreased with increasing
moving speed in the Carbopol-supported gel. Furthermore, quan-
titative analysis showed that the line width of the printed fibers
ranged from 200–750 μm, demonstrating that multiscale hydro-
gel fibers were easily printed in the Carbopol-supported gel by
tuning the printing pressure and moving speed. The printing pa-
rameters of 5 mm/s moving speed and 1. 5 bar printing pressure
were chosen for further investigation because the fibers printed
using these parameters were more stable and suitable for cre-
ating topographical cues to induce cell growth. To confirm the
printing feasibility of the Carbopol-supported gel, a 3D structure
“SMU” was printed in the Carbopol-supported gel at 15 °C using
the stress relaxation FG hydrogel as the bioink, and the printed
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Fig. 4. 3D Printing of the stress-relaxed FG IPN hydrogel in the Carbopol-supported gel. (a) Scheme of the FG solution showing sol-gel transition at different tem peratures.
(b) The FG physical crosslinked gel was formed at 15 °C; however, it turned into a solution when the temperature increased to 37
°C. (c) The viscosity of the FG phys-
ical crosslinked gel at 15 °C. (d) Viscosities of the FG gel with the shear rate ranged from 0.1 to 100 1/s. (e) Gross images of the Carbopol-supported gel with different
concentrations (0.2 wt%, 0.8 wt%, and 1.4 w t% ) . (f) Viscosities of the Carbopol-supported gel with the shear rate ranging from 0.1 to 10 1/s. 3D printed FG hydrogel fibers
in the Carbopol-supported gel (g) under different printing parameters (h). (i) Statistical diagram of printed fiber diameters in the Carbopol-supported gel under different
parameters; the X-axis is the printing pressure, the Y-a xi s is
the printing speed, and the z-axis is the line width of printed fibers. (j) the FG gel-based “SMU” was printed in
the Carbopol-supported gel at 15 °C. (k–m) Schematic and procedure of 3D aligned skeletal muscle scaffold printing in the Carbopol-supported gel and without the support
gel. Scale bar: 50 0
μm.
structures were well maintained without any collapse (
Fig. 4 j and
movie S1). Furthermore, other structures, such as a hollow hexag-
onal prism scaffold and trigeminal vessels were also designed and
printed in the Carbopol-supported gel (Fig. S8a and b, movies S2
and S3). The printing gel-in-gel strategy has received increasing
attention for the development of hydrogel-based scaffolds with
complex 3D structures, while most previous studies used natural
hydrogels such as collagen or dECM as bioinks [ 17 , 32 , 49 ]. How-
ever, the inconvenient gelation process limits their practical use
in 3D printed scaffold development. In contrast, photo-crosslinking
provides an easily controllable gelation ability, while most of the
present photocurable hydrogels lack suitable viscoelasticity for 3D
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Fig. 5. FG IPN Hydrogel-based 3D bioprinted C2C12 cells-laden aligned scaffold as in vitro engineered skeletal muscle model. (a) Schematic of the formation of the stress-
relaxed FG IPN hydrogel-based 3D bioprinted aligned scaffold by printed gel-in-gel strategy, and the cellular elongation in the 3D aligned scaffold. (b) The live/dead staining
images of C2C12 cells cultured within the 3D aligned scaffolds. Quantit ative analysis of cell viability (c) and the cell aspect ratio (d) in the 3D aligned scaffolds and 3D gels
at days 1, 3, and 5. (e) The view and fluorescence images of C2C12 cells within the 3D aligned
scaffold by staining with the F-actin and DAPI after seven days of cultivation;
the right image is the inset of a yellow coil. Quantified analysis of the relative percentage of area coverage (f) and the cells aligned within ±10 °(g) in the 3D gel and 3D
aligned scaffold group, respectively. (h) Fluorescence staining image and the quantitative analysis of cellular orientation distribution of C2C12 cells within the 3D aligned
scaffold group. (i) Fluorescence staining image and the quantitative analysis of the orientation distribution of C2C12 cells within the 3D gel group. (j) Immunofluorescence
image of muscle satellite cell differentiation in
the 3D aligned scaffold. (k) The calcium transient immunofluorescence image of the muscle satellite cells in the 3D aligned
scaffolds. (l) Line scanning across each ROI captured spontaneous calcium transients, scale bar: 30 s. The dynamic changes of Ca
2 +
(F/F0) showed according to the normalized
fluorescence intensity of each ROI.
cellular spreading and proliferation [50–52] . Comparatively, in our
study, the FG IPN hydrogel with fast stress relaxation, as the bioink
for 3D bioprinting skeletal muscle scaffold, would not only per-
form in the Carbopol-supported gel and could maintain a highly
accurate shape and structure, but also exhibited the effect of vis-
coelastic properties on 3D cellular behavior control. Therefore, the
3D printed biomimetic aligned skeletal muscle scaffold was fabri-
cated with a size of 15 ×6 ×4 mm (length ×width ×height)
( Fig. 4 k, l), and the diameter of the printed fibers was set to 300
μm. In addition, the printed fibers were kept at a certain distance
(20 0–50 0 μm) that could provide several macrospaces for cells to
exchange nutrients and oxygen. The 3D aligned scaffold was re-
moved from the Carbopol-supported gel after photo-crosslinking
and enzyme-crosslinking ( Fig. 4 m). Upon continuous stretching of
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Fig. 6. Inflammation and angiogenesis of cells in the stress-relaxed FG IPN hydrogel-based 3D bioprinted skeletal muscle scaffold group in vivo after implantation for 14
days; the 3D gel group was used as the control (a) and (e). Gross views of the 3D gel and 3D aligned scaffold subcutaneous implantation in vivo after 14 days of surgery.
Scale bar: 500 mm. (b–d) and (f–h) H&E images for the 3D gel and the 3D aligned scaffold implantation after 14 days; yellow arrow: microvessels. Scale bar: 20 0
μm, 10 0
μm, and 30 μm. Quantitative analysis of the number of inflammatory cells (i) the generated capillaries (j), and the capillary density (k) in the region of implanted 3D gel
group and the 3D aligned scaffold group on the 14 th day after surgery.
the stress-relaxed FG IPN hydrogel-based scaffolds, it was found
that the printed aligned scaffolds showed adequate flexibility and
viscoelasticity (Movie S4).
3.4. FG IPN Hydrogel-based 3D bioprinted aligned engineered skeletal
muscle as an in vitro model
Engineered skeletal muscle tissues can be used in in vitro stud-
ies, including disease models and drug screening [53] . Herein, FG
IPN hydrogel-based 3D aligned skeletal muscle scaffolds were pre-
pared as engineered skeletal muscle tissues in vitro . The cell-laden
FG bioink was bioprinted in a Carbopol-supported gel medium
at 15 °C, following which the cell-laden scaffold was formed af-
ter photo-crosslinking and enzyme-crosslinking and then removed
from the Carbopol-supported gel ( Fig. 5 a). Previous studies have
shown that shear stress formed by the hydrogel bioink passing
through the nozzle affects cell viability during the 3D bioprinting
process [54–56] . The viability and morphology of C2C12 myoblast
cells within 3D aligned scaffolds after bioprinting were investi-
gated during culturing for five days. C2C12 cells cultured within
the stress-relaxed FG hydrogel bulk served as the control group
(3D gel group). As shown in Fig. 5 b, the live/dead staining images
showed that most of the cells encapsulated within the scaffolds
presented green fluorescence after one, three, and five days, indi-
cating that most cells were alive. The quantitative analysis also in-
dicated approximately 90% cell viability in the 3D aligned scaffold
group, which was as high as in the 3D gel group ( Fig. 5 c). These
data illustrate that the FG IPN hydrogel bioink with the shear-
thinning property was able to avoid the relatively high shear stress,
which could lead to cellular damage during the bioprinting pro-
cess. In addition, C2C12 myoblast showed extended growth and an
elongated morphology within the 3D aligned scaffold after cultur-
ing for five days compared with the cells that did not fill the fibers
at days 1 and 3 ( Fig. 5 d). The aspect ratios of C2C12 cells within
the 3D aligned scaffold increased from 5.8 ±1. 9 at day 1 to 27.3 ±
9.9 at day 3, and the aspect ratio was as high as 32 ±5.8 after cul-
turing for five days ( Fig. 5 d). The significant increase in cellular as-
pect ratios during five days of cultivation further indicated that the
use of the 3D bioprinted aligned biomimetic scaffold may lead to
cellular elongation. To further investigate the cellular morphologies
of the stress-relaxed FG IPN hydrogel-based 3D bioprinted aligned
biomimetic scaffold, the cytoskeleton (F-actin) and nuclei (DAPI)
were stained after culturing for seven days. The stereomicroscopic
fluorescence images showed complete cell distribution within the
3D aligned scaffold ( Fig. 5 e); further details were obtained using a
confocal microscope. The quantitative analysis results also revealed
that the area coverage of the cytoskeleton showed no significant
difference between the 3D aligned scaffold group and the 3D gel
group ( Fig. 5 f). Moreover, quantitative analysis results showed that
the percentage of cells aligned within ±10 °orientation of the 3D
aligned scaffold group was as high as 26.2 ±8.7%. In contrast, cells
within the 3D gel had an alignment index of 12.7 ±1% ( Fig. 5 g).
Furthermore, the confocal images and the polar distribution of cel-
lular orientation showed elongated cytoskeletons in the 3D aligned
scaffold, which were not apparent in the 3D gel group ( Fig. 5 h and
i). This indicated that the combination of the stress-relaxed FG IPN
hydrogel and the 3D aligned structures fabricated via the printing
gel-in-gel strategy could induce C2C12 cell elongation.
Next, the effect of the 3D aligned scaffold on muscle satellite
cell differentiation was studied. After differentiation for nine days,
MHC staining was performed. Many myotubes were formed and
elongated in the 3D aligned scaffolds, which exhibited the ability
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Fig. 7. S keletal muscle injury repair after scaffold implantation of the 3D aligned scaffold in a rat model for two weeks. (a) H&E staining image of the regeneration muscle
after 14 days of surgery; yellow circle: centronucleated myofibers, arrowed circle: microvessels. Scale bar: 200
μm and 100 μm. Quantitative analysis of the number (b) of
regenerating myofibers (centronucleated myofibers ) and the diameter (c) of centronucleated myofibers in the region of regenerated muscle on the 14t h day after surgery. (d)
Masson staining image of the regenerated muscle after 14 days of surgery. Scale bar: 400
μm. (e) Quantitative analysis of the coverage of the regenerated muscle by collagen.
(f) Immunofluorescence staining image (red: myosin, green: CD31, blue: DAPI) of the regenerated muscle after 14 days of surgery. Scale bar: 10 0
μm. (g) Quantitative analysis
of the generated capillaries in the region of the regenerated muscle on the 14 th day after surgery.
∗p < 0 .05 and
∗∗p < 0.01, compared with the blank group.
to induce myotube fusion formation ( Fig. 5 j). Moreover, bioprinted
fibers were also observed for swaying owing to the contraction re-
sulting from myotubes in the 3D aligned scaffolds (Movies S5 and
S6). The calcium transients in the 3D aligned scaffolds of myotubes
were also investigated using the calcium-sensitive dye Fluo-4 AM.
Spontaneous calcium oscillations in the bioprinted fibers ( Fig. 5 k
and movie S7). In addition, the line scanning image also showed
spontaneous calcium transients across the myotubes in the ROI
and calcium spikes during contraction ( Fig. 5 l). Moreover, dynamic
changes in Ca
2 +
corresponding to calcium spikes also revealed syn-
chronized calcium transients of the myotubes in the 3D aligned
scaffolds ( Fig. 5 l). Thus far, these results indicate that the FG IPN
hydrogel with fast stress relaxation provides a suitable microen-
vironment for 3D cellular proliferation and myotube formation,
while the printing gel-in-gel strategy provides a 3D anisotropic
macrostructure for 3D cellular alignment and elongation, which
presents a potential strategy for the development of in vitro 3D
engineered tissue models.
3.5. Inflammation and micro-vascularization of the FG IPN
hydrogel-based 3D bioprinted scaffold in vivo
We further investigated the potential of FG IPN hydrogel-based
cell-free 3D bioprinted scaffold for in vivo VML repair. First, the
inflammation and i n vivo micro-vascularization of the FG IPN
hydrogel-based 3D bioprinted scaffold were evaluated after subcu-
taneous implantation on the backs of SD rats for two weeks. To
reduce the risk of deformation of the implanted 3D aligned scaf-
fold, a 3D aligned scaffold within an extended border was fabri-
cated (Fig. S9 and S10). The H&E staining images showed that the
3D scaffold group exhibited lower levels of inflammation than the
3D gel groups ( Fig. 6 a–c, e–g). Statistical analysis results demon-
strated that the number of inflammatory in the 3D aligned scaffold
groups was lower than that in the 3D gel groups ( Fig. 6 i). There-
fore, the results of the subcutaneous implantation assessments
suggest that the 3D scaffold f ormulation could increase nutrient
circulation, thereby reducing inflammation and increasing biocom-
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Fig. 8. Functional muscle tests for regeneration in a rat model of skeletal muscle injury after implantation of a stress relaxed FG IPN hydrogel-based 3D bioprinted aligned
scaffold for six months. (a) Ultrasound images of the regenerative muscles with implantation of 3D gel and 3D scaffold; the normal and blank groups were used as controls.
The quantitative analysis of cross-section (b) and side-section (c) area in the region of regeneration muscle for the blank, 3D gel, and 3D aligned scaffold groups during
the sixth month after surgery. The scheme of treadmill experiment (d) and the quantitative analysis of regenerated rats (e)
after surgery of six months. The scheme of
electromyography test (f) and the quantitative analysis of contraction of the regenerative muscles (g) at different stimulus frequencies of the normal, blank, 3D gel, and 3D
scaffold groups.
∗p < 0 .05 and
∗∗p < 0.01, compared with the blank group.
patibility in vivo . Moreover, one of the challenges in the clinical
application of hydrogel scaffolds is their limited ability to pro-
vide sufficient blood supply [57] . Inadequate vascularization causes
cell damage or death in the tissue-engineered scaffolds. There-
fore, 3D biomimetic scaffolds with a suitable macroenvironment
and microenvironment to exchange nutrients could improve pre-
vascularization in vivo [17] . The gross appearance images showed
that some micro-vessels appeared in the implanted regions in the
3D aligned scaffold group after implantation for 14 days ( Fig. 6 a),
whereas the 3D gel group showed no significant microvessel for-
mation ( Fig. 6 e). Furthermore, H&E staining indicated the forma-
tion of blood microvessel walls in the 3D aligned scaffold group. In
contrast, the 3D gel group contained few vessels on day 14 ( Fig. 6 d,
h). To quantitatively count the capillary numbers and coverage, the
H&E images of the 3D aligned scaffold group and the 3D gel group
were further analyzed on day 14. The results indicated that the
number of capillaries in the 3D aligned scaffold group was signif-
icantly higher than that in the 3D gel group on day 14 ( Fig. 6 j).
Furthermore, the coverage of the capillary in the 3D aligned scaf-
fold group was significantly enhanced compared with that in the
3D gel group, which indicated that the pore nanofiber structure
of the stress-relaxed FG IPN hydrogel and the distance between
13

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the bioprinted aligned fibers of the 3D aligned scaffold were use-
ful for vascularization in vivo ( Fig. 6 k). This revealed that stress-
relaxed FG IPN hydrogel-based 3D aligned scaffolds could facilitate
the aggregation of endothelial cells surrounding the scaffold and
are suitable carriers of cells or growth factors to promote capillary
formation in vivo . According to previous studies, the aligned scaf-
fold structures can generate a better-organized muscle-like tissue
compared to bulk hydrogels and promote the formation of more
microvessels [ 11 , 58 ]. Therefore, we propose that the stress-relaxed
FG IPN hydrogel-based 3D bioprinted aligned scaffold can provide
a suitable biomimetic environment for the exchange of nutrition
and efficiently enhance vessel formation in vivo .
3.6. FG IPN hydrogel-based 3D bioprinted scaffold for inducing VML
repair
To further investigate the effect of FG IPN hydrogel-based 3D
bioprinted scaffold on VML repair, we prepared a VML defect in the
rat tibialis anterior muscle, and the 3D bioprinted cell-free scaffold
was then implanted into the VML defect. The VML defects were
treated with normal saline or FG IPN hydrogel bulk and named
the blank and 3D gel groups, respectively. After surgery for two
weeks, H&E staining images showed that the 3D aligned scaffold
group showed more formation of new myofibers than the other
groups ( Fig. 7 a, b), which were evaluated following the measure-
ment methods in previous studies (newly formed myofibers had
nuclei in the center of the cytoplasm, while mature myofibers had
nuclei at the periphery [59] ). Correspondingly, according to the
Masson staining results, where the blue partition represents col-
lagen, while a red represents myofibers in Fig. 8 d, the density
around the fibrotic area was significantly enhanced in the blank
group and the 3D gel group, particularly, in the blank group com-
pared with that in the 3D scaffold group ( Fig. 7 e). The formation of
myotubes and myofiber function is very important in tissue engi-
neering [ 60 , 61 ]. In this study, we found that the number of newly
formed myofibers in the 3D scaffold group was significantly higher
than that in the blank and 3D gel groups, which indicated the in-
creased ability of the 3D scaffolds to induce myofiber formation.
Furthermore, more microvessels were formed in the 3D scaffold
group than in the blank and 3D gel groups ( Fig. 7 a, f). Quantita-
tive analysis also showed that the number of capillaries in the 3D
scaffold group was higher than that in the blank and 3D gel groups
( Fig. 7 g). Recently, several studies have shown that viscoelastic ma-
terials promote tissue regeneration in situ [ 39 , 62 ]. Similarly, the
results of this study suggest that the FG IPN hydrogel-based 3D
aligned scaffold was able to provide a suitable micro-macroscopic
biomimetic structure for recruiting myocytes to differentiate into
myotubes because of the fast stress relaxation properties of the
viscoelastic bioink. Therefore, the stress-relaxed hydrogel-based 3D
bioprinted skeletal muscle scaffold has an exceptional ability to
precisely regulate tissue regeneration and could be applied in the
clinical engineering of skeletal muscle scaffolds in vivo .
The functional recovery of VML repair after surgery for six
months was further investigated. The repaired area of the skele-
tal muscle injury was observed and analyzed quantitatively using
ultrasound imaging ( Fig. 8 a). The cross-sectional and side-sectional
areas of the different groups were not significantly different, indi-
cating that myofibers were filled in the injured area ( Fig. 8 b and c).
In addition, to observe the functional behavior of skeletal muscle
after six months of repair, the treadmill running trial was applied
to test the physical activity of rats; the 3D scaffold group showed
a greater distance than the blank and 3D gel groups ( Fig. 8 d and
e). Both the ultrasound imaging and treadmill running trial re-
sults demonstrated that most myofibers in the blank and 3D gel
groups may not be functionally recovered compared with those
in the 3D scaffold group after six months. Furthermore, the myo-
physiological response was measured by electrostimulation, where
a pulse was applied to the skeletal muscle fibers ( Fig. 8 f). These
results showed that myofibers in the skeletal muscle injury with
the application of the 3D aligned scaffold in the parallel direc-
tion were significantly higher than those in the myofibers in the
blank and 3D gel groups ( Fig. 8 g). Hence, we confirmed that the
stress-relaxed FG IPN hydrogel-based 3D bioprinted aligned scaf-
fold promoted skeletal muscle regeneration. It further proved that
the stress-relaxed FG IPN hydrogel-based 3D bioprinted aligned
scaffold could promote native cell infiltration because of the vis-
coelastic microenvironment, which has the unique ability to en-
hance tissue regeneration and may be applied clinically as an en-
gineered skeletal muscle tissue treatment strategy in vivo .
4. Conclusions
In summary, we developed an IPN hydrogel by the combina-
tion of fibrinogen and GelMA, which showed tunable stress relax-
ation properties by changing the concentration ratio of fibrinogen
and GelMA. Especially, these IPN hydrogels with faster stress re-
laxation showed higher 3D cellular proliferation and better differ-
entiation. Furthermore, we bioprinted an anisotropic biomimetic
scaffold using such IPN hydrogel as bioink materials via the print-
ing gel-in-gel strategy, which was applied for in vitro 3D skele-
tal muscle tissue models and in vivo VML repair, respectively. The
cultivation results of myoblast-laden FG IPN viscoelastic hydrogel-
based 3D bioprinted aligned scaffolds indicated its ability to guide
3D myoblast alignment and elongation in vitro owing to the stress
relaxation characteristics of the FG IPN hydrogel and the macro-
and microenvironment. In addition, the effect of stress-relaxed FG
IPN hydrogel-based 3D bioprinted aligned scaffolds on enhancing
microvascularization and promoting injury repair in vivo was in-
vestigated. We expect that such a stress-relaxed hydrogel based
on fibrinogen and GelMA can provide a mimicking native mi-
croenvironment for enhancing 3D myoblast proliferation and dif-
ferentiation in vitro . Moreover, the stress-relaxed FG IPN hydrogel-
based 3D bioprinted aligned scaffolds also provided a microenvi-
ronment conducive to the recruitment of native cells to promote
skeletal muscle regeneration following VML. These data demon-
strate that fabricating stress-relaxed FG IPN hydrogel-based 3D bio-
printed aligned scaffolds is beneficial not only for building a phys-
iologically accurate 3D in vitro model but also for a variety of ap-
plications in skeletal muscle regeneration in vivo .
Declaration of Competing Interest
The authors declare that they have no known competing finan-
cial interests or personal relationships that could have appeared to
influence the work reported in this paper.
Acknowledgments
This work was supported by the National Natural Science
Foundation of China ( 31900960 , 31972915 , and 21773199 ), the
Science and Technology Program of Guangzhou ( 202102020909 ),
the Guangdong Basic and Applied Basic Research Foundation
( 2019A1515011413 , 2020B1515120 0 01 ), the Science and Technology
Project of Guangdong Province ( 2018B090944002 ), and the San-
ming Project of Medicine in Shenzhen ( SZSM201612019 ).
Supplementary materials
Supplementary material associated with this article can be
found, in the online version, at doi: 10.1016/j.actbio.2022.08.037 .
14

T. Li, J. Hou, L. Wa ng et al. Acta Biomaterialia xxx (xxxx) xxx
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JID: ACTBIO [m5G; August 25, 2022;13:5 ]
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