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\begin{document}

\title{Biomimicking Trilayer Scaffolds with Controlled Estradiol Release for
Uterine Tissue Regeneration}



\author[1]{Shangsi Chen}%
\author[1]{Junzhi  Li}%
\author[1]{Liwu  Zheng}%
\author[1]{Min  Wang}%
\affil[1]{The University of Hong Kong}%


\vspace{-1em}



  
  \date{June 17, 2024}


\begingroup
\let\center\flushleft
\let\endcenter\endflushleft
\maketitle
\endgroup





\selectlanguage{english}
\begin{abstract}
Scaffold-based tissue engineering provides an efficient approach for
repairing uterine tissue defects and restoring fertility. In the current
study, a novel trilayer tissue engineering scaffold with high similarity
to the uterine tissue in structure was designed and fabricated via 4D
printing, electrospinning and 3D bioprinting for uterine regeneration.
Highly stretchable poly(L-lactide-co-trimethylene carbonate)
(PLLA-co-TMC, ``PTMC'' in short)/thermoplastic polyurethane (TPU)
polymer blend scaffolds were firstly made via 4D printing. To improve
the biocompatibility, porous poly(lactic acid-co-glycolic acid)
(PLGA)/gelatin methacryloyl (GelMA) fibers incorporated with
polydopamine (PDA) particles were produced on PTMC/TPU scaffolds via
electrospinning. Importantly, estradiol (E2) was encapsulated in PDA
particles. The bilayer scaffolds thus produced could provide controlled
and sustained release of E2. Subsequently, bone marrow derived
mesenchymal stem cells (BMSCs) were mixed with gelatin methacryloyl
(GelMA)-based inks and the formulated bioinks were used to fabricate a
cell-laden hydrogel layer on the bilayer scaffolds via 3D bioprinting,
forming ultimately biomimicking trilayer scaffolds for uterine tissue
regeneration. The trilayer tissue engineering scaffolds thus formed
exhibited a shape morphing ability by transforming from the planar shape
to tubular structures when immersed in the culture medium at 37 . The
developed trilayer tissue engineering scaffolds would provide a new
insight for uterine tissue regeneration.%
\end{abstract}%



\sloppy


\par\null

\textbf{Biomimicking Trilayer Scaffolds with Controlled Estradiol
Release for Uterine Tissue Regeneration}

Shangsi Chen \textsuperscript{a}, Junzhi Li \textsuperscript{a}, Liwu
Zheng \textsuperscript{b}, Min Wang \textsuperscript{a, *}

\textsuperscript{a} Department of Mechanical Engineering The University
of Hong Kong Pokfulam Road, Hong Kong\textsuperscript{b} Faculty of
Dentistry The University of Hong Kong 34 Hospital Road, Sai Ying Pun,
Hong Kong

\textbf{Keywords:} 4D printing, controlled release, shape morphing

* Corresponding Author:

Professor Min Wang, at the University of Hong Kong, Hong Kong

Email: memwang@hku.hk Tel: +852 3971 7903 Fax: +852 28585415

\textbf{Abstract}

Scaffold-based tissue engineering provides an efficient approach for
repairing uterine tissue defects and restoring fertility. In the current
study, a novel trilayer tissue engineering scaffold with high similarity
to the uterine tissue in structure was designed and fabricated via 4D
printing, electrospinning and 3D bioprinting for uterine regeneration.
Highly stretchable poly(L-lactide-co-trimethylene carbonate)
(PLLA-co-TMC, ``PTMC'' in short)/thermoplastic polyurethane (TPU)
polymer blend scaffolds were firstly made via 4D printing. To improve
the biocompatibility, porous poly(lactic acid-co-glycolic acid)
(PLGA)/gelatin methacryloyl (GelMA) fibers incorporated with
polydopamine (PDA) particles were produced on PTMC/TPU scaffolds via
electrospinning. Importantly, estradiol (E2) was encapsulated in PDA
particles. The bilayer scaffolds thus produced could provide controlled
and sustained release of E2. Subsequently, bone marrow derived
mesenchymal stem cells (BMSCs) were mixed with gelatin methacryloyl
(GelMA)-based inks and the formulated bioinks were used to fabricate a
cell-laden hydrogel layer on the bilayer scaffolds via 3D bioprinting,
forming ultimately biomimicking trilayer scaffolds for uterine tissue
regeneration. The trilayer tissue engineering scaffolds thus formed
exhibited a shape morphing ability by transforming from the planar shape
to tubular structures when immersed in the culture medium at 37 . The
developed trilayer tissue engineering scaffolds would provide a new
insight for uterine tissue regeneration.

\textbf{Introduction}

The uterus in females provides essential biological functions in human
reproduction, such as the implantation and growth of embryos. Congenital
anomalies and acquired diseases caused by intrauterine adhesion (IUA),
infection or hysteromyoma may lead to uterus dysfunction and hence
compromise a woman's ability to be pregnant and/or carry a healthy fetus
to term \textsuperscript{{[}1{]}}. Currently, infertility has been a
severe problem in our society. Infertile male-and-female couples at
reproductive ages suffer significantly from both emotional and mental
problems. About 0.2\% of women have been diagnosed for absolute uterine
factor infertility (AUFI) and approximately 6\% of them need uterine
repair treatments \textsuperscript{{[}2{]}}. Fortunately, the first
birth of a healthy child following uterus transplantation was reported
in the United States in 2018 \textsuperscript{{[}3{]}}, giving infertile
couples a glimpse of hope. Afterwards, successful cases have been
continuously reported about women giving birth to healthy children after
receiving transplanted uterus \textsuperscript{{[}4{]}}. Although uterus
transplantation has been an effective treatment for AUFI, the problems
of donor shortage, possible diseases transmission and use of
antirejection drugs limit its wide applications. Therefore, new
solutions to regenerate the structure and restore the functions of
injured uterus need to be found or developed.

Tissue engineering has been shown to be a promising approach to repair
damaged tissues/organs in the body by using biological substitutes
through combining three-dimensional (3D) scaffolds, cells/stem cells and
biomolecules and therefore overcome the hurdles in tissue or organ
transplantation \textsuperscript{{[}5{]}}. Tissue engineering has
already made great success in treating problems for various tissues and
organs, including bone \textsuperscript{{[}6{]}}, blood
vessel\textsuperscript{{[}7{]}}, skin \textsuperscript{{[}8{]}}, and
bladder\textsuperscript{{[}9{]}}. Several types of natural or synthetic
biodegradable polymer scaffolds encapsulated with appropriate stem cells
or biomolecules have been made and studied for repairing damaged uterine
tissue \emph{in vitro} or \emph{in vivo} owing to their good
biocompatibility and biodegradability \textsuperscript{{[}10{]}}. For
example, collagen-based scaffolds are an attractive option for uterine
regeneration \textsuperscript{{[}11{]}}. Ding \emph{et al.} showed that
3D collagen membrane loaded with bone marrow derived mesenchymal stem
cells (BMSCs) promoted the healing of severe uterine injury in
rats\textsuperscript{{[}12{]}}. They found that collagen/BMSCs
constructs facilitated proliferative abilities of uterine endometrial
and muscular cells and restored the ability of endometrium to receive
embryo and support its development to a viable stage. However, the
progress of using tissue engineering scaffolds for uterine regeneration
is still very limited. Most of studies in this area has remained in the
preclinical stage \textsuperscript{{[}5b, 13{]}}. There is a lack of
clinical studies to examine the therapeutic efficacy of these tissue
engineering scaffolds for uterine tissue regeneration.

On the other hand, another deficiency is that the current efforts for
uterine treatments appear to have centered on endometrium
regeneration\textsuperscript{{[}14{]}}. Indeed, endometrium in uterus is
essential for embryo implantation and pregnancy maintenance because of
its dynamic remodeling process and significant regenerative capacity.
The damage to endometrium does cause infertility. But the uterine tissue
has a hierarchical structure and consists of three layers: outlayer of
perimetrium, interlayer of myometrium and innerlayer of endometrium.
Myometrium containing uterine smooth muscle cells plays an important
role in inducing uterine contraction and supporting stromal and vascular
tissues. An injury of myometrium also significantly affects the uterine
structure and functions and further results in
infertility\textsuperscript{{[}15{]}}. In this context, the reported
studies focusing on endometrium repair may have indicated to some extent
the inability to restore or regenerate the structure and functions of a
whole, multilayered uterine tissue. Additionally, although a recent
study reported the construction of a tissue-engineered uterine scaffold
seeded with autologous stem cells to support live births in
rabbit\textsuperscript{{[}1{]}}, the scaffold had a limited elasticity,
which would hinder its further applications because human uterus has
excellent stretchable properties. Therefore, the development of tissue
engineering scaffolds possessing a multilayered structure and high
elasticity to mimic the inherent structure and properties of natural
uterus would be significantly beneficial for uterine tissue
regeneration.

3D printing has been extensively investigated in the tissue engineering
field because of its great ability to manufacture objects of complex
shapes and structures, as well as products of customized
designs\textsuperscript{{[}16{]}}. To overcome the drawbacks (mainly, in
relation to time, static shapes, properties or functions of printed
objects) of 3D printed structures, 4D printing emerged recently, for
which time as the fourth dimension is integrated with 3D
printing\textsuperscript{{[}17{]}}. On the other hand, 3D bioprinting is
now increasingly used for fabricating cell-laden tissue engineering
scaffolds \textsuperscript{{[}18{]}}. In the current study, a tissue
engineering scaffold mimicking the structure and properties of native
uterine tissue with a multilayered structure and high elasticity was
designed and scaffolds of this design were fabricated via 4D printing,
electrospinning and 3D bioprinting. (Fig.1). Firstly, for mimicking the
highly stretchable myometrium layer of the uterine tissue,
poly(L-lactide-co-trimethylene carbonate) (PLLA-co-TMC, ``PTMC'' in
short) and thermoplastic polyurethane (TPU) were homogenously mixed to
fabricate the PTMC/TPU scaffold layer via fused deposition modelling
(FDM). PTMC/TPU scaffolds thus produced exhibited high stretchability.
Next, poly(lactic acid-co-glycolic acid) (PLGA) and gelatin methacryloyl
(GelMA) mixed solution was employed to fabricate a PLGA/GelMA fibrous
layer on the PTMC/TPU scaffold through electrospinning. Estradiol (E2),
an essential steroid hormone, was encapsulated in the PLGA/GelMA fibrous
layer. E2 could be controllably and sustainably released to regulate
cell behavior. Furthermore, BMSC-laden GelMA/Gel hydrogel was 3D
bioprinted on the PLGA/GelMA fibrous layer to form the complete trilayer
tissue engineering scaffolds. BMSCs showed very high survival rate and
were homogenously distributed in the 3D bioprinted hydrogel layer. The
trilayer scaffolds possessed the layered structure similar to that of
human uterus and were highly elastic. Moreover, the trilayer scaffolds
could evolve from the planar shape to tubular structures when cultured
at 37. Therefore, these trilayer tissue engineering scaffolds with a
hierarchical structure, high stretchability, and controlled and
sustained E2 release would have a high potential for uterine tissue
regeneration.\selectlanguage{english}
\begin{figure}[H]
\begin{center}
\includegraphics[width=0.70\columnwidth]{figures/image1/image1}

\end{center}
\end{figure}

Fig.1 Schematic illustration for the fabrication of trilayer tissue
engineering scaffolds for uterine tissue regeneration. (A) 3D printing
of the PTMC/TPU scaffold layer. (B) Electrospinning to fabricate the
E2-containing PLGA/GelMA fibrous layer on the PTMC/TPU scaffold layer.
(C) 3D bioprinting to construct the BMSC-laden hydrogel layer on the
PLGA/GelMA fibrous layer to complete the fabrication of designed
biomimicking trilayer tissue engineering scaffold.

\textbf{Materials and methods}

\textbf{Materials}

PTMC (molar ratio of LA:TMC = 8:2 and dl/g 0.9) and PLGA (molar ratio of
LA:GA =50:50 and average molecular weight of 90 kDa) were purchased from
Jinan Daigang Ltd., China. TPU 95A was supplied by BASF SE, Germany.
Estradiol (E2) was bought from Aladdin Co. Ltd., China. Gelatin (porcine
skin, type A), bovine serum albumin (BSA), methacrylate anhydride (MA,
94\%), ammonium hydroxide solution (NH\textsubscript{4}OH) (28\%),
absolute ethanol, 2-hydroxy-2-methylpropiophenone (97\%),
1,1,1,3,3-hexafluoro-2-propanol (HFIP) and dopamine hydrocholoride
(98\%) were bought from Sigma-Aldrich (St. Louis, MO, USA). Dialysis
tubing cellulose membrane (MWCO 10kDa) was supplied by Thermo Fisher
Scientific, USA. All reagents were used as received without further
purification.

\textbf{2.2 Fabrication and characterization of PTMC/TPU scaffolds}

Different amounts of PTMC and TPU at specific PTMC:TPU ratios were
homogeneously mixed in round-bottom flasks by mechanical stirring after
being heated in an oil bath to 170  under nitrogen protection.
Afterwards, PTMC/TPU mixtures were cut into pieces and then transferred
to a stainless steel cartridge in a 3D bioprinter (regenHU,
Switzerland). 4D printing was conducted using these printing parameters:
the inner diameter of the printing nozzle was 410 \selectlanguage{greek}μ\selectlanguage{english}m, the printing
temperature was 170 , and the printing speed was 8 mm/s. Subsequently,
4D printed PTMC/TPU scaffolds were rolled up to tubular structures using
a stainless steel rod (rod diameter: 8mm) in an oven at 80 for 30min.
The tubular scaffolds were then flattened at room temperature for
further use.

A scanning electron microscope (SEM, Hitachi S-4800, Japan) was used to
observe the surface morphology and microstructure of 4D printed PTMC/TPU
scaffolds. Differential scanning calorimetry (DSC, Pyris 6, PerkinElmer,
USA) was conducted to analyze the glass transition temperature of
PTMC/TPU scaffolds, and thermogravimetric analysis (TGA, DZ-TGA101,
Nanjing Dazhan Testing Instrument CO., Ltd, China) was performed to
determine the real PTMC:TPU ratios for PTMC/TPU scaffolds. The
mechanical properties of PTMC/TPU scaffolds were studied via tensile
tests using a universal testing machine (Model 5848, Instron Ltd., USA)
at room temperature and at 37 , respectively. The surface wettability
and BSA adsorption of PTMC/TPU scaffolds were studied. Moreover, the
shape morphing behavior of PTMC/TPU scaffolds was investigated by
immersing scaffolds in a water bath at 37 , with shape morphing
processes being recorded by a digital camera.

\textbf{2.3 Synthesis of GelMA, PDA and PDA@E2 particles}

GelMA synthesis was similar to previous
studies\textsuperscript{{[}19{]}}. Briefly, 10 g gelatin was dissolved
in 100 ml phosphate buffered saline (PBS, 0.01M, pH 7.4) at 50 . 0.8 ml
MA was added dropwise into the gelatin solution. After reaction for 1 h,
500 ml PBS was added to stop the reaction and the obtained solution was
transferred into dialysis tubes to dialyze against deionized (DI) water
for 7 days at 40 . The DI water was refreshed daily. GelMA was obtained
after lyophilization. To quantify the degree of substitution of
GelMA,\textsuperscript{1}H-NMR spectroscopy (Bruker Avance III 400, USA)
analysis was conducted.

The synthesis of PDA and PDA@E2 particles followed previously reported
procedures \textsuperscript{{[}20{]}} but with some small modifications.
Because E2 is poorly water-soluble, a water/ethanol mixture was used to
synthesize PDA@E2 particles. Different amounts of E2 (0 or 0.75 g) were
dissolved in absolute ethanol under constant magnetic stirring for at
least 2 h. 10 ml ammonia and 290 ml DI water were added into ethanol
solution under stirring for 30 min. Subsequently, 50 ml 5 w/v\% dopamine
hydrocholoride solution was poured into the ethanol/water mixture under
constant stirring for 24 h at room temperature in open air. PDA and
PDA@E2 particles were obtained after centrifugation at 12,000 rpm/min
for 10 min. The morphology and microstructure of PDA and PDA@E2
particles were characterized using SEM and a transmission electron
microscope (TEM, Tecnai G2 20, USA). The diameter distribution of
particles was analyzed using Image J software.

\textbf{2.4 Fabrication of PLGA/GelMA fibers on PTMC/TPU scaffolds}

15 w/v\% PLGA, 5 w/v\% GelMA and 0.5 w/v\%
2-hydroxy-2-methylpropiophenone were dissolved in HFIP solvent to
prepare a PLGA/GelMA solution for electrospinning. The solution was then
transferred to a 10 ml syringe equipped with a steel needle with an
inner diameter of 0.8 mm. The syringe was fixed to a pump (LongerPump,
LSP01-3A, UK). The steel needle was connected to a high-voltage power
supply (Kou Hing Hong Scientific Supplies Ltd., HK). To determine the
optimal parameters for electrospinning of PLGA/GelMA fibers, different
applied voltages (10, 15, 20 kV) and feeding rates (0.5, 1.0, 2.0 ml/h)
were investigated. The 3D printed PTMC/TPU scaffolds were attached to an
aluminum foil. The aluminum foil was connected to the high-voltage power
supply and was placed 10 cm away from the needle tip to collect
electrospun PLGA/GelMA fibers on the PTMC/TPU scaffolds. The bilayer
scaffolds thus formed were exposed to a UV light (365 nm) for 10 min and
then incubated in an oven at 50  overnight. In the current study, these
bilayer scaffolds were designated as S+F scaffolds. The fabrication of
S+F-PDA scaffolds was similar to the above process but with different
concentrations (0.5, 1.0, 2.5 and 5.0\%, with the concentration
referring to the PDA mass to the total mass of PLGA and GelMA) of PDA
particles dispersed in the PLGA/GelMA electrospinning solution.

The surface morphology and diameters of electrospun PLGA/GelMA fibers
and PLGA/GelMA-PDA fibers were studied using SEM. The mechanical
properties of PLGA/GelMA and PLGA/GelMA-PDA fibers were investigated via
tensile tests. Moreover, the surface and cross-sectional morphology of
S+F-PDA scaffolds were examined through SEM. The mechanical performance
and shape morphing behavior of bilayer scaffolds were also investigated.

\textbf{2.5 Evaluation of photothermal effect and controlled E2 release
from bilayer scaffolds}

The encapsulation of PDA particles in PLGA/GelMA fibers would produce
photothermal effect for the scaffolds. To study the photothermal effect,
bilayer S+F-PDA scaffolds were exposed to a near-infrared (NIR) laser
(808 nm wavelength) at different densities (0.5
W/cm\textsuperscript{2}and 1.0 W/cm\textsuperscript{2}). The temperature
changes were revealed by thermal images which were recorded using an
infrared camera (GUIDE\selectlanguage{ngerman}® EasIR-9, AutoNavi, China). Additionally, to
study the effect of NIR laser irradiation on E2 release, S+F-PDA@E2
scaffolds were exposed to different power densities (0, 0.5, and 1.0
W/cm\textsuperscript{2}) of the NIR laser for 30min every 1 hour in PBS
(pH 7.4) at 37 \selectlanguage{english}. PBS was collected at each predetermined time point,
and an equal amount of fresh PBS was added to the system. The amount of
E2 released in PBS was determined using a UV-vis spectrophotometer
(UV-2600, Shimadzu, Japan) at the wavelength of 280 nm, and cumulative
release curves for E2 were then plotted.

Since human uterus has a dynamic pH environment, it is very important to
investigate E2 release kinetics in solutions of different pH values.
S+F-PDA@E2 scaffold samples were therefore immersed in different buffer
solutions (at pH 4.5, pH 7.4, and pH 9.0) at 37 \selectlanguage{english}. At each predetermined
time point, a small amount of buffer solution was extracted from the
release system and an equal amount of fresh buffer solution was
supplemented. The amount of E2 released was determined using the UV-vis
spectrophotometer. The cumulative amounts of E2 released in different
buffer solutions were calculated, and E2 release curves were
established.

\textbf{2.6 \emph{In vitro} biological evaluation of bilayer scaffolds}

To evaluate the biological properties of bilayer scaffolds, BMSCs
extracted from healthy adult rats were used in \emph{in
vitro}experiments. BMSCs were propagated in high-glucose Dulbecco's
modified Eagle's medium (DMEM, Gibco) with 10\% (v/v) fetal bovine serum
(FBS, Gibco) and 1\% (v/v) penicillin-streptomycin (Gibco). The cell
culture work was performed in a CO\textsubscript{2} incubator
supplemented with 5\% CO\textsubscript{2} and 95\% humidity at 37 \selectlanguage{english}.

The attachment, survival, proliferation and morphology of BMSCs on
PTMC/TPU (designated as ``S''), S+F and S+F-PDA@E2 scaffolds were
studied through SEM observation, live/dead assay (Invitrogen, Thermo
Fisher Scientific), MTT assay (Invitrogen, Thermo Fisher Scientific),
and confocal laser scanning microscopy (LCSM, Leica, Germany) after
phalloidin/DAPI staining (Invitrogen, Thermo Fisher Scientific),
respectively. BMSCs at the density of 5x10\textsuperscript{4} cells per
well were seeded on S, S+F and S+F-PDA@E2 scaffold samples in a 24-well
cell culture plate and cultured in DMEM at 37 \selectlanguage{english}. After being immobilized
by 4\% paraformaldehyde for 30 min, samples were dehydrated by a serial
gradient ethanol solution (50\%, 60\%, 70\%, 80\%, 90\%, and 100\%) for
10 min and then dried in a vacuum oven overnight. The dried samples were
then sputtered with a thin layer of gold, and cell morphology was
observed under SEM.

For cell survival analysis, BMSCs at the density of
5x10\textsuperscript{4} cells per well were seeded on S, S+F and
S+F@-PDAE2 scaffolds in a 24-well cell culture plate and cultured in
DMEM at 37 \selectlanguage{english}. After culture for 24 h and 48 h, respectively, live/dead
assay was used to stain BMSCs. Living cells were stained green and dead
cells were stained red when observed under a fluorescence microscope
(Leica DMi8, Germany). Moreover, cell survival rate was calculated using
Image J software. On the other hand, BMSCs at density of
1x10\textsuperscript{4} cells per well were seeded on scaffold samples
to study the BMSC proliferation rate and cell morphology. After culture
for 1, 3 and 7 days, respectively, BMSC viability on S, S+F and
S+F-PDA@E2 scaffolds was determined using MTT assay. Furthermore, BMSCs
were visualized under CLSM after phalloidin/DAPI staining.

\textbf{2.7 Rheological properties of BMSC-laden GelMA/Gel bioinks}

7 w/v\% GelMA, 3 w/v\% gelatin and BMSCs at the density of
1x10\textsuperscript{6} per ml were mixed to prepare BMSC-laden
GelMA/Gel bioinks for 3D bioprinting. The rheological properties of
GelMA/Gel inks and BMSC-laden GelMA/Gel bioinks were investigated at 20
\selectlanguage{english} using a rheometer (MCR 302, 176 Anton Paar, Austria) equipped with a
parallel plate unit with a 20 mm diameter. In accordance with the
preliminary strain sweep test results, GelMA/Gel inks and BMSC-laden
GelMA/Gel bioinks with or without UV crosslinking were loaded onto a
parallel plate and subjected to a maximum strain (\selectlanguage{greek}γ\selectlanguage{english}) of 1.0\% under
continuous oscillation. In the frequency sweep mode, the storage moduli
(G') and loss moduli (G'') of GelMA/Gel inks and BMSC-laden GelMA/Gel
bioinks were measured in the range of 0.1 - 100 rad/s. Furthermore, the
shear thinning behavior of GelMA/Gel inks and BMSC-laden GelMA/Gel
bioinks was studied. The viscosity of each sample was measured over the
temperature range of 15-50  at 1 s\textsuperscript{-1} shear rate.
Thixotropic tests were conducted in two repeated steps by varying the
shear rate. First, a low shear rate (1 s\textsuperscript{-1}) was
applied for 120 s. Then, a high shear rate (500 s\textsuperscript{-1})
was applied for 60 s.

\textbf{2.8 Fabrication of the complete biomimicking trilayer tissue
engineering scaffolds}

BMSC-laden GelMA/Gel bioinks were used to fabricate the cell-laden
hydrogel layer on S+F-PDA@E2 bilayered scaffolds via 3D bioprinting for
the construction of designed biomimicking trilayer tissue engineering
scaffolds. In the current study, GelMA/Gel inks were homogeneously mixed
with BMSCs at room temperature. The resulting cell-laden bioinks were
subsequently transferred to a cartridge in a 3D bioprinter. The printing
parameters were set as: the inner diameter of the printing nozzle was
210\selectlanguage{greek}μ\selectlanguage{english}m, the printing temperature was 20, and the printing speed was
6mm/s. After adjusting the height between the S+F-PDA@E2 scaffold and
printing nozzle, BMSC-laden GelMA/Gel hydrogels were 3D printed on
S+F-PDA@E2 scaffolds. Subsequently, a UV light (wavelength: 365 nm;
power: 365 mW) was used to crosslink the hydrogel layer for 2 min. The
trilayer tissue engineering scaffolds thus produced were transferred to
6-well cell culture plate and cultured with DMEM at 37 . The
printability of BMSC-laden GelMA/Gel bioinks was evaluated by using the
following formula:

Printability = L\textsuperscript{2}/16A (1)

where L is the perimeter and A is the area of the pore. If the
Printability value is close to 1, it signifies the square shape of the
pores, suggesting good printability of the ink or
bioink\textsuperscript{{[}21{]}}.

\textbf{2.9 Cell survival rate and shape morphing of trilayer tissue
engineering scaffolds}

Live/dead assay was used to examine the BMSC survival rate in the
GelMA/Gel hydrogel layer. After printing and then culturing for 1, 3 and
7 days, respectively, BMSCs in the printed hydrogel layer were stained
by live/dead assay according to the manufacturer's protocol.
Additionally, the BMSC distribution in the final trilayer tissue
engineering scaffolds was visualized via CLSM. The shape morphing
behavior of trilayer scaffolds was recorded using a digital camera when
scaffolds were immersed in the culture medium at 37 .

\textbf{2.10 Statistical analysis}

The results presented in this article were obtained from at least 3
separate samples and were expressed as the mean \selectlanguage{ngerman}± SD. One-way ANOVA was
performed for statistical analysis. Statistically significant difference
existed when: *p\textless{}0.05, **p\textless{}0.01, and
***p\textless{}0.001.

\textbf{3 Results and discussion}

\textbf{3.1 Fabrication, structure and properties of PTMC/TPU scaffolds}

The high stretchability of human uterus is mainly due to the intrinsic
elasticity of myometrium. Elastomeric TPU has been extensively used to
restore damaged vascular and skeletal muscles\textsuperscript{{[}22{]}}.
Moreover, TPU elastomers possess controllable mechanical properties that
can be used to match the targeted body tissues
\textsuperscript{{[}23{]}}. In the current study, to mimic the functions
and properties of myometrium, TPU, as a highly elastic material, is a
good biomaterial for scaffold fabrication. On the other hand, given the
curved shape of the uterus, it is important to fabricate scaffolds that
have the shape morphing ability to form curved or tubular shape to match
the curvature of the repair site in uterus. Previous studies indicated
that biodegradable PTMC polymer exhibited programmed shape morphing from
the planar shape to tubular structure after being incubated at the human
body temperature (37 \selectlanguage{english})\textsuperscript{{[}24{]}}. Therefore, in the
current study, PTMC was chosen to mix homogeneously with TPU for
fabricating PTMC/TPU scaffolds via FDM. To determine the optimal
properties of PTMC/TPU scaffolds (i.e., comparable mechanical strength
with the uterus and good shape-morphing ability), mixtures of different
PTMC:TPU ratios were made under mechanical stirring at 170 \selectlanguage{english}.
Subsequently, PTMC/TPU mixtures were 3D printed to form polymer blend
scaffolds. As shown in Fig.2(A), when the PTMC:TPU ratio was below
0.25:1, PTMC/TPU scaffolds were unable to transform from planar shape to
tubular structures after immersion in water at 37 \selectlanguage{english}. On the contrary, at
higher PTMC:TPU ratios, PTMC/TPU scaffolds exhibited shape morphing
ability {[}Fig.2(B), Video S1{]}. The shape memory effect of PTMC/TPU
scaffolds could be attributed to the suitable glass transition (Tg)
temperature of PTMC. DSC results shown in Fig.2(C) indicated that Tg
temperatures of PTMC and PTMC/TPU scaffolds were at about 36 \selectlanguage{english}. The
shape morphing behavior of these scaffolds could be attributed to the
amorphous nature of the PTMC polymer. After 3D printing, scaffolds were
shaped into tubular structures using a stainless-steel rod in an oven at
80\selectlanguage{english} for 30min. The gradient increase of temperature from the surface to
interior during the heating process affected the degree of molecular
orientation, thereby resulting in anisotropic birefringence and
inhomogeneous transparency. Therefore, PTMC/TPU polymer blend scaffolds
could retain their temporary planar shape at room temperature and could
completely recover their permanent tubular structure at 37 \selectlanguage{english} when the
glass transition of PTMC/TPU scaffolds started to take place.
Furthermore, TGA results {[}Fig.2(D) and Table S1{]} showed that the
real PTMC percentage in PTMC/TPU scaffolds was almost the same as the
nominal percentage, suggesting that PTMC and TPU were homogenously mixed
under mechanical stirring at 170 \selectlanguage{english}.

Since the PTMC/TPU scaffolds were made to mimic the functions of
myometrium, suitable mechanical properties from these scaffolds are
essential. Previous \emph{ex vivo} studies indicated that the ultimate
tensile strength of porcine uterine tissues was 320 +- 176 kPa with a
corresponding strain of 30 +- 9.0\% and that the human uterus exhibited a
better mechanical performance with an average ultimate strength of
656.3\selectlanguage{english} +-\selectlanguage{english} 483.9\selectlanguage{english} kPa at a strain of 32 +- 11.2\% \textsuperscript{{[}25{]}}.
Another study also claimed that the mechanical strength of native
uterine tissues was 0.258 +- 0.071 MPa \textsuperscript{{[}26{]}}.
Moreover, \emph{in vivo} tensile tests suggested that the strain of
human uterus could be up to 110 - 130 \% \textsuperscript{{[}27{]}}. In
this context, for adequately biomimicking the mechanical performance of
uterus, PTMC/TPU scaffolds should have an ultimate strength over 200 KPa
and a strain over 100 \% at the human body temperature. In the current
study, the mechanical strength of PTMC/TPU scaffolds was measured via
tensile tests at room temperature and at human body temperature,
respectively. Fig.3(B-G) shows the mechanical behavior and properties of
PTMC/TPU scaffolds at room temperature. When the PTMC:TPU ratio was 2:1,
PTMC/TPU scaffolds had a high ultimate strength of 1.16 +- 0.31 MPa but a
very low strain of 11.52 +- 2.27 \%, which was not suitable for mimicking
the functions of human uterus. When the PTMC:TPU ratio was at 1:1, 0.5:1
and 0.25:1, PTMC/TPU scaffolds possessed the desirable mechanical
strength of \textasciitilde{}0.6 MPa and an appropriate strain of over
100 \%. Particularly, for scaffolds at the PTMC:TPU ratio of 0.25:1, the
strain was about 400 \%. Because the glass transition temperatures of
PTMC/TPU scaffolds were below 37 \selectlanguage{english}, these scaffolds would become more
elastic at the human body temperature. As shown in Fig.S1 for tests
conducted at 37 \selectlanguage{english}, the strain of PTMC/TPU scaffolds dramatically
increased, while the mechanical strength of PTMC/TPU scaffolds
decreased. The highest mechanical strength of PTMC/TPU scaffolds (at the
PTMC:TPU ratio of 0.25:1) was 0.25 +- 0.02 MPa at 37 \selectlanguage{english}. Consequently,
owing to the comparable mechanical strength to native uterine tissues,
PTMC/TPU scaffolds at the PTMC:TPU ratio of 0.25:1 were used in
subsequent experiments.

The surface morphology of 3D printed PTMC/TPU scaffolds are displayed in
Fig.3(A), showing that scaffolds at different PTMC:TPU ratios all had
relatively smooth surfaces, especially for scaffolds at the PTMC:TPU
ratio of 0.25:1. The smooth surface makes scaffolds less hydrophilic,
which is unfavorable for cells, including their adhesion, proliferation
and differentiation \textsuperscript{{[}28{]}}. Additionally, due to the
high content of TPU, PTMC/TPU (0.25:1) scaffolds had a hydrophobic
surface and the water contact angle was 112.3 +- 6.3deg {[}Fig.S2(A){]}.
The BSA adsorption experiment also indicated that PTMC/TPU (0.25:1)
scaffolds had the lowest BSA adsorption amount {[}Fig.S2(B){]},
suggesting poor surface properties of PTMC/TPU (0.25:1) scaffolds.\selectlanguage{english}
\begin{figure}[H]
\begin{center}
\includegraphics[width=0.70\columnwidth]{figures/image2/image2}

\end{center}
\end{figure}

Fig.2 (A) A diagram illustrating the effect of PTMC:TPU ratio on the
shape morphing ability of PTMC/TPU scaffolds. (B) Shape morphing
behavior of PTMC/TPU scaffolds (PTMC:TPU of 1:0.25) when immersed in the
culture medium at 37 . (C) DSC curves and (D) TGA results of 3D printed
PTMC/TPU scaffolds.\selectlanguage{english}
\begin{figure}[H]
\begin{center}
\includegraphics[width=0.70\columnwidth]{figures/image3/image3}

\end{center}
\end{figure}

Fig.3 Surface morphology and mechanical properties of 3D printed
PTMC/TPU scaffolds. (A) SEM images showing the surface morphology of
PTMC/TPU scaffolds. (B) Tensile stress-strain curves, (C) plastic
strain, (D) ultimate tensile strength, (E) elastic strain, (F) yield
strength and (G) elastic modulus of 3D printed PTMC/TPU scaffolds.

\textbf{3.2 Fabrication, structure and properties of bilayer scaffolds}

The hydrophobic surface of a scaffold is not beneficial for
scaffold-cell interactions and can cause poor tissue regeneration
outcomes \textsuperscript{{[}29{]}}. To improve surface properties,
PLGA/GelMA fibers were constructed on PTMC/TPU scaffolds via
electrospinning. PLGA and GelMA are two popular biomaterials used in
tissue engineering \textsuperscript{{[}30{]}}. Previous studies by a few
groups had employed PLGA and GelMA scaffolds or hydrogels to act as
smart cell or drug delivery system for promoting endometrium
regeneration \textsuperscript{{[}10b, 31{]}}. For example, Chen \emph{et
al.} fabricated GelMA hydrogels carrying human umbilical cord
mesenchymal stem cells (HUMSCs) which could release HUMSCs for
facilitating endometrium regeneration and restoring
fertility\textsuperscript{{[}32{]}}. Therefore, in the current study,
PLGA/GelMA fibers were used not only to improve biocompatibility but
also to function as a smart drug delivery system to controllably and
sustainably release E2 for modulating cell behavior and promoting
uterine tissue regeneration.

As evidenced by the \textsuperscript{1}H-NMR and FTIR results shown in
Fig.S3, GelMA was successfully synthesized. PLGA and GelMA were then
dissolved in HFIP solvent to prepare electrospinning solutions. Many
factors such as solution concentration, applied voltage and feeding rate
can significantly affect the morphology, microstructure and diameter of
resulting electrospun fibers, \textsuperscript{{[}33{]}}. To determine
the optimal electrospinning parameters for PLGA/GelMA fibers, different
applied voltages and feeding rates were used. It was observed that the
average diameter of electrospun PLGA/GelMA fibers increased with the
increase in feeding rate and with the decrease in applied voltage
(Fig.S4-6). In the current study, the electrospun PLGA/GelMA fibers on
PTMC/TPU scaffolds could significantly improve surface wettability and
biocompatibility (Fig.S7). Also, PLGA/GelMA fibers could work as E2
loading and delivering system to controllably and sustainably release
E2. Directly incorporating E2 in PLGA/GelMA electrospinning solution
could impair E2 bioactivity, and also the PLGA/GelMA fibers thus made
may not provide controlled and sustained release of E2. Previous studies
indicated that biomolecules directly incorporated in electrospun fibers
had an initial burst release and exhibited low efficiency for long-term
therapeutic effect \textsuperscript{{[}34{]}}. Therefore, PDA@E2
particles were synthesized in the current study for delivering E2 in the
designed and desired manner. PDA particles could work as an E2 delivery
vehicle, protect its bioactivity and maintain its effectiveness. PDA
particles have been popularly functioned as a drug delivery system
because of their excellent biocompatibility, pH-sensitive properties and
photo-thermal effect \textsuperscript{{[}5d, 35{]}}. Because E2 is
poorly water-soluble, ethanol/water mixture was used to increase the
solubility of E2 in the current study. The incorporation of E2 in PDA
particles had little effect on PDA particle morphology and diameter. PDA
particles had an average diameter of about 1,035 \selectlanguage{ngerman}± 15 nm while the
diameter of PDA@E2 particles was around 947 ± 50 nm {[}Fig.4(A)(B) and
Fig.S8(B){]}. In UV-vis spectra, E2 had a characteristic absorption peak
at 261 nm, while the absorption peak of PDA@E2 particles had a
blue-shift to 256 nm {[}Fig.S8(A){]}.

Different concentrations of PDA particles were mixed with the PLGA/GelMA
solution to construct electrospun fibers on PTMC/TPU scaffolds. As shown
in Fig.4(C), high concentration of PDA particles (5.0\%) could cause
particle aggregations in PLGA/GelMA-PDA fibers, which would dramatically
increase fiber diameter (Fig.S9) and affect mechanical
properties\textsuperscript{{[}36{]}}. Canales \emph{et al.} showed that
the diameter and morphology of electrospun poly(lactic acid) (PLA)
fibers were significantly influenced by the incorporation of high
bioglass particle concentration \textsuperscript{{[}37{]}}. They pointed
out that due to the formation of bioglass particle aggregates, the
mechanical strength of electrospun fibers encapsulated with 20\%
bioglass particles decreased from 0.2 MPa to 0.04 MPa. In the current
study, the PDA encapsulation in PLGA/GelMA fibers increased the tensile
strength but decreased the elongation of fibers. When the concentration
of PDA particles was at 5\%, PLGA/GelMA-PDA fibers had the lowest
elongation at fracture (around 60\%). When the PDA particle
concentration was at 2.5\%, PLGA/GelMA-PDA fibers had the highest
tensile strength (about 4.5 MPa), but its elongation at fracture (about
100\%) did not dramatically decrease. Based on these results,
electrospun PLGA/GelMA fibers having 2.5\% PDA or PDA@E2 particles were
chosen for subsequent experiments.

The bilayer scaffolds were constructed by fabricating PLGA/GelMA-PDA
fibers on PTMC/TPU scaffolds {[}Fig.1(B){]}. In previous studies by
others, there were problems in the integration of two distinct layer for
fabricating bilayer scaffolds \textsuperscript{{[}38{]}}. To tackle the
interface integration problem, after fabricating PLGA/GelMA-PDA fibers
on PTMC/TPU scaffolds, the bilayer scaffolds were incubated in an oven
at 50 \selectlanguage{english} overnight. Since PTMC/TPU scaffolds had low Tg temperatures
{[}Fig.2(C){]}, the glass transition of PLGA took place at 33 \selectlanguage{english} and PLGA
became very viscous at about 50 \selectlanguage{english} (Fig.S10), the high temperature
incubation would make PLGA/GelMA-PDA fibers and PTMC/TPU scaffolds
attached to each other. SEM images in Fig.5(B)(C) show the surface and
cross-sectional views of bilayer scaffolds. It could be seen that
electrospun fibers fully covered the PTMC/TPU scaffold surface. The
cross-sectional view showed that fibers were closely attached to the
PTMC/TPU scaffold surface. In addition, as shown in Fig.S11, at the
initial stage of electrospinning, PLGA/GelMA-PDA fibers merged with the
PTMC/TPU scaffold surface. Subsequently, mechanical properties of
bilayer scaffolds were investigated using tensile tests. The tensile
strain-stress curve in Fig.5(D) showed that the mechanical behavior of
bilayer scaffolds could be divided into two phases: (1) the break of
electrospun fiber, and (2) the break of PTMC/TPU scaffolds. The tensile
strain and strength of S+F-PDA bilayer scaffolds (Table S2) were 315.54
+- 57.47 \% and 0.6 +- 0.06 MPa, respectively, which are superior to
natural uterine tissues. The shape morphing behavior of bilayer
scaffolds was also investigated. Video S2 recorded the shape morphing
process of the S+F-PDA bilayer scaffold, showing that electrospun fibers
had little effect on the shape morphing ability of bilayer scaffolds.
The bilayer scaffolds could easily return to the permanent tubular shape
when immersed in the culture medium at 37 \selectlanguage{english} {[}Fig.5(F){]}.\selectlanguage{english}
\begin{figure}[H]
\begin{center}
\includegraphics[width=0.70\columnwidth]{figures/image4/image4}

\end{center}
\end{figure}

Fig.4 Particle morphology and size distribution of (A) PDA and (B)
PDA@E2 particles. (C) Morphology and size distribution of electrospun
PLGA/GelMA-PDA fibers having different concentrations of PDA particles.\selectlanguage{english}
\begin{figure}[H]
\begin{center}
\includegraphics[width=0.70\columnwidth]{figures/image5/image5}

\end{center}
\end{figure}

Fig.5 (A) Mechanical behavior and properties of electrospun
PLGA/GelMA-PDA fibers. (B) SEM images showing the surface morphology of
a bilayer S+F-PDA scaffold. (C) SEM image showing the cross-sectional
view of a bilayer S+F-PDA scaffold. (D) Characteristic tensile
stress-strain curve of bilayer S+F-PDA scaffolds. (E) Shape morphing of
a bilayer S+F-PDA scaffold from its planar shape to tubular structure
when immersed in the culture medium at 37 . (The black layer was the
PLGA/GelMA-PDA@E2 fibrous layer.)

\textbf{3.3 Photothermal effect and E2 release from bilayer scaffolds}

The application of PDA particles in smart drug delivery systems for
controlled and sustained release of biomolecules has been considered as
an efficient way to promote tissue
regeneration\textsuperscript{{[}39{]}}. Previous studies have already
indicated that oral administration of E2 would have a very poor
therapeutic effect in treating uterine injuries as compared to \emph{in
situ} controlled release \textsuperscript{{[}40{]}}. PDA is an excellent
photothermal agent with good biocompatibility and biodegradability.
Moreover, due to the inherent molecule structure, PDA particles exhibit
a pH-sensitive behavior. Therefore, the release of biomolecules from PDA
particles can be easily regulated by NIR laser irradiation and
environment pH. In the current study, since E2 as a biomolecule for
uterine regeneration was encapsulated in PDA particles and the
synthesized PDA@E2 particles were incorporated in electrospun fibers,
the release of E2 from S+F-PDA@E2 bilayer scaffolds could be precisely
tuned by NIR laser irradiation and/or environment pH. Firstly, the
photothermal effect of S+F and S+F-PDA@E2 bilayer scaffolds were
investigated. Fig.6(A)(B) displays temperature changes of S+F and
S+F-PDA@E2 bilayer scaffolds under the irradiation of the NIR laser at
the energy density of 0.5 and 1.0 W/cm\textsuperscript{2}, respectively.
S+F-PDA@E2 bilayer scaffolds could be heated up to 35  and 49  after
being irradiated by 0.5 W/cm\textsuperscript{2} and 1.0
W/cm\textsuperscript{2} NIR laser, respectively, for 3 min. Such high
temperatures caused by NIR laser could accelerate Brownian movements of
E2 molecules. The accelerated movement of E2 molecules would further
result in more E2 released from bilayer scaffolds. Indeed, with the
irradiation of NIR laser, more E2 was released in comparison to
unirradiated samples {[}Fig.6(C){]}.

On the other hand, because human uterus presents a dynamic pH
environment, it is essential to study E2 release kinetics in buffer
solutions of different pH values. In the current study, the E2 release
behavior were studied in solutions of pH 4.5, 7.4, and 9.0,
respectively. As shown in Fig.6(D), E2 could be sustainably released in
long-term (over 28 days), and the E2 release kinetics could match the
Higuchi model (Fig.S12). The sustained release of E2 would have
significant therapeutic effect on uterine regeneration. It is well-known
that a long-term biomolecule delivery can prolong the beneficial effect
and thus promote tissue regeneration \textsuperscript{{[}41{]}}.
Moreover, in contrast to an acidic environment, E2 could be quickly
released in a basic solution, which may be attributed to the
disintegration of PDA particles and degradation of PLGA/GelMA
electrospun fibers in an alkaline environment.

Estrogen plays a critical role in regulating menstruation and
endometrium regeneration. The safety and efficacy of E2 for endometrium
regeneration has been clinically proved \textsuperscript{{[}42{]}}.
Endometrium stromal cells and glandular epithelial cells in the
endometrium have many estrogen receptors. When E2 binds to their
receptors, it enables increases in expressions of angiogenic growth
factors such as VEGF, bFGF and TGF-\selectlanguage{greek}β\selectlanguage{english}1 \textsuperscript{{[}43{]}}. These
angiogenic growth factors would facilitate endothelial cell migration,
proliferation, differentiation, tube formation, and thus increase
angiogenesis and blood vessel density, thereby improving the
reconstruction of uterine tissues. Previous studies have investigated
the possibility in using hydrogels or scaffolds to deliver E2 \emph{in
situ} for uterus injury treatment \textsuperscript{{[}40, 44{]}}.
However, the problem was that E2 released from these hydrogels or
scaffolds could not reach a controlled and sustained long-term delivery
state. In the current study, the in-site release of E2 from bilayer
scaffolds could not only be regulated by the NIR laser irradiation but
also be modulated by environment pH, which is highly beneficial for
uterine regeneration.\selectlanguage{english}
\begin{figure}[H]
\begin{center}
\includegraphics[width=0.70\columnwidth]{figures/image6/image6}

\end{center}
\end{figure}

Fig.6 Photothermal effect and E2 release behavior of bilayer scaffolds.
(A) Temperature changes of S+F and S+F-PDA@E2 bilayer scaffolds under
the irradiation of an 808 nm NIR laser at different energy densities
(0.5 and 1.0 W/cm\textsuperscript{2}). (B) Infrared thermal images of
S+F and S+F-PDA@E2 bilayer scaffolds after a 3-min NIR laser
irradiation. (C) \emph{In vitro} E2 cumulative release curves of
S+F-PDA@E2 bilayer scaffolds with (0.5 and 1.0 W/cm\textsuperscript{2},
respectively) or without NIR irradiation (pH 7.4). (D) \emph{In vitro}E2
cumulative release curves of S+F-PDA@E2 bilayer scaffolds in pH 4.5,
7.4, and 9.0 solutions, respectively.

\textbf{3.4 \emph{In vitro} biological performance of bilayer scaffolds}

Many \emph{in vitro} and \emph{in vivo} studies have demonstrated the
safety and efficacy of menstrual-derived stem cells (MenSCs) obtained
from women's menstrual fluids for repairing damaged uterine
tissues\textsuperscript{{[}45{]}}. Although MenSCs attract growing
interest in clinical applications because they are multipotent and have
high proliferation rate, the limited source and complex extraction
procedure is problematic. A recent study has indicated that BMSCs had
comparable properties as MenSCs had \textsuperscript{{[}46{]}},
suggesting that BMSCs could replace MenSCs for use in uterine
regeneration. Additionally, many studies have demonstrated the efficacy
of BMSCs for uterine regeneration in preclinical and clinical trials.
BMSCs could differentiate into endometrium epithelial cells so as to
improve gland and blood vessel formations and activate the resident
endometrium stem cells to promote uterine regeneration via the paracrine
effect\textsuperscript{{[}10b, 47{]}}. Therefore, in the current study,
BMSCs were used to evaluate the \emph{in vitro} biological performance
of bilayer scaffolds. Cell attachment results, as shown in Fig.S13,
revealed that BMSCs cultured on S, S+F and S+F-PDA@E2 scaffolds all
exhibited spread morphology, suggesting that all scaffolds were
biocompatible. The live/dead assay results indicated that the BMSC
survival rate in S, S+F and S+F-PDA@E2 scaffolds was over 90 \% after
cultured for 24 and 48 h, respectively {[}Fig.7(A)(B){]}. However, the
proliferation rates of BMSCs on those scaffolds were significantly
different {[}Fig.7(C){]}. Compared to S scaffolds (i.e., PTMC/TPU), S+F
bilayer scaffolds and S+F-PDA@E2 bilayer scaffolds had higher
proliferation rates, suggesting that electrospun PLGA/GelMA fibers on
PTMC/TPU scaffolds had made scaffold surface hydrophilic and improved
scaffold-cell interface behavior and thus promoted cell proliferation.
Moreover, due to the PDA particles and sustained release of E2,
S+F-PDA@E2 bilayer scaffolds had the highest proliferation rate, which
was consistent with previous studies \textsuperscript{{[}40a, 44{]}}.
Furthermore, phalloidin/DAPI staining results in Fig.7(D) indicated that
BMSCs maintained their phenotypes and showed spindle-like morphology on
S+F and S+F-PDA@E2 bilayer scaffolds. Overall, S+F-PDA@E2 bilayer
scaffolds had excellent biocompatibility and promoted BMSC growth.\selectlanguage{english}
\begin{figure}[H]
\begin{center}
\includegraphics[width=0.70\columnwidth]{figures/image7/image7}

\end{center}
\end{figure}

Fig.7 \emph{In vitro} biological performance of S+F-PDA@E2 scaffolds.
(A) Live/dead fluorescence images for BMSCs cultured on bilayer
scaffolds. (B) Cell viability of BMSCs on bilayer scaffolds after
culture for 24 and 48h, respectively. (C) Proliferation of BMSCs on
bilayer scaffolds. (D) F-actin and nucleus fluorescence images of BMSCs
on bilayer scaffolds after culture for 1, 3 and 7 days, respectively.

\textbf{3.5 Fabrication and performance of biomimicking trilayer tissue
engineering scaffolds}

S+F-PDA@E2 bilayer scaffolds providing controlled and sustained release
of E2 are beneficial for endometrium stem cell migration, growth,
proliferation and differentiation and hence can promote uterine
regeneration. Damaged uterine tissues could significantly affect the
growth of endometrium stem cells, which are essential for the dynamic,
cyclic processes of growth, differentiation, sloughing, and renewal of
endometrium. Therefore, introducing autologous stem cells to the damaged
site in uterus could stimulate and activate endometrium stem cells to
grow, proliferate and differentiate \textsuperscript{{[}48{]}}. After
implantation, autologous stem cells would secrete matrix proteins to
build up extracellular matrix (ECM) and attract surrounding cells to
migrate towards the damaged site via auto/paracrine effect, thereby
facilitating tissue remodeling. Meanwhile, new tissue engineering
strategies can combine scaffolds, stem cells and biomolecules for
treating tissue damages. 3D bioprinting, a recently emerged
technological cluster in additive manufacturing (AM), shows increasing
applications in tissue engineering because of its ability to construct
hydrogel-based 3D structures with high spatial distribution of
cells\textsuperscript{{[}18a, 49{]}}. Compared with conventional
biomanufacturing technologies, 3D bioprinting provides a high
cell-loading efficiency and more homogenous cell distribution within the
constructs and can load and release cells to the damaged sites
on-target. Therefore, in the current study, to introduce abundant
autologous stem cells at uterine tissue repair sites, BMSC-laden
hydrogels were 3D printed on bilayer scaffolds for achieving the
designed ultimate trilayer scaffolds for uterine tissue regeneration.

GelMA-based bioinks have been widely used for 3D bioprinting due to
their high biocompatibility, controllable biodegradation rate and
photopolymerizable ability \textsuperscript{{[}50{]}}. So far, many
studies have used GelMA-based bioinks to carry different cells for
repairing various damaged body tissues, including bone, articular
cartilage, blood vessel and skin. The application of 3D bioprinted GelMA
hydrogels loaded with human induced pluripotent stem cell-derived
mesenchymal stem cells (hiMSCs) for uterine endometrium regeneration was
initially reported in 2020 \textsuperscript{{[}10c{]}}. Ji \emph{et
al.}showed that hiMSC-loaded GelMA-based hydrogels could significantly
increase the survival duration of incorporated hiMSCs and promote the
recovery of the endometrial histomorphology and the regeneration of
endometrial cells and endothelial cells. Therefore, in the current
study, GelMA-based bioinks loaded with BMSCs were prepared to fabricate
a cell-laden hydrogel layer on S+F-PDA@E2 bilayer scaffolds, forming the
designed biomimicking trilayer scaffolds. Since pure GelMA inks usually
exhibit poor printability, GelMA/Gel inks were used. Fig.8 provides
rheological properties of GelMA/Gel inks and GelMA/Gel-BMSC bioinks.
Fig.8(A) and (B) show that the G' and G'' of GelMA/Gel-BMSC bioinks
dramatically decreased in comparison with GelMA/Gel inks owing to the
addition of BMSCs. Moreover, because of the loading of BMSCs in
GelMA/Gel bioinks, the shear-thinning behavior became weak and the
gel-sol transition temperature decreased from 32.4  to 25.3 
{[}Fig.8(C) and (D){]}. Thixotropic tests were conducted to evaluate the
recoverability of bioinks. Once a bioink is extruded from the nozzle
tip, it is relieved from the high shear stress in the nozzle and should
have the ability to quickly recover to its initial
state\textsuperscript{{[}51{]}}. Compared to GelMA/Gel inks,
GelMA/Gel-BMSC bioinks were less able to recover their initial state.
The loading of BMSCs in GelMA/Gel bioinks therefore weakened their
rheological properties, which may be attributed to the interference of
BMSCs. The introduced cells in polymer solutions should have interfered
the interaction between polymer chains.

After the rheological studies, GelMA/Gel-BMSC bioinks were 3D printed on
S+F-PDA@E2 bilayer scaffolds. As shown in Fig.9(A) and (B), BMSCs were
homogeneously distributed in the printed hydrogel layer after 3D
bioprinting. However, due to the incorporation of BMSCs, GelMA/Gel-BMSC
hydrogels exhibited lower printability than GelMA/Gel hydrogels
(Fig.S14). Previously, Schwartz \emph{et al} . investigated the effect
of cell encapsulation on the printability of
bioinks\textsuperscript{{[}52{]}}. They found that cell encapsulation in
gelatin bioinks impaired 3D bioprinting resolution and that a high cell
density could significantly affect the printability of gelatin bioinks.
As discussed above, this phenomenon could have resulted from the loose
connections among neighboring polymer chains as the cells may have
blocked their direct contact, causing the reduction in bioink viscosity.
In the cell survival study, the live/dead assay results shown in
Fig.9(C) indicated that BMSCs had very high cell survival rates in the
hydrogels after 3D bioprining, suggesting that the shear stress in the
3D bioprinting process had little effect on cell apoptosis. Moreover,
the cell survival rate of BMSCs dramatically decreased to
\textasciitilde{}82\% after cultured for 1 day but recovered to
\textasciitilde{}95\% after cultured for 7 days {[}Fig.9(E){]}. The
structure of BMSC-laden hydrogels started to disintegrate as the culture
time increased {[}Fig.9(D){]}. Also, the cell survival rates of BMSCs in
the 3D printed hydrogel layer on S+F, S+F-PDA and S+F-PDA@E2 layered
scaffolds did not show significant difference.

The uterus has a hierarchical and curved structure. Although a basic
trilayer tissue engineering scaffold can have a multilayered
biomimicking structure and comparable mechanical strength with natural
uterus, the planar, static shape of a trilayer scaffold is not
sufficient to meet the implantation requirement. Shape memory polymers
with the ability to change shapes of their products upon suitable
stimuli have distinct advantages for obtaining curved or tubular
scaffolds \textsuperscript{{[}53{]}}. Owing to programmed shape morphing
of the PTMC/TPU scaffold layer, the trilayer scaffolds produced in the
current study could transform from the planar shape to tubular
structures when cultured at 37 , as shown in Fig.9(F) and Video S3. To
better visualize the trilayer structure of the scaffolds, a red dye was
added in the hydrogel layer for scaffolds in the control group. The
trilayer structure could therefore be seen clearly in the tissue
engineering scaffolds {[}Fig.9(F){]}: the white outlayer was 3D printed
PTMC/TPU scaffold layer with high elasticity, the black interlayer was
electrospun PLGA/GelMA fibers incorporated PDA@E2 particles and the red
innerlayer was 3D bioprinted BMSC-laden GelMA/Gel hydrogel. Furthermore,
the CLSM images in Fig.9(G) revealed that BMSCs were homogeneously
distributed in the inner GelMA/Gel hydrogel layer of trilayer scaffolds.
As a result, compared to scaffolds made and investigated by other
researchers, the trilayer scaffolds fabricated in this study had (1)
comparable mechanical strength with native uterus, (2) a trilayer
structure that mimicked the hierarchical structure of uterus, (3)
controlled and sustained release of E2 to regulate cell behavior, and
(4) shape morphing ability to form curved or tubular structure after
implantation. Such trilayer scaffolds have the high potential for
uterine regeneration applications.\selectlanguage{english}
\begin{figure}[H]
\begin{center}
\includegraphics[width=0.70\columnwidth]{figures/image8/image8}

\end{center}
\end{figure}

Fig.8 Rheological properties of GelMA/Gel inks and GelMA/Gel-BMSC
bioinks. (A) Storage modulus (G') and loss modulus (G'') of GelMA/Gel
inks at 20  with and without UV crosslinking. (B) G' and G'' of
GelMA/Gel-BMSC bioinks at 20  with and without UV crosslinking. (C)
Shear thinning behavior of GelMA/Gel inks and GelMA/Gel-BMSC bioinks.
(D) Variation of G' and G'' in terms of temperature for GelMA/Gel inks
and GelMA/Gel-BMSC bioinks. (E) Viscosity of GelMA/Gel inks and
GelMA/Gel-BMSC bioinks in thixotropic tests.\selectlanguage{english}
\begin{figure}[H]
\begin{center}
\includegraphics[width=0.70\columnwidth]{figures/image9/image9}

\end{center}
\end{figure}

Fig.9 Characteristic of biomimicking trilayer scaffolds containing a 3D
bioprinted BMSC-laden hydrogel layer. (A) An optical image of 3D printed
GelMA/Gel hydrogel. (B) Optical images and (C) fluorescence images of 3D
bioprinted BMSC-laden hydrogel. (D) Live/dead assay results of BMSCs in
trilayer scaffolds after culturing for 1, 3 and 7 days. (E) Cell
survival rates of BMSCs in trilayer scaffolds after culturing for 1, 3
and 7 days. (F) Shape morphing of a trilayer scaffold from the planar
shape to tubular structure when cultured at 37 . (The red layer in the
scaffold was the 3D bioprinted BMSC-laden hydrogel layer.) (G) CLSM
images showing the 3D distribution of BMSCs in trilayer scaffolds after
scaffold shape morphing at 37 .

\textbf{4 Conclusions}

In the current study, a trilayer tissue engineering scaffold mimicking
the hierarchical structure of native uterine tissue was designed and
trilayer scaffolds of this design were successfully constructed through
4D printing, electrospinning and 3D bioprinting. The trilayer scaffolds
had three distinct layers: 3D printed PTMC/TPU scaffold outlayer with
shape morphing ability and high stretchability, electrospun
PLGA/GelMA-PDA@E2 fibrous interlayer with photothermal effect and
controlled and sustained E2 release, and BMSC-laden hydrogel innerlayer
providing abundant stem cells at the repair/tissue regeneration site.
These trilayer scaffolds were highly stretchable and exhibited
comparable mechanical properties with native uterine tissue.
Additionally, they could deliver E2 controllably and sustainably. The E2
release from trilayer scaffolds was also pH-responsive and could be
tuned by NIR laser irradiation. PLGA/GelMA-PDA@E2 fibers fabricated on
the PTMC/TPU scaffold made the scaffold surface more hydrophilic and
improved biological performance. BMSCs encapsulated in GelMA/Gel bioinks
exhibited high survival rates in the 3D bioprinted hydrogel layer of the
trilayer scaffolds. Importantly, the biomimicking trilayer scaffolds had
the shape morphing ability, transforming from the planar shape to curved
or tubular structures when immersed in the culture medium at 37 . The
biomimicking trilayer tissue engineering scaffolds designed and
fabricated in the current study have the high potential for uterine
tissue regeneration.

\textbf{Supporting information}

Fig.S1 (A) Tensile stress-strain curves, (B) plastic strain and (C)
ultimate tensile strength of PTMC/TPU scaffolds tested at 37 .

Fig.S2 (A) Water contact angle and (B) relative amount of BSA adsorption
by PTMC/TPU scaffolds of different PTMC:TPU ratios.

Fig.S3 (A) Schematic illustration of GelMA synthesis.
(B)\textsuperscript{1}H-NMR and (C) FTIR spectra of Gel and GelMA.

Fig.S4 SEM images and diameter distribution of electrospun PLGA/GelMA
fibers at different feeding rates (Applied voltage: 10 kV).

Fig.S5 SEM images and diameter distribution of electrospun PLGA/GelMA
fibers at different applied voltages (Feeding rate: 2.0 ml/h).

Fig.S6 Fiber diameter analysis for electrospun PLGA/GelMA fibers at
different (A) feeding rates and (B) applied voltages.

Fig.S7 (A) Water contact angle and (B) relative amount of BSA adsorption
by S and bilayer S+F-PDA scaffolds.

Fig.S8 (A) UV-vis spectra and (B) average particle size of PDA and
PDA@E2 particles.

Fig.S9 Average fiber diameter of electrospun PLGA/GelMA fibers having
different concentrations of PDA particles.

Fig.S10 DSC curve of PLGA.

Fig.S11 An SEM image providing the cross-sectional view of a bilayer
scaffold at the initial electrospinning stage.

Fig.S12 Analysis of E2 release using the Higuchi model for S+F-PDA@E2
scaffolds in pH 4.5, 7.4, and 9.0 environments.

Fig.S13 SEM images showing BMSC morphology on S, S+F and S+F-PDA@E2
scaffolds after 1d culture.

Fig.S14 Printability of GelMA-Gel inks and GelMA-Gel-BMSC bioinks.

Table S1 Nominal and real percentages of PTMC in PTMC/TPU scaffolds.

Table S2 Tensile properties of S+F-PDA bilayer scaffolds.

Video S1 The shape morphing process of PTMC/TPU scaffold.

Video S2 The shape morphing process of S+F-PDA bilayer scaffold.

Video S3 The shape morphing process of trilayer scaffold.

\textbf{Author Contribution Statement}

Shangsi Chen: design of the research, methodology, investigation,
writing; Jiahui Lai: methodology, investigation; Liwu Zheng: reviewing,
editing; Min Wang: design of the research, research supervision,
writing, reviewing and editing.

\textbf{Conflict of interest statement}

There are no conflicts of interest to declare.

\textbf{Acknowledgments}

M. Wang thanks Hong Kong's Research Grants Council (RGC) for the
financial support for this work through research grants 17200519,
17202921, 17201622 and N\_HKU749/22 and also The University of Hong Kong
(HKU) for research grants in its Seed Fund for Basic Research Scheme. S.
Chen and J Lai thank HKU for providing them with PhD scholarships.
Assistance provided by members of M. Wang's research group and by
technical staff in HKU's Department of Mechanical Engineering, Faculty
of Dentistry and Electron Microscopy Unit is acknowledged.

\textbf{References}

{[}1{]} R. S. Magalhaes, J. K. Williams, K. W. Yoo, J. J. Yoo, A.
Atala,\emph{Nature Biotechnology} \textbf{2020} , 38, 1280.

{[}2{]} I. Kisu, K. Banno, M. Mihara, N. Suganuma, D.
Aoki,\emph{Fertility and Sterility} \textbf{2013} , 100, 280.

{[}3{]} G. Testa, G. J. McKenna, R. T. Gunby, Jr., T. Anthony, E. C.
Koon, A. M. Warren, J. M. Putman, L. Zhang, G. dePrisco, J. M. Mitchell,
K. Wallis, G. B. Klintmalm, M. Olausson, L. Johannesson, \emph{American
Journal of Transplantation} \textbf{2018} , 18, 1270.

{[}4{]} a) L. Johannesson, G. Testa, J. M. Putman, G. J. McKenna, E. C.
Koon, J. R. York, J. Bayer, L. Zhang, Z. S. Rubeo, R. T. Gunby, A. R.
Gregg, \emph{Obstetrics \& Gynecology} \textbf{2021} , 137, 241; b) E.
G. Richards, R. M. Farrell, S. Ricci, U. Perni, C. Quintini, A. Tzakis,
T. Falcone, \emph{Journal of Assisted Reproduction and
Genetic}\textbf{2021} , 38, 2251; c) M. Brannstrom, C. Racowsky, M.
Carbonnel, J. Wu, A. Gargiulo, E. Y. Adashi, J. M. Ayoubi, \emph{Human
Reproduction Update} \textbf{2023} , DOI: 10.1093/humupd/dmad012dmad012.

{[}5{]} a) D. Brownell, S. Chabaud, S. Bolduc, \emph{International
Journal of Molecular Sciences} \textbf{2022} , 23; b) X. Liu, K. Wu, L.
Gao, L. Wang, X. Shi, \emph{Bioactive Materials} \textbf{2022} , 14, 86;
c) P. V. Torres-Ortega, R. Del Campo-Montoya, D. Plano, J. Paredes, J.
Aldazabal, M. R. Luquin, E. Santamaria, C. Sanmartin, M. J.
Blanco-Prieto, E. Garbayo, \emph{Biomacromolecules} \textbf{2022} , 23,
4629; d) S. Chen, Y. Wang, X. Zhang, J. Ma, M. Wang, \emph{Colloids and
Surfaces B: Biointerfaces} \textbf{2022} , 213, 112364.

{[}6{]} a) M. C\selectlanguage{ngerman}ámara-Torres, P. Fucile, R. Sinha, C. Mota, L.
Moroni,\emph{International Materials Reviews} \textbf{2022} , DOI:
10.1080/09506608.2022.21532191; b) S. Chen, Y. Shi, X. Zhang, J.
Ma,\emph{Materials Science \& Engineering C} \textbf{2020} , 112,
110893; c) S. Chen, Y. Shi, X. Zhang, J. Ma, \emph{Journal of Biomedical
Materials Research Part A} \textbf{2019} , 107, 2512.

{[}7{]} a) L. Debbi, B. Zohar, M. Shuhmaher, Y. Shandalov, I.
Goldfracht, S. Levenberg, \emph{Biomaterials} \textbf{2022} , 280,
121286; b) Y. Wu, L. Song, M. Shafiq, H. Ijima, S. H. Kim, R. Wei, D.
Kong, X. Mo, K. Wang, \emph{Composites Part B: Engineering}\textbf{2023}
, 252.

{[}8{]} a) S. H. Tan, D. A. C. Chua, J. R. J. Tang, C. Bonnard, D.
Leavesley, K. Liang, \emph{Acta Biomaterialia} \textbf{2022} , 153, 13;
b) T. Weigel, C. Malkmus, V. Weigel, M. Wussmann, C. Berger, J.
Brennecke, F. Groeber-Becker, J. Hansmann, \emph{Advanced
Materials}\textbf{2022} , 34, e2106780.

{[}9{]} a) S. Sharma, S. Rajani, J. Hui, A. Chen, T. Bivalacqua, A.
Singh, \emph{ACS Applied Materials \& Interfaces} \textbf{2022} , 14,
37301; b) B. Chen, X. Chen, W. Wang, J. Shen, Z. Song, H. Ji, F. Zhang,
J. Wu, J. Na, S. Li, \emph{Journal of Tissue Engineering} \textbf{2021}
, 12, 2041731420986796.

{[}10{]} a) X. Yi, F. Liu, K. Gao, F. Chen, Y. Wang, H. Li, X. Wang, Y.
Huang, H. Fu, W. Zhou, J. B. Fan, S. Wang, Y. Gao, \emph{Advanced
Materials} \textbf{2022} , 34, e2106510; b) B. Xiao, W. Yang, D. Lei, J.
Huang, Y. Yin, Y. Zhu, Z. You, F. Wang, S. Sun, \emph{Advanced
Healthcare Materials} \textbf{2019} , 8, e1801455; c) W. Ji, B. Hou, W.
Lin, L. Wang, W. Zheng, W. Li, J. Zheng, X. Wen, P. He, \emph{Acta
Biomaterialia} \textbf{2020} , 116, 268.

{[}11{]} a) Z. Fang, C. Lu, W. Du, X. Wang, H. Yang, M. Shi, T. Liu, Y.
Xie, S. Wang, X. Xu, H. Li, H. Wang, Y. Zheng, \emph{International
Journal of Biological Macromolecules} \textbf{2023} , 236, 123943; b) Y.
Rezaeipour, P. Alizadeh, M. Keshavarz, \emph{Applied Materials
Today}\textbf{2023} , 30.

{[}12{]} L. Ding, X. Li, H. Sun, J. Su, N. Lin, B. Peault, T. Song, J.
Yang, J. Dai, Y. Hu, \emph{Biomaterials} \textbf{2014} , 35, 4888.

{[}13{]} X. Li, H. F. Lv, R. Zhao, M. F. Ying, A. T. Samuriwo, Y. Z.
Zhao, \emph{Materials Today Bio} \textbf{2021} , 11, 100101.

{[}14{]} a) S. Lu, X. Wang, W. Li, Y. Zu, J. Xiao, \emph{Advanced
Functional Materials} \textbf{2023} , DOI: 10.1002/adfm.202303368; b) L.
Xin, X. Zheng, J. Chen, S. Hu, Y. Luo, Q. Ge, X. Jin, L. Ma, S.
Zhang,\emph{Advanced Healthcare Materials} \textbf{2022} , 11, e2201680.

{[}15{]} P. Ciarmela, M. S. Islam, F. M. Reis, P. C. Gray, E. Bloise, F.
Petraglia, W. Vale, M. Castellucci, \emph{Human Reproduction
Update}\textbf{2011} , 17, 772.

{[}16{]} a) P. S. Zielinski, P. K. R. Gudeti, T. Rikmanspoel, M. K.
Wlodarczyk-Biegun, \emph{Bioact Mater} \textbf{2023} , 19, 292; b) Q.
Gao, J. S. Lee, B. S. Kim, G. Gao, \emph{International Journal of
Bioprinting} \textbf{2023} , 9, 638.

{[}17{]} a) Y. Wang, H. Cui, T. Esworthy, D. Mei, Y. Wang, L. G.
Zhang,\emph{Advanced Materials} \textbf{2022} , 34, e2109198; b) Y. S.
Lui, W. T. Sow, L. P. Tan, Y. Wu, Y. Lai, H. Li, \emph{Acta
Biomaterialia}\textbf{2019} , 92, 19; c) J. Lai, C. Wang, M. Wang,
\emph{Applied Physics Reviews} \textbf{2021} , 8; d) X. Xu, Z. Fang, B.
Jin, H. Mu, Y. Shi, Y. Xu, G. Chen, Q. Zhao, N. Zheng, T. Xie,
\emph{Advanced Materials} \textbf{2023} , 35, e2209824.

{[}18{]} a) F. Krujatz, S. Dani, J. Windisch, J. Emmermacher, F. Hahn,
M. Mosshammer, S. Murthy, J. Steingrower, T. Walther, M. Kuhl, M.
Gelinsky, A. Lode, \emph{Biotechnology Advances} \textbf{2022} , 58,
107930; b) Z. Fu, L. Ouyang, R. Xu, Y. Yang, W. Sun, \emph{Materials
Today} \textbf{2022} , 52, 112.

{[}19{]} a) J. W. Nichol, S. T. Koshy, H. Bae, C. M. Hwang, S. Yamanlar,
A. Khademhosseini, \emph{Biomaterials} \textbf{2010} , 31, 5536; b) S.
Chen, Y. Wang, J. Lai, S. Tan, M. Wang,
\emph{Biomacromolecules}\textbf{2023} , 24, 2928.

{[}20{]} a) Y. Liu, K. Ai, J. Liu, M. Deng, Y. He, L. Lu, \emph{Advanced
Materials} \textbf{2013} , 25, 1353; b) R. Batul, M. Bhave, J. M. P, A.
Yu, \emph{Molecules} \textbf{2020} , 25.

{[}21{]} L. Ouyang, R. Yao, Y. Zhao, W. Sun,
\emph{Biofabrication}\textbf{2016} , 8, 035020.

{[}22{]} a) S. Gokyer, E. Yilgor, I. Yilgor, E. Berber, E. Vrana, K.
Orhan, Y. A. Monsef, O. Guvener, M. Zinnuroglu, C. Oto, P. Yilgor
Huri,\emph{ACS Biomaterials Science \& Engineering} \textbf{2021} , 7,
5189; b) Z. Fang, Y. Xiao, X. Geng, L. Jia, Y. Xing, L. Ye, Y. Gu, A.-y.
Zhang, Z.-g. Feng, \emph{Biomaterials Advances} \textbf{2022} , 133,
112628.

{[}23{]} J. Tang, Y. Wu, S. Ma, T. Yan, Z. Pan, \emph{Composites Part B:
Engineering} \textbf{2022} , 232.

{[}24{]} a) X. Wan, Y. He, Y. Liu, J. Leng, \emph{Additive
Manufacturing} \textbf{2022} , 53; b) Q. Zhao, J. Wang, H. Cui, H. Chen,
Y. Wang, X. Du, \emph{Advanced Functional Materials} \textbf{2018} , 28;
c) Y.-D. Zhao, J.-H. Lai, M. Wang, \emph{Nano LIFE} \textbf{2021} , 11.

{[}25{]} a) A. Baah-Dwomoh, J. McGuire, T. Tan, R. De Vita,\emph{Applied
Mechanics Reviews} \textbf{2016} , 68; b) S. J. Manoogian, J. A.
Bisplinghoff, A. R. Kemper, S. M. Duma, \emph{Journal of Biomechanics}
\textbf{2012} , 45, 1724.

{[}26{]} E. G. Santoso, K. Yoshida, Y. Hirota, M. Aizawa, O. Yoshino, A.
Kishida, Y. Osuga, S. Saito, T. Ushida, K. S. Furukawa, \emph{PLoS
One}\textbf{2014} , 9, e103201.

{[}27{]} a) Q. Yao, Y. W. Zheng, H. L. Lin, Q. H. Lan, Z. W. Huang, L.
F. Wang, R. Chen, J. Xiao, L. Kou, H. L. Xu, Y. Z. Zhao,\emph{Artificial
Cells, Nanomedicine, and Biotechnology} \textbf{2020} , 48, 218; b) M.
Brannstrom, \emph{Journal of Materials Science: Materials in Medicine}
\textbf{2017} , 28, 70.

{[}28{]} a) Y. Liu, Y. Wang, X. Cheng, Y. Zheng, M. Lyu, P. Di, Y.
Lin,\emph{Materials Science \& Engineering C} \textbf{2021} , 121,
111801; b) B. Abar, C. Kelly, A. Pham, N. Allen, H. Barber, A. Kelly, A.
J. Mirando, M. J. Hilton, K. Gall, S. B. Adams, \emph{Journal of
Biomedical Materials Research Part A} \textbf{2021} , 109, 1792.

{[}29{]} a) Z. C. Chen, T. L. Chang, Q. X. Wu, C. C. Liu, H. C. Chen, C.
H. Huang, \emph{Journal of Alloys and Compounds} \textbf{2023} , 938; b)
J. O. Abaricia, N. Farzad, T. J. Heath, J. Simmons, L. Morandini, R.
Olivares-Navarrete, \emph{Acta Biomaterialia} \textbf{2021} , 133, 58.

{[}30{]} C. Gao, Y. Li, X. Liu, J. Huang, Z. Zhang, \emph{Chemical
Engineering Journal} \textbf{2023} , 451.

{[}31{]} Y. Zhu, S. Li, Y. Li, H. Tan, Y. Zhao, L. Sun, \emph{Chemical
Engineering Journal} \textbf{2022} , 449.

{[}32{]} L. Chen, L. Li, Q. Mo, X. Zhang, C. Chen, Y. Wu, X. Zeng, K.
Deng, N. Liu, P. Zhu, M. Liu, Y. Xiao, \emph{Bioengineering \&
Translational Medicine} \textbf{2023} , 8, e10328.

{[}33{]} H. M. Ibrahim, A. Klingner, \emph{Polymer Testing}\textbf{2020}
, 90.

{[}34{]} a) M. A. Wsoo, S. I. Abd Razak, S. P. M. Bohari, S. Shahir, R.
Salihu, M. R. A. Kadir, N. H. M. Nayan, \emph{International journal of
biological macromolecules} \textbf{2021} , 181, 82; b) M. A. M. Hussein,
E. Guler, E. Rayaman, M. E. Cam, A. Sahin, M. Grinholc, D. S.
Mansuroglu, Y. M. Sahin, O. Gunduz, M. Muhammed, \emph{Carbohydrate
polymers} \textbf{2021} , 270, 118373.

{[}35{]} X. Yang, S. Wang, X. Zhang, C. Ye, S. Wang, X.
An,\emph{Biomaterials Advances} \textbf{2022} , 135, 112677.

{[}36{]} E. K. Tsekoura, T. Dick, P. Pankongadisak, D. Graf, Y. Boluk,
H. Uludag, \emph{Pharmaceuticals} \textbf{2021} , 14.

{[}37{]} D. A. Canales, F. Reyes, M. Saavedra, L. Peponi, A. Leones, H.
Palza, A. R. Boccaccini, A. Grunewald, P. A. Zapata, \emph{International
Journal of Biological Macromolecules} \textbf{2022} , 210, 324.

{[}38{]} M. Yu, D. Luo, J. Qiao, J. Guo, D. He, S. Jin, L. Tang, Y.
Wang, X. Shi, J. Mao, S. Cui, Y. Fu, Z. Li, D. Liu, T. Zhang, C. Zhang,
Z. Li, Y. Zhou, Y. Liu, \emph{Bioactive Materials} \textbf{2022} , 10,
93.

{[}39{]} a) H. Li, D. Yin, W. Li, Q. Tang, L. Zou, Q.
Peng,\emph{Colloids and Surfaces B: Biointerfaces} \textbf{2021} , 199,
111502; b) D. Wu, J. Zhou, X. Chen, Y. Chen, S. Hou, H. Qian, L. Zhang,
G. Tang, Z. Chen, Y. Ping, W. Fang, H. Duan,
\emph{Biomaterials}\textbf{2020} , 238, 119847.

{[}40{]} a) S. S. Zhang, X. X. Xu, W. W. Xiang, H. H. Zhang, H. L. Lin,
L. E. Shen, Q. Lin, F. Lin, Z. Y. Zhou, \emph{The FASEB
Journal}\textbf{2020} , 34, 446; b) B. Li, L. Zhang, Y. Xie, L. Lei, W.
Qu, L. Sui, \emph{Regenerative Therapy} \textbf{2022} , 21, 494.

{[}41{]} a) N. Lin, X. Li, T. Song, J. Wang, K. Meng, J. Yang, X. Hou,
J. Dai, Y. Hu, \emph{Biomaterials} \textbf{2012} , 33, 1801; b) M.
Gruppuso, B. Guagnini, L. Musciacchio, F. Bellemo, G. Turco, D.
Porrelli, \emph{ACS Applied Materials \& Interfaces} \textbf{2022} , 14,
27599; c) J. Kim, B. B. Seo, K. H. Hong, S. E. Kim, Y. M. Kim, S. C.
Song, \emph{Acta Biomaterialia} \textbf{2022} , 144, 183.

{[}42{]} a) D. A. Gibson, I. Simitsidellis, F. Collins, P. T. K.
Saunders, \emph{Journal of Endocrinology} \textbf{2020} , 246, R75; b)
L. Liu, X. Huang, E. Xia, X. Zhang, T. C. Li, Y. Liu, \emph{Human
Fertility} \textbf{2019} , 22, 191.

{[}43{]} J. Johary, M. Xue, X. Zhu, D. Xu, P. P. Velu, \emph{Journal of
Minimally Invasive Gynecology} \textbf{2014} , 21, 44.

{[}44{]} Q. Yao, Y. W. Zheng, Q. H. Lan, L. F. Wang, Z. W. Huang, R.
Chen, Y. Yang, H. L. Xu, L. Kou, Y. Z. Zhao, \emph{European Journal of
Pharmaceutical Sciences} \textbf{2020} , 148, 105316.

{[}45{]} a) L. Chen, L. Guo, F. Chen, Y. Xie, H. Zhang, P. Quan, L.
Sui,\emph{American Journal of Translational Research} \textbf{2020} ,
12, 4941; b) Y. T. Song, P. C. Liu, J. Tan, C. Y. Zou, Q. J. Li, J.
Li-Ling, H. Q. Xie, \emph{Stem Cell Research \& Therapy} \textbf{2021} ,
12, 556.

{[}46{]} F. Alcayaga-Miranda, J. Cuenca, P. Luz-Crawford, C.
Aguila-Diaz, A. Fernandez, F. E. Figueroa, M. Khoury, \emph{Stem Cell
Research \& Therapy} \textbf{2015} , 6, 32.

{[}47{]} Y. Zhou, H. Shen, Y. Wu, X. Zhao, J. Pei, Z. Mou, J. Dong, X.
Hua, \emph{Frontiers in Cell and Developmental Biology} \textbf{2020} ,
8, 52.

{[}48{]} C. E. Gargett, H. P. Nguyen, L. Ye, \emph{Reviews in Endocrine
and Metabolic Disorders} \textbf{2012} , 13, 235.

{[}49{]} a) J. M. Unagolla, A. C. Jayasuriya, \emph{Applied Materials
Today} \textbf{2020} , 18; b) Z. Wang, W. Kapadia, C. Li, F. Lin, R. F.
Pereira, P. L. Granja, B. Sarmento, W. Cui, \emph{Journal of Controlled
Release} \textbf{2021} , 329, 237.

{[}50{]} a) G. L. Ying, N. Jiang, S. Maharjan, Y. X. Yin, R. R. Chai, X.
Cao, J. Z. Yang, A. K. Miri, S. Hassan, Y. S. Zhang, \emph{Advanced
Materials} \textbf{2018} , 30, e1805460; b) M. T. Tavares, V. M. Gaspar,
M. V. Monteiro, J. P. S. Farinha, C. Baleizão, J. F.
Mano,\emph{Biofabrication} \textbf{2021} , 13, 035012.

{[}51{]} H. Rastin, R. T. Ormsby, G. J. Atkins, D. Losic, \emph{ACS
Applied Bio Materials} \textbf{2020} , 3, 1815.

{[}52{]} R. Schwartz, M. Malpica, G. L. Thompson, A. K.
Miri,\emph{Journal of the Mechanical Behavior of Biomedical
Materials}\textbf{2020} , 103, 103524.

{[}53{]} a) W. Yuan, S. Xu, C. Yu, M. Ding, Y. Zheng, J. Zhou, G. Shan,
Y. Bao, P. Pan, \emph{Chemical Engineering Journal} \textbf{2022} , 446;
b) W. Peng, G. Zhang, Q. Zhao, T. Xie, \emph{Advanced
Materials}\textbf{2021} , 33, e2102473.

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