Transplantation of Stem Cell Spheroid-Laden 3-Dimensional Patches with Bioadhesives for the Treatment of Myocardial Infarction

Abstract

Myocardial infarction (MI) is treated with stem cell transplantation using various biomaterials and methods, such as stem cell/spheroid injections, cell sheets, and cardiac patches. However, current treatment methods have some limitations, including low stem cell engraftment and poor therapeutic effects. Furthermore, these methods cause secondary damage to heart due to injection and suturing to immobilize them in the heart, inducing side effects. In this study, we developed stem cell spheroid-laden 3-dimensional (3D) patches (S_3DP) with biosealant to treat MI. This 3D patch has dual modules, such as open pockets to directly deliver the spheroids with their paracrine effects and closed pockets to improve the engraft rate by protecting the spheroid from harsh microenvironments. The spheroids formed within S_3DP showed increased viability and expression of angiogenic factors compared to 2-dimensional cultured cells. We also fabricated gelatin-based tissue adhesive biosealants via a thiol-ene reaction and disulfide bond formation. This biosealant showed stronger tissue adhesiveness than commercial fibrin glue. Furthermore, we successfully applied S_3DP using a biosealant in a rat MI model without suturing in vivo, thereby improving cardiac function and reducing heart fibrosis. In summary, S_3DP and biosealant have excellent potential as advanced stem cell therapies with a sutureless approach to MI treatment.

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Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 1
RESEARCH ARTICLE
Transplantation of Stem Cell Spheroid-Laden
3-Dimensional Patches with Bioadhesives for
the Treatment of Myocardial Infarction
Hye Ran Jeon1†, Jeon Il Kang2†, Suk Ho Bhang3, Kyung Min Park2,4*,
and Dong-Ik Kim1,5*
1Department of Health Sciences and Technology, Samsung Advanced Institute for Health Sciences
and Technology (SAIHST), Sungkyunkwan University, Seoul 06355, Republic of Korea. 2Department
of Bioengineering and Nano-Bioengineering, College of Life Sciences and Bioengineering, Incheon
National University, 119 Academy-ro, Yeonsu-gu, Incheon 22012, Republic of Korea. 3School of
Chemical Engineering, Sungkyunkwan University, Suwon 16419, Republic of Korea. 4Research Center
for Bio Materials & Process Development, Incheon National University, 119 Academy-ro, Yeonsu-gu,
Incheon 22012, Republic of Korea. 5Division of Vascular Surgery, Sungkyunkwan University School of
Medicine, Samsung Medical Center, Seoul 06351, Republic of Korea.
*Address correspondence to: kmpark@inu.ac.kr (K.M.P.); dikim@skku.edu (D.-I. K.)
†These authors contributed equally to this work.
Myocardial infarction (MI) is treated with stem cell transplantation using various biomaterials and methods,
such as stem cell/spheroid injections, cell sheets, and cardiac patches. However, current treatment methods
have some limitations, including low stem cell engraftment and poor therapeutic effects. Furthermore, these
methods cause secondary damage to heart due to injection and suturing to immobilize them in the heart,
inducing side effects. In this study, we developed stem cell spheroid-laden 3-dimensional (3D) patches
(S_3DP) with biosealant to treat MI. This 3D patch has dual modules, such as open pockets to directly deliver
the spheroids with their paracrine effects and closed pockets to improve the engraft rate by protecting the
spheroid from harsh microenvironments. The spheroids formed within S_3DP showed increased viability
and expression of angiogenic factors compared to 2-dimensional cultured cells. We also fabricated gelatin-
based tissue adhesive biosealants via a thiol-ene reaction and disulfide bond formation. This biosealant
showed stronger tissue adhesiveness than commercial fibrin glue. Furthermore, we successfully applied
S_3DP using a biosealant in a rat MI model without suturing invivo, thereby improving cardiac function
and reducing heart fibrosis. In summary, S_3DP and biosealant have excellent potential as advanced stem
cell therapies with a sutureless approach to MI treatment.
Introduction
Cardiovascular diseases, responsible for 17.7 million deaths
annually, have emerged as the primary global cause of mortality.
is number is expected to increase to 23.6 million by 2030,
surpassing that of cancer and other ailments [1]. Among car-
diovascular diseases, myocardial infarction (MI; abbreviations
are listed in Table S1) is a representative cardiac ischemic disease
caused by coronary artery occlusion [2–5]. MI induces a lack
of oxygen and nutrient supply, causing myocardial necrosis, loss
of cardiac function, and eventu al heart failure [2,3]. Furthermore,
unlike other organs, the regenerative ability of the damaged
heart is limited because adult cardiomyocytes have low prolif-
erative and self-renewal capacities [3,5–8]. Conventionally, MI
is treated using a combination of pharmacological and surgical
interventional therapies, such as balloon angioplasty, coronary
bypass, stent insertion, and heart transplantation, to enhance
patient prognosis [2,8–10]. However, these therapies have not
been able to be a fundamental treatment to regenerate injured
myocardium with its function, and postoperative management
is important to reduce infarct size [4]. erefore, it is necessary
to develop treatments that restore the underlying damaged tis-
sue and its function.
Over the last few decades, stem cell transplantation has been
highlighted as an alternative source of the myocardium to treat
ischemic hearts because of their pluri- or multi-potency [8,9].
Mesenchymal stem cells (MSCs) have been extensively used as
promising therapeutic agents for various diseases, extending
beyond MI [9]. Although stem cells have been transplanted to
treat MI, the low retention of stem cell delivery (5% to 10%) in
the myocardium limits their applications [4–6]. erefore, vari-
ous technologies, including spheroid formations, injectable
biomaterials, cell sheets, and cardiac patches, have emerged as
more eective approaches to deliver stem cells with improved
Citation: JeonHR, KangJI, BhangSH,
ParkKM, KimDI. Transplantation
of Stem Cell Spheroid-Laden
3-Dimensional Patches with
Bioadhesives for the Treatment of
Myocardial Infarction. Biomater. Res.
2024;28:Article 0007. https://doi.
org/10.34133/bmr.0007
Submitted 25 September 2023
Accepted 3 January 2024
Published 4 March 2024
Copyright © 2024 Hye Ran Jeon etal.
Exclusive licensee Korean Society
for Biomaterials, Republic of Korea.
No claim to original U.S. Government
Works. Distributed under a Creative
Commons Attribution License 4.0
(CC BY 4.0).

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 2
retention rates to promote the restoration of infarcted myocar
-
dial functions [4,11–14].
Stem cell spheroids are dense 3-dimensional (3D) cell clus-
ters formed by cell-to-cell aggregation. Fabricating spheroids
promotes stem cell survival and retention, and physiologi-
cal and metabolic functions compared to single stem cells
[15,16]. Injectable biomaterials, such as polymeric hydrogels,
have attracted considerable attention as delivery carriers for
cells and spheroids. ey can promote retention rates and
biological functions by providing a 3D articial extracellular
matrix (ECM) that can recapitulate cell-to-matrix interac-
tions and physically support cells and spheroids [1,16,17].
However, these methods inherently require syringe injection,
which causes shear stress on the cells and spheroids, leading
to structural damage to the spheroid and reduced cell viability
[17,18]. Furthermore, physical injury caused by the injection
can cause secondary damage, bleeding, and myocardial per-
foration, leading to possible adverse cardiac events [7,19]. As
with other strategies, cell sheets and cardiac patches have
been widely utilized due to their improved cell-to-cell com-
munication and more powerful paracrine eects, improving
prognosis compared to stem cell solution therapy [9,13].
Although these strategies can directly or indirectly deliver
stem cells and their cytokines/growth factors to treat MI,
suturing is essential to immobilize them onto the epicardial
surface and induce secondary damage [9,20,21]. erefore,
it is necessary to develop advanced stem cell therapies to
improve the stem cell engrament rate and paracrine eects
using a sutureless approach.
Polymeric hydrogel-based biosealants have recently emerged
as a promising alternative for suturing and stapling because of
their sol-gel phase transition, bioactivity, and strong tissue adhe-
sion [22,23]. erefore, they have been used to prevent gas/
liquid leakage and bleeding and immobilize implantable medi-
cal devices on target tissues [24]. Currently, various tissue adhe-
sive biosealants, such as Tisseel, Coseal, Duraseal, Progel, and
others, have been successfully approved by the Food and Drug
Administration and are widely used in clinics [24]. However,
these commercialized biosealants have relatively lower tissue
adhesiveness when applied to myocardial tissue owing to the
presence of the body or pericardial uid and the dynamic pulsa-
tion of the heart [20,25–27]. erefore, there is a need to develop
advanced functional biosealants with strong tissue adhesion.
Toward this, the biosealant that will be applicated for heart tis-
sue has to meet some requirements, such as (a) rapid and easy
application for clinical compliance, (b) biocompatibility of poly-
mer and chemistry, (c) mechanical similarity to surrounding
tissue and elasticity to heartbeats, (d) strong tissue adhesion on
the heart surface under the pericardial uid and the dynamic
heartbeats, and (e) biodegradability.
In this study, we developed 2 types of stem cell spheroid-laden
3D patches (S_3DP): (a) open pockets to directly deliver stem
cell spheroids with an increased paracrine eect, and (b) open/
closed pockets to enhance the engra rates of spheroids by
protecting them from harsh microenvironments (Fig. 1A). e
stem cell spheroids formed within these patches revealed high
cell viability with improved bioactivities, such as the expres-
sion of genes and proteins and paracrine ability related to
Fig.1.Schematic illustration of open pocket and open/closed pocket 3D patches laden with human adipose-derived stem cell (hADSC) spheroids for effective MI treatment.
(A) By forming hADSC spheroids in the open or closed pocket of the 3D patches, cell viability and expression of the angiogenic growth factors were increased. (B) The
transplantation of S_3DP using a biosealant onto the infarcted region of the rat heart and its therapeutic effects.

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 3
angiogenesis, in vitro. We also developed a tissue adhesive
biosealant as a sutureless technique to immobilize functional
S_3DP on the heart tissue without physical damage. Finally, we
transplanted S_3DP into a rat MI model using a biosealant and
demonstrated its eect on cardiac function and brosis in vivo
(Fig. 1B). e entire S_3DP preparation and transplantation
process for MI treatment is depicted in Fig. 1.
Materials and Methods
Materials
For fabrication of the patch and stem cell spheroid, Dulbecco’s
modified Eagle’s medium (DMEM), fetal bovine serum
(FBS), penicillin/streptomycin antibiotics (P/S), and trypsin
were purchased from Gibco BRL (Gaithersburg, MD, USA).
Phosphate-buered saline (PBS) was purchased from Biosesang
(Seongnam, Korea). Human adipose-derived stem cell (hADSC)
was obtained from Lonza (Walkersville, MD, USA). e biopsy
punch was purchased from Kai Medical (Gifu, Japan). For the
in vitro spheroid study, TRIzol was purchased from Ambion
(Austin, TX, USA). Chloroform, isopropyl alcohol, ethanol,
and β- mercaptoethanol were purchased from Sigma-Aldrich
(St. Louis, MO, USA). AccuPower RocketScript Cycle RT
PreMix and AccuPower 2X Greenstar qPCR MasterMix were
purchased from Bioneer (Daejeon, Korea). PRO-PREP Protein
Extraction Solution was obtained from iNtRON Biotechnology
(Seongnam, Korea). Bradford reagent and 4× Laemmli sample
buer were obtained from Bio-Rad (Hercules, CA, USA). Five
percent skim milk was purchased from BD Difco (Detroit, MD,
USA). Antibodies of glyceraldehyde 3-phosphate dehydroge-
nase (GAPDH), Bcl-2-associated X (BAX), and hepatocyte
growth factor (HGF) were purchased from Abcam (Cambridge,
MA, USA). Antibodies against E-cadherin and B-cell lymphoma
2 (BCL-2) were purchased from Cell Signaling Technology
(Danvers, MA, USA), and the vascular endothelial growth fac-
tor (VEGF) antibody was purchased from Santa Cruz Bio-
technology (Dallas, TX, USA). Goat anti-mouse IgG-HRP and
goat anti-rabbit IgG-HRP antibodies were purchased from
Bethyl Laboratories (Montgomery, TX, USA). e ECL reagent
WESTSAVE UP was purchased from AbFrontier (Seoul, Korea).
e x-ray lms were obtained from AGFA HealthCare NV
(Mortsel, Belgium). e Proteome Proler Human Angiogenesis
Array Kit was purchased from R&D Systems (Minneapolis,
MN, USA). For in vivo animal study, 8- to 10-week-old male
Sprague–Dawley rats were supplied by Orient Bio Inc. (Seong nam,
Korea). 6-0 polypropylene, 4-0 surgifit, and 6-0 black silk
sutures were obtained from Ailee (Busan, Korea). Ten percent
formalin and paran were purchased from Sigma-Aldrich
(St. Louis, MO, USA). Xylene was obtained from Daejung
(Siheung, Korea). e hematoxylin and eosin (H&E) staining
kit was purchased from Abcam (Cambridge, MA, USA), and
Masson’s trichrome was purchased from Dako (Hamburg,
Germany). A Millipore Milli-Q purication system treated with
water was used for all the experiments.
For polymer synthesis and characterization, gelatin (Gtn,
type A from porcine skin, <300 bloom), cystamine dihydro-
chloride (CYS), 6-maleimidohexanoic acid (MHA), 1-ethyl-
3-(3-dimethylaminopropyl)carbodiimide hydrochloride (EDC),
N-hydroxysuccinimide (NHS), DL-dithiothreitol (DTT), and
deuterium oxide (D2O) were purchased from Sigma-Aldrich
(St. Louis, MO, USA). Dulbecco’s phosphate-buered saline
(DPBS) was obtained from Gibco (Grand Island, NY, USA).
Dialysis membranes (molecular cuto = 3.5 kDa) were sup-
plied by Spectrum Laboratories (Rancho Dominguez, CA,
USA). Hydrochloric acid (HCl, 1N) was obtained from Daejung
(Siheung, Korea). 5,5′-Dithio-bis(2-nitrobenzoic acid) (Ellman’s
reagent) was provided by ermo Fisher Scientic (Rockford,
IL, USA). For the in vitro cell study, human dermal broblast
(HDF) was purchased from Lonza (Walkersville, MD, USA).
DMEM, newborn calf serum (NBCS), P/S, and 0.25% trypsin-
EDTA were supplied by Gibco (Grand Island, NY, USA). e
Live/Dead kit was obtained from Invitrogen (Grand Island,
NY, USA). e WST-1 cell proliferation kit was purchased
from Roche (Basel, Switzerland). For the in vivo animal study,
5-week-old female C57BL mice were obtained from Koatech
(Gyeonggi-do, Korea). Isourane (Forane Solution) was pur-
chased from Hana Pharm Co. (Seongnam, Korea). Formalin
(10%) was purchased from Sigma-Aldrich (St. Louis, MO, USA).
Hematoxylin H and aqueous eosin Y solutions were purchased
from Sigma-Aldrich (St. Louis, MO, USA).
Fabrication of open and open/closed pocket patches
for stem cell spheroid encapsulation
In our previous study, we developed an elastomeric bioscaold
(TPU-CEC363) with tunable biodegradability and elasticity for
noninvasive stem cell-based therapy [28]. In this study, devel-
oped elastomers were used as bioinks to fabricate stem cell
spheroid-encapsulating patches. Briey, we synthesized TPU-
CEC363 using poly(ethylene glycol) (PEG; Sigma-Aldrich),
ε-caprolactone (Tokyo Chemical Industry, Tokyo, Japan), iron
(III) acetylacetonate (Sigma-Aldrich), and toluene (Daejung).
TPU-CEC363 was then synthesized using the tri-block copo-
lymer, toluene, and hexamethylene diisocyanate. As a result,
TPU-CEC363 containing 50% each of PEG and polycaprolac-
tone was prepared, and the synthesized TPU-CEC363 was
extruded to a diameter of 1.75 mm [28]. To fabricate the 3D
patches, we set the printing program to continuously eject the
bioink under the following nozzle conditions. We used a print-
ing nozzle with a diameter of 300 μm to print the patches
(25 × 25 mm) with a thickness of 100 μm in each layer. e
3D patches were fabricated using FDM Vis Power Plus (Vision
Technology Korea, Daejeon, Korea) with a ow rate of 1.196
and a nozzle temperature of 230 °C.
Characterization of open and open/closed
pocket patches
e 3D-printed patches with a size of 25 × 25 mm were supplied
by the Korea Institute of Science and Technology and Vision
Technology Korea. e 3D patches were cut using a biopsy
punch with a diameter of 4 mm, and the number of pockets in
each patch was observed under a light microscope (CKX41,
Olympus, Tokyo, Japan) at ×40 magnication. e swelling abil-
ity was performed with 1× PBS, and photographs of the pockets
were captured using a light microscope (CKX41, Olympus) at
×100 magnication and Innity analyze soware (Lumenera
Corporation, Ottawa, Canada). e pocket size was analyzed
using ImageJ soware (National Institutes of Health, Bethesda,
MD, USA).
Preparation of 3D patch and spheroid formation in
3D patch
e 3D patches with a diameter of 4 mm were placed into each
well of a 96-well plate, and the 96-well plate without a lid was

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 4
placed in a sterilized pack and sterilized with ethylene oxide gas
at 37 °C for 3 h. A 96-well plate containing sterilized 3D patches
was refrigerated and protected from light to slow degradation.
Before use, the 3D patches were swollen by adding 250 μl of 1×
PBS to the wells of a 96-well plate. e 3D patches in PBS were
centrifuged at 1,500 rpm for 5 min to remove air within the
pockets, and the PBS was replaced with 200 μl of serum-free
DMEM. e hADSCs were cultured with DMEM containing
10% (v/v) FBS and 1% (v/v) P/S in a 5% carbon dioxide (CO2)
incubator at 37 °C. e medium was changed every 2 days, and
hADSCs from passages 4 to 7 were used in the experiments.
Serum-free DMEM (50 μl) containing 3 × 105 cells was dis-
pensed over the swollen 3D patches within the wells of a 96-well
plate. e hADSCs entered the pockets of the 3D patches aer
centrifugation at 1,500 rpm for 5 min, and spheroids were formed
in the open and open/closed pockets aer 24 h. e sample codes
and experimental conditions are described in Table 1.
Characterization of stem cell spheroid
We observed the spheroid morphology under a light micro-
scope (CKX41, Olympus) at ×40 magnication aer 24 h.
Images were captured to conrm the diameter of the spheroids
using a light microscope (CKX41, Olympus) at ×100 magni-
cation and Innity analyze soware (Lumenera Corporation),
and the diameter was analyzed using ImageJ soware.
Quantitative reverse transcription-polymerase
chain reaction
Total RNA was extracted using 1 ml of TRIzol and 200 μl of
chloroform. Aer centrifugation at 11,000 rpm for 10 min at
4 °C, 80% (v/v) isopropyl alcohol (in water) and 75% (v/v)
ethanol (in water) were used for washing. e samples were
then dissolved in RNase-free water. Total RNA (1 μg) and the
AccuPower RocketScript Cycle RT PreMix were used to syn-
thesize complementary DNA. e AccuPower 2X Greenstar
qPCR MasterMix and QuantStudio 6 Flex Real-Time PCR
System (Applied Biosystems, Waltham, MA, USA) were used
for quantitative reverse transcription-polymerase chain reac-
tion (qRT-PCR). e relative gene expression was analyzed
using the 2−ΔΔCT method, and GAPDH served as the internal
control. e primer sequences are described in Table S2.
Western blot analysis
Spheroids were extracted using PRO-PREP Protein Extraction
Solution to prepare protein samples. Protein concentration was
determined using the Bradford assay. Protein samples were
boiled at 100 °C for 10 min in 4× Laemmli sample buer con-
taining β-mercaptoethanol, and samples were loaded onto a
sodium dodecyl sulfate-polyacrylamide gel electrophoresis
(SDS-PAGE) gel. Aer gel electrophoresis, proteins were trans-
ferred from gel to nitrocellulose membrane, and membranes
were continuously blocked with 1× TBS-T containing 5% skim
milk for 1 h at room temperature (RT). e membranes were
incubated overnight at 4 °C with primary antibodies: anti-
GAPDH (1:3,000), anti-E-cadherin (1:500), anti-BAX (1:250),
anti-BCL-2 (1:250), anti-HGF (1:500), and anti-VEGF (1:250).
e membranes were washed with 1× TBS-T for 45 min and
incubated for 1 h at RT with goat anti-mouse IgG-HRP (1:5,000)
and goat anti-rabbit IgG-HRP (1:10,000) secondary antibodies.
Aer incubation, the membranes were washed with 1× TBS-T
for 45 min, and protein bands were detected using the ECL
reagent WESTSAVE UP and an x-ray lm. To analyze protein
expression, GAPDH and ImageJ soware were used as a house-
keeping control and an analysis tool.
Angiogenesis antibody array
To prole angiogenesis-related proteins in the conditioned
medium (CM) of the spheroids, we used the Proteome Proler
Human Angiogenesis Array. Aer spheroid formation in the
3D patches, the cells were carefully washed with 1× PBS, and
the spheroids in the 3D patches were incubated with fresh
serum-free DMEM for 24 h to harvest the CM. e protein
concentration of the harvested CM was measured by Bradford
assay. e array membranes were blocked according to the
manufacturer’s instructions, and the buer was aspirated. e
sample/antibody mixtures were added to the membranes and
incubated overnight at 4 °C. e following day, we washed each
membrane with 1× wash buer, and membranes were incubated
with streptavidin-HRP for 30 min at RT. Aer washing, the
membranes were covered with Chemi reagent mix and exposed
to x-rays. Pixel densities on the developed x-ray lm were quan-
tied at each spot using ImageJ soware.
Synthesis and characterization of
biosealant polymers
To fabricate the biosealant for xing S_3DP on the heart surface,
we synthesized thiolated gelatin (GtnSH) and maleimide-
conjugated gelatin (GtnMI) via EDC/NHS-mediated conjuga-
tion of thiol and maleimide groups to Gtn backbone. To
synthesize GtnSH, we dissolved 250 mg of gelatin in 125 ml
of deionized water (DIW) at 37 °C. Aer Gtn dissolution,
450.5 mg of CYS (2.0 mmol) was dissolved in 2.5 ml of DIW
and added into the Gtn solution. We homogenously mixed them
for 10 min at 37 °C. Next, 192.0 mg of EDC (1.0 mmol) and
116.0 mg of NHS (1.0 mmol) were dissolved in DIW (2.5 ml) in
a 10-ml vial, respectively. EDC and NHS solutions were sequen-
tially added to Gtn/CYS and allowed to react for 2 h at 500 rpm.
Aer the conjugation reaction, we injected a DTT solution
(617.0 mg, 4.0 mmol) dissolved in 5 ml of DIW into the reactant
and incubated it for 24 h at 500 rpm. is reactant was serially
dialyzed using a dialysis membrane bag (molecular weight cut-
o = 3.5 kDa) against a 5 mM HCl solution for 36 h, followed
by a 1 mM HCl solution for 24 h. e solution was changed twice
to remove unconjugated residual chemicals, such as unconju-
gated CYS, by-products of the EDC/NHS reaction, and unreacted
DTT molecules. To obtain the GtnSH polymer, we froze the dia-
lyzed GtnSH solution at −80 °C and lyophilized it for 5 to 6 days.
e obtained polymers were stored in a desiccator until use.
To synthesize GtnMI, we dissolved 250 mg of Gtn and
211.2 mg of MHA (1.0 mmol) in 50 ml and 120 ml of DPBS
solution at 37 °C, respectively. Each solution was mixed for
10 min at 37 °C. en, 230.1 mg of EDC (1.2 mmol) and
161.1 mg of NHS (1.4 mmol) were dissolved in 5 ml of DPBS
within a 10-ml vial, respectively. e EDC solution was mixed
with the NHS solution and added to the Gtn/MHA mixture.
Aer 2 h of reaction, the reactant was dialyzed using a dialysis
membrane bag (molecular weight cuto = 3.5 kDa) against
DIW for 72 h to remove the unconjugated MHA molecules
and by-products of the EDC/NHS reaction. Aer dialysis, the
GtnMI solution was frozen at −80 °C and lyophilized for 5 to
6 days. Lyophilized GtnMI were stored in a deep freezer at
−80 °C without light.

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 5
A proton nuclear magnetic resonance (
1
H-NMR) spectrom-
eter (Agilent 400–MR, Agilent Technologies, CA) was used to
characterize the chemical structures of GtnSH and GtnMI. We
prepared 600 μl of GtnSH and GtnMI solutions (10 mg/ml)
using D2O and analyzed each chemical structure.
Next, Ellman’s assay was performed to determine the func-
tional groups of the biosealant polymers. To measure the thiol
content of GtnSH, we prepared 0 to 1 mg/ml L-cysteine solu-
tions using DIW to obtain a standard curve. Next, 100 μl of
GtnSH solution (1 mg/ml) dissolved in DIW and standard
solution were mixed with 100 μl of Ellman’s reagent solution.
Aer 20 min of reaction in the dark, the absorbance of the
samples was measured at 405 nm using an ultraviolet (UV)
spectrophotometry (Multiskan EX, ermo Fisher Scientic,
Rockford, IL, USA). iol content was calculated from a thiol
standard curve at known L-cysteine concentrations (0.001 to
0.04 mg/ml). To measure the maleimide content of GtnMI, we
performed a modied Ellman’s assay by introducing a thiol-ene
reaction. We prepared a 0.04 mg/ml L-cysteine solution as
a substrate for the thiol-ene reaction. Aer mixing 50 μl of
0.04 mg/ml L-cysteine solution with 50 μl of standard MHA
solution or 50 μl GtnMI solution, we incubated them to induce
the thiol-ene reaction for 10 min without light. We next added
50 μl of Ellman’s reagent solution to quantify the degree of thiol
group reduction depending on the content of the maleimide
group in the standard solution and GtnMI solution. After
20 min of incubation in the dark, we measured the absorbance
at 405 nm using UV spectrophotometr y (Multiskan EX, ermo
Fisher Scientic, Rockford, IL, USA). e maleimide content
was calculated from the maleimide standard curve at known
MHA concentrations (0.005 to 0.04 mg/ml).
Fabrication of tissue adhesive biosealant
We fabricated a tissue adhesive biosealant by simply mixing
GtnSH, GtnMI, and calcium peroxide (CaO2) solutions (vol-
ume ratio of GtnSH:GtnMI:CaO2 = 10:9:1). In this system,
CaO2 was used as an additional crosslinker to enhance the
mechanical property of the biosealant. We rst dissolved 20 mg
of GtnSH and GtnMI in 200 μl and 180 μl of prewarmed DPBS,
respectively. We loaded the GtnSH solution into a 1-ml com-
mercial syringe (Korea Vaccine Co., Ltd., Gyeonggi-do, Korea).
Next, a 20-μl CaO2 solution (0 to 2 wt%) prepared using 1 M
Tris-HCl was mixed with 180 μl of GtnMI solution and loaded
into another 1-ml commercial syringe. Finally, both syringes
were equipped with a dual syringe kit and the solution was
injected to fabricate the tissue adhesive biosealant.
Rheological analysis of biosealant
We determined the elastic modulus (G′) of the biosealant using
a rheometric uid spectrometer (DHR-1, TA instruments, New
Castle, DE, USA) in oscillatory mode. We prepared each 150-μl
GtnSH and GtnMI/CaO2 solution and equipped them into a
dual syringe kit. Next, we injected a biosealant solution on the
parallel plate (diameter, 20 mm) with a gap of 600 μm and
performed dynamic time sweeps on samples depending on
CaO2 concentrations at a frequency of 0.1 Hz and a strain of
0.1% at 37 °C. To prevent solvent evaporation, a solvent trap
was lled with DIW.
Tissue adhesive test of biosealant
e tissue adhesiveness of the biosealant was measured using
a universal testing machine (UTM, UNITEST M1, TEST
ONE, Busan, Korea) with a load cell sensor (LCK1205-K010)
according to the modied ASTM standard F2255-05 method.
We prepared the decellularized porcine skin (round shape,
10 mm diameter) using a 0.1% SDS solution and a freeze
dryer. Next, the decellularized porcine skin was attached to
an acrylic plate (1 × 4 cm) using cyanoacrylate. Aer swelling
of decellularized porcine skin, we injected a biosealant solu-
tion using a dual syringe lled with 100 μl of GtnSH and
GtnMI/CaO2 solutions between 2 pieces of decellularized
porcine skin. We stabilized samples for 30 min under a force
of 100 g at 37 °C within a humid chamber. e tissue adhesion
force was measured at a speed of 1 mm/min. To evaluate the
feasibility of the tissue adhesive biosealant in vivo, we treated
biosealant to various mouse tissues, such as the heart, liver,
spleen, lung, kidney, and skin. Various organs and skin tissues
were harvested and rinsed with DPBS to remove blood and
body uids. Next, we cut the tissues in half and treated the
tissue surfaces or cut edges with biosealants. We stabilized
them for 5 min at 37 °C within an incubator and lied them
to conrm the tissue adhesion of the biosealant. e animal
study was performed according to a protocol approved by the
Institutional Animal Care and Use Committee (IACUC) of
the Incheon National University (INU-ANIM-2022-02).
Cytocompatibility of biosealant
We evaluated the cytocompatibility of polymers and bioseal-
ants using HDFs. All reagents and solutions used in this experi-
ment were sterilized by UV irradiation for 15 min or syringe
ltration with a pore size of 0.2 μm. To assess the cytotoxicity
of GtnSH and GtnMI, the WST-1 assay (Roche) was performed
according to the manufacturer’s instructions and our previous
reports [29,30]. Briey, we seeded 1.0 × 104 cells in a 96-well
plate (SPL Life Sciences, Korea) with 200 μl of DMEM-HG
supplemented with 10% (v/v) NBCS and 1% (v/v) P/S at 37 °C
in 5% CO2. We prepared different concentrations (0.5 to
2.0 mg/ml) of GtnSH and GtnMI media using DMEM- HG.
Aer 24 h of cell seeding, we treated polymer media to HDFs
and cultured for 24 h. Next, we treated 10% WST-1 solution
to HDFs and incubated for 20 min. e absorbance of the
suspension was measured at 450 nm using a microplate reader.
e cell viability of the polymers was calculated as a percentage
of the control cells (untreated cells, TCPS). To determine the
cytotoxicity of the biosealant, we performed a WST-1 assay as
described above. In brief, HDFs (1.0 × 104 cells/well) were
seeded in a 96-well plate (SPL Life Sciences, Korea) with 100 μl
of DMEM-HG supplemented with 10% (v/v) NBCS and 1%
(v/v) P/S at 37 °C in 5% CO
2
. Next, we prepared the biosealant
eluates by incubating 40 μl of biosealant droplets with 193 μl
of media for 24 h. Next, we treated 100 μl of serially diluted
eluates to HDFs and incubated them for 24 h. en, HDFs
were incubated with 10% WST-1 solution for 20 min, and the
absorbance was measured at 450 nm. e cell viability of the
biosealant was calculated as a percentage of the control cells
(TCPS). e quantitative analysis of cell viability was analyzed
using a Live/Dead assay kit (Invitrogen). Aer 24 h of bioseal-
ant eluate treatment, we treated 193 μl of 2 μM calcein-AM
(the acetomethoxy derivative of calcein) and 4 μM ethidium
homodimer-1 (EthD-1) mixtures for 15 min. e cellular mor-
phology was qualitatively conrmed using a Nikon Eclipse
Ti-2 uorescence microscope (Nikon, Japan) at ×40 (low mag-
nication) and ×100 (high magnication).

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 6
In vivo biodegradability and tissue compatibility of
biosealant
To evaluate the biodegradability of the biosealant, we subcuta-
neously injected the biosealant solution into mice (6-week-old
female C57BL/6N mice) using a dual syringe kit in vivo. All
polymers, reagents, and solutions were sterilized by UV irradia-
tion for 15 min or ltered through a syringe lter (pore size:
0.2 μm). Aer a week of stabilization, the mice were anesthe-
tized with isourane in balanced oxygen. Next, we shaved dor-
sal hair and cleaned their skin with 70% ethanol and povidone.
We prepared 200 μl of biosealant solution and subcutaneously
injected it into mice (6-week-old female C57BL/6N mice) using
a dual syringe kit in vivo. For 8 weeks, we monitored the change
in biosealant volume and harvested the biosealant with the
surrounding tissues from the mice. e tissue compatibility of
the biosealant was evaluated using subcutaneous implantation
and H&E staining. Aer 8 weeks of subcutaneous injection, we
harvested various organs and xed them using 10% formalin.
e xed organs were dehydrated in graded ethanol (80% to
100%), embedded in paran, and serially sectioned using a
microtome (4 μm). e tissue slides were stained with H&E
and observed using a light microscope (Leica DM1000, Leica,
Germany) at ×200 magnication. e animal study was per-
formed according to a protocol approved by the IACUC of the
Incheon National University (INU-ANIM-2022-02).
Acute MI model of rats and transplantation
of S_3DP in MI model invivo
All animal experiments and surgical procedures were approved
by the IACUC of the Samsung Medical Center (SMCIACUC2021-
03-23-001). Male Sprague–Dawley rats (8 to 10 weeks old, 200
to 250 g) were anesthetized with isourane gas, endotracheally
intubated, and ventilated with air using a small-animal ventila-
tor (Harvard Apparatus, Hopkinton, MA, USA). e S_3DP
cultured in a 96-well plate was prepared in a 5% CO
2
incubator
at 37 °C during rat MI modeling. Aer opening the chest, MI
was induced by permanent ligation of the le anterior descend-
ing coronary artery (LAD) using 6-0 polypropylene. Ischemia
of the anterior wall of the le ventricle (LV) was conrmed by
the myocardium turning red to pale pink, and the animals were
randomly divided into 5 groups (n = 5 per group). e sample
codes and compositions are described in Table 2. Before trans-
planting S_3DP into the heart of the rat, pericardial uid was
removed using sterilized gauze, and the medium in the well
containing S_3DP was carefully removed by pipetting to pre-
vent the loss of spheroids from the S_3DP. en, S_3DP was
located on the ischemic heart surface and immobilized using
the biosealant. e biosealant was applied along the outside of
the 3D patch and allowed to stabilize for 3 min. Aer conrm-
ing that the biosealant and the 3D patch were xed to the sur-
face of the heart, the chest cavity was closed using 4-0 surgit
sutures. e chest muscles were then closed using 6-0 poly-
propylene, and the skin was sutured using 6-0 black silk
sutures.
Echocardiography
e rats were anesthetized aer 28 days of MI and S_3DP trans-
plantation, and echocardiography was performed for cardiac
function evaluation using a VisualS onics Ve vo 2100 (VisualSonics
Inc., Toronto, ON, Canada). All echocardiography measure-
ments were performed by a single-blind investigator. e le
ventricular internal dimension at end-diastole (LVIDd) and le
ventricular internal dimension at end-systole (LVIDs) were mea-
sured using the 2-dimensional (2D) targeted M-mode to cal-
culate the le ventricular ejection fraction (LVEF) and le
ventricular fractional shortening (LVFS). LVEF and LVFS were
calculated using the following equations [4,11,12]:
Histological analysis
Aer echocardiography, whole hearts were harvested from the
sacriced rats and xed with 10% formalin. e hearts were
sliced into 3 sections along the transverse axis below the ligature
to the apex at 2-mm intervals. e samples were dehydrated in
a series of ethanol, embedded in paran, and cut into 4-μm-thick
sections. e sections were deparanized in xylene, ethanol, and
distilled water and stained with H&E and Masson’s trichrome to
visualize cardiac brosis. Stained slides were scanned using the
Aperio ScanScope AT Slide Scanner (Leica Biosystems Inc.,
Bualo Grove, IL, USA) at ×200 magnication, and all images
were captured using ImageScope soware (Leica Biosystems
Inc.). e infarction size and brosis area of the LV were deter-
mined using the following formula [13]:
e LV wall thickness was measured in the infarcted and
border zones. Each zone was divided into 6 equal segments,
and LV wall thickness was calculated by averaging the thick-
nesses of the 6 segments [6]. All histological analyses were
quantied using ImageJ soware.
Statistical analysis
e GraphPad Prism 5 and 7 soware (GraphPad Soware, San
Diego, CA, USA) was used for all statistical analyses. Statistical
analysis was performed using a t test and one-way analysis of
variance (ANOVA) with the Bonferroni test, and all quantitative
data were presented as mean ± standard deviation (SD). Statistical
signicance was set at P < 0.05.
Results and Discussion
Fabrication of open and open/closed pocket patches
In our previous study, we synthesized TPU-CEC363 using PEG,
ε-caprolactone, and iron (III) acetylacetonate to develop a
(1)
LVEF
=
[(
LVIDd
3
−LVIDs
3)
∕LVIDd
3
×100%
]
(2)
LVFS
=
[
(LVIDd −LVIDs)∕LVIDd ×100%
]
(3)
Infarction size =
[(
epicardial circumference of the total LV area
+endocardial circumference of the total LV area)∕
(
epicardial circumference of the total LV area
+endocardial circumference of the total LV area)×100%
]
(4)

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 7
3D-printable patch with dual pockets [28]. TPU-CEC363 has
elasticity, exibility, thermal property, water uptake, and bio-
degradability, enabling the patch to withstand repeated contrac-
tions and relaxations of the heart. We punched a patch with a
diameter of 4 mm to obtain a uniform shape (Fig. 2A). In the
swelling test, the 3D patches swelled owing to the hydrophilic
properties of the TPU-CEC363. We conrmed that the diam-
eter of the patches increased slightly from 4 to 5 mm aer swell-
ing (Fig. 2B). We randomly measured the side lengths of the
open and closed pockets to conrm dierences in pocket size
aer swelling. As a result, the side length of the pockets was
similarly increased by approximately 70 μm aer swelling, and
there was no signicant dierence between open and closed
designs (Fig. 2C). ese results indicate that we successfully
fabricated the 3D-printed patch and constructed uniform dual
pocket structures under moist conditions. Furthermore, the
3D-printed patches provided a uniform and sucient cavity
aer swelling for spheroid formation.
Fabrication and characterization of hADSC
spheroids in open and open/closed pocket patches
To fabricate the stem cell spheroid within the 3D-printed patches,
we optimized the number of hADSCs with the open/closed
pocket patch. We loaded suspensions from 2 × 105 to 5 × 105
cells onto the 3D-printed patches and centrifuged them to enter
the cells within pockets (Fig. 3A and Fig. S1A). Aer 24 h of cell
seeding, we conrmed the morphology of the formed spheroids
according to the cell number (Fig. S1A).
When 2 × 105 cells were loaded, insucient cells were
loaded in the closed pockets, forming small-sized spheroids
(Fig. S1A and B). At 3 × 105 cells, we observed that spheroids
with uniform morphology were formed in the pockets (Fig.
Fig.2.Characteristics and preparation of the 3D patches for hADSC seeding. (A) The photograph of the 25 × 25 mm 3D patch and an illustration of 3D patch preparation for
the experiments. (B and C) Swelling characteristic of the 3D-printed patch. The photographs, micrographs, and graphs of the open pocket patch and open/closed pocket
patch before and after swelling. The result in (C) is shown as the average values ± SD (n = 24). Scale bar: 200 μm.

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 8
S1A and Fig. 3B). e size of the spheroids did not signicantly
dier depending on the pocket type (Fig. S1B and Fig. 3C).
However, above 3 × 105 cells, the restricted entrance size of the
closed pockets inhibited the homogeneous distribution of cells
in both the open and closed pockets. As a result, the size of the
spheroids in the open pockets gradually increased as the num-
ber of cells increased (Fig. S1A and B). e focus is on uni-
formly loading the cells into the fabricated 3D patches regardless
of pocket type.
Next, we evaluated the viability of spheroids and their expres-
sion of angiogenic growth factors depending on cell numbers.
e gene expression of BCL-2, an anti-apoptotic factor, was
highest in spheroids formed with 3 × 105 cells (Fig. S1C), and
the gene levels of apoptotic factors (BAX and Caspase-3) showed
relatively high expression in spheroids formed with 2 × 105 cells
(Fig. S1D). Furthermore, the gene expression of VEGF was not
signicantly dierent depending on the cell number, but the
gene expression patterns of HGF and broblast growth factor
(FGF)-2 were dierent (Fig. S1E). Nevertheless, the gene expres-
sions of spheroids formed by 3 × 10
5
cells were relatively higher
than those of other groups for both factors. erefore, we chose
the 3 × 105 cells/patch as the appropriate cell number for the
open/closed pocket patch and evaluated the treatment ecacy
through comparison with 2D cells.
Fig.3. Morphology and viability of hADSC spheroids in the open pocket and open/closed pocket patches. (A) Schematic illustration of the hADSC spheroids fabrication.
(B) Representative optical images of hADSCs within the 3D patches and spheroids after 24 h. Scale bar: 200 μm. (C) The average diameter of spheroids formed in the 3D
patches. The result in (C) is shown as the average values ± SD (n = 16). (D) Western blot analysis and quantification data comparing E-cadherin between 2D cells and spheroid
within 3D patches. The gene expression of (E) apoptotic factors (BAX, BAK, Caspase-3, and Caspase-9) and (F) anti-apoptotic factors (BCL-2 and BCL-xL). (G) Western blot
analysis of the representative apoptotic factor (BAX) and anti-apoptotic factor (BCL-2). The results in (D) and (G) are shown as the average values ± SD (n = 3). * indicates
a statistical significance compared to the 2D group (*P < 0.05 and ***P < 0.005). The results in (E) and (F) are shown as the average values ± SD (n = 4). * indicates a
statistical significance compared to the 2D group (*P < 0.05 and **P < 0.01).

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 9
During spheroid formation, it is well known that cell-to-cell
interactions increase with the accumulation of cadherin protein
on the cell membrane. In particular, E-cadherin plays a key role
in spheroid formation, and cadherin-to-cadherin binding com-
pacts cell aggregation to form spheroids [31–33]. erefore, we
evaluated the eect of the pocket system on spheroid formation
by analyzing E-cadherin expression. e protein expression of
E-cadherin was increased in the spheroid-laden open pocket
patch (S_OP) and spheroid-laden open/closed pocket patch
(S_OPCL) groups compared to that in the 2D group because
the pocket system structurally induced the aggregation of
hADSCs, increasing cell-to-cell interactions (Fig. 3D).
In fabricating a spheroid, the size of a spheroid is crucial in
determining its viability because it impacts oxygen and nutrient
supply within the spheroid. If the spheroid size is >500 μm, the
inner core accumulates the toxic metabolic waste and forms a
necrotic core [31,33]. erefore, we examined spheroid viability
using qRT-PCR and Western blotting. Compared to the 2D
group, the gene expression of apoptotic factors, such as BAX,
Bcl-2 homologous antagonist killer (BAK), and Caspase-3,
tended to decrease in the S_OP and S_OPCL groups. Moreover,
the expression of Caspase-9 was signicantly lower in these
groups (Fig. 3E). Cell apoptosis is initiated by sequential activa
-
tion of several caspases. Among various caspases, Caspase-9,
an initiator caspase, is activated by forming an apoptosome
with cytochrome c and apoptotic protease activating factor-1
(Apaf-1). e activated Caspase-9 stimulates Caspase-3, an
eector caspase, to initiate cell apoptosis [34,35]. In contrast,
the gene expression of anti-apoptotic markers, such as BCL-2
and B-cell lymphoma extra-large (BCL-xL), in the S_OP and
S_OPCL groups tended to be higher than those in 2D cells (Fig.
3F). Furthermore, regarding protein expression, we conrmed
that BCL-2 expression was signicantly increased in the S_OP
and S_OPCL groups, whereas BAX expression was not signi-
cantly dierent from that in 2D cells (Fig. 3G). Consequently,
our 3D patches with a dual module pocket demonstrated their
potential as a culture system for fabricating stem cell spheroids
without cytotoxic issues.
Paracrine angiogenic effect of S_3DP
Stem cell spheroids have been reported to exhibit superior thera-
peutic eects compared with conventional monolayer stem cells
because of their upregulated paracrine secretion [11,32]. Previous
studies reported that the expression of hypoxia-inducible factor
1-alpha (HIF-1α) increases in the mildly hypoxic inner zone of
spheroids [31,32]. HIF-1α acts as a transcription factor and regu-
lates various target genes related to angiogenesis [32]. With this
rationale in mind, we compared the angiogenesis-related factor
expression and paracrine eects between our system and 2D
cultured stem cells in vitro.
We rst evaluated gene expression of HIF-1α, resulting in
the S_OP and S_OPCL groups upregulated gene expression of
HIF-1α compared to the 2D group (Fig. 4A). Next, we analyzed
the expression of VEGF, FGF-2, and HGF as representative
pro-angiogenic markers [9,31]. e S_OP and S_OPCL groups
showed increased gene expression of each factor compared to
that in the 2D group (Fig. 4A). Notably, the gene expression
of VEGF was significantly increased by more than 3 times.
Following gene expression analysis, we performed Western blot-
ting to assess the protein expression of angiogenesis-related
growth factors, such as HGF and VEGF (Fig. 4B). e levels
of both growth factors were higher in the S_OP and S_OPCL
groups than those in the 2D group. Based on these results, our
spheroid-laden patch system promotes the expression of growth
factors related to angiogenesis.
hADSCs secrete a wide range of factors that promote angio-
genesis, reduce apoptosis, and regulate inammatory responses.
e paracrine eects of hADSCs have been validated in numerous
studies of MI treatment [5,31]. Consequently, the secretome
released from hADSCs is considered a promising source for
cytokine-based therapy, and its secretion and angiogenic poten-
tial have been reported to be promoted under 3D culture condi-
tions [15,31,33]. To investigate whether our 3D culture system
could induce a stronger paracrine eect by releasing increased
angiogenesis-related factors compared to 2D culture, we ana-
lyzed the collected CM using an angiogenesis array kit. e
expression of various angiogenic factors was increased in the
CM obtained from the S_OP and S_OPCL groups compared
to that in the CM from the 2D group (Fig. 4C). Notably, pro-
angiogenic factors such as angiogenin, endothelin-1, FGF-1,
HGF, insulin-like growth factor-binding protein (IGFBP)-3,
pentraxin-3, persephin, and urokinase-type plasminogen acti-
vator (uPA) were modestly increased in the S_OP and S_OPCL
groups compared to those in the 2D group (Fig. 4D). Interestingly,
endocrine gland-derived vascular endothelial growth factor
(EG-VEGF), IGFBP-2, and interleukin-1 beta (IL-1β) were
detected only in the S_OP and S_OPCL groups, but not expressed
in the 2D group (Fig. 4E). ese proteins have the potential
angiogenic capacity, and their expression is known to be regu-
lated by HIF-1α [36–39]. In particular, EG-VEGF and IL-1β
are inuenced by the expression of HIF-1α because they have
HIF-1α binding sites in their promoter regions [40,41]. Further-
more, IGFBP-2 acts as a transcriptional enhancer that promotes
the activity of the VEGF gene promoter through nuclear trans-
location, leading to enhanced VEGF expression [36,37]. Conse-
quently, we speculate that the increased expression of HIF-1α
within the hypoxic core of the spheroids may stimulate the
expression of EG-VEGF, IGFBP-2, and IL-1β, resulting in the
paracrine angiogenic eect of S_3DP.
Based on these, the cytokines expressed only in the CM of
the S_OP and S_OPCL groups showed more marked therapeu-
tic potential than conventional 2D cell transplantation.
Consequently, we demonstrated that our system could improve
the secretion of various angiogenic factors and their paracrine
eects through CM.
Fabrication and characterization of tissue adhesive
biosealant
To immobilize S_3DP on the heart surface, we developed a
gelatin-based biosealant via the dual crosslinking reactions
involving the thiol-ene reaction and disulde bond formation
(Fig. 5A and B). In this system, the biosealant is rst crosslinked
via a thiol-ene reaction between the functional groups, thiol
and maleimide groups, in the polymers. We used CaO2 as an
additional crosslinker to reinforce the mechanical stiness
of the biosealant. CaO
2
and its derivative, calcium hydroxide
(Ca(OH)2), have been extensively used in food, cosmetic, and
biomedical applications [29]. Notably, it is well known that
CaO2 produced hydrogen peroxide (H2O2) during its decom-
position in aqueous media, inducing disulde bond formation
via H
2
O
2
-mediated thiol oxidation. Furthermore, these chemi
-
cal reactions of the hydrogel enabled attachment to the heart
via the interaction of maleimide and thiol groups with the pri-
mary amine and thiol groups on the heart surface. erefore,

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 10
GtnSH and GtnMI were synthesized as polymers of biosealant
using the EDC/NHS chemistry (Fig. S2A). Gtn was selected as
the polymer backbone owing to its biocompatibility, biodegrad-
ability, bioactivity, and easy modication [29]. Next, we char-
acterized the chemical structures of GtnSH and GtnMI by
1
H-NMR spectroscopy. As a result, GtnSH exhibited new peaks
at 2.8 and 3.5 ppm, indicating the conjugation of cystamine
to the carboxyl group of Gtn [42] (Fig. S2B). For GtnMI,
we confirmed a sharp peak appearance at 6.8 ppm and a
decreased peak area at 3.0 ppm, which indicates the intro-
duction of the maleimide group to a primary amine group
of Gtn [30,43] (Fig. S2C). Next, we determined the functional
groups of GtnSH and GtnMI using Ellman’s assay. GtnSH
and GtnMI revealed 158.3 ± 42.7 and 87.8 ± 13.9 μmol/g of
polymers, respectively. Based on these results, we success-
fully synthesized polymers for fabricating tissue adhesive
biosealant.
e viscoelasticity of the biosealant is crucial for patch xa-
tion and sealing, particularly in resisting external forces, such
as the heartbeat. Accordingly, we fabricated tissue adhesive
biosealants with various CaO2 concentrations and measured
their G′ using a rheometer. As increasing CaO2 concentration,
Fig.4.Paracrine effect of angiogenic factors of hADSC spheroids in the open pocket and open/closed pocket patches. The relative expression of HIF-1α and angiogenic growth
factors (VEGF, FGF-2, and HGF) of the 2D group and experimental groups was evaluated using (A) qRT-PCR and (B) Western blot analysis. (C) Representative images of the
human angiogenesis antibody array data of the 2D group and the experimental groups. (D) The relative expression of angiogenesis-related proteins in CM. (E) The relative
expression of angiogenic proteins that were only expressed in the experimental groups, not in the 2D group. The results in (A) and (B) are shown as the average values ± SD (n =
5 and 3). * indicates a statistical significance compared to each group (*P < 0.05 and ***P < 0.005). The results in (D) and (E) are shown as the average values ± SD (n = 2).

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 11
the G′ of hydrogels was reinforced (0 wt%, 230 Pa; 0.01 wt%,
360 Pa; 0.05 wt%, 420 Pa; 0.1 wt%, 920 Pa) (Fig. 5C). Typically,
the stiness of the rat myocardium is known to range from 0.1
to 140 kPa [44]. Additionally, it has been reported that the opti-
mal G′ value of the hydrogel matrix for cardiac repair ranges
from 380 to 600 Pa, representing a relatively so matrix. is
so matrix facilitates the transmission of mechanical signals and
the coordination of myocardial tissue, improving the reconstruc-
tion of cardiac function [45]. Consequently, we controlled the
stiness of our biosealant using CaO
2
concentrations, providing
a suitable matrix in the dynamic cardiac environment for eec-
tive cardiac repair.
e tissue adhesive force of the biosealant should withstand
external forces such as heartbeat and blood pressure and stably
immobilize the patch in place. erefore, we measured the tis-
sue adhesive force of our biosealant at various CaO
2
concentra-
tions using decellularized porcine skin and a UTM (Fig. 5D).
We observed that CaO2 addition reinforced the tissue adhesive
strength of the biosealant compared to that of the control (CaO
2
0 wt%, 11.3 kPa; 0.01 wt%, 15.7 kPa; 0.05 wt%, 15.4 kPa; 0.1 wt%,
Fig.5.Fabrication and characterization of biosealant. (A) Schematic illustration of biosealant for fixing S_3DP. (B) Mechanism of biosealant formation and tissue adhesion.
(C) Mechanical property of biosealant with various CaO2 concentrations. The result in (C) is shown as the average values ± SD (n = 3). * indicates a statistical significance
compared to CaO2 0 wt% (*P < 0.05). (D) Schematic illustration of tissue adhesiveness. Tissue adhesiveness of biosealant on (E) decellularized porcine skin and (F) various
organ surfaces. The result in (E) is shown as the average values ± SD (n = 6). * indicates a statistical significance compared to CaO2 0 wt% (*P < 0.05).

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 12
14.0 kPa) (Fig. 5E). ese tissue adhesive strengths were higher
than commercially available brin glue [46,47] (5 to 10 kPa).
Our biosealant could strongly attach to tissue surfaces through
a thiol-ene reaction and disulde bond formation, in contrast
to brin glue, which relies on hydrogen bond formation during
coagulation [48] (Fig. 5B). Interestingly, we conrmed that the
increasing tissue adhesive strength was saturated between
CaO2 0.01 and 0.05 wt% and decreased at CaO2 0.1 wt%.
e number of tissue adhesive functional groups, such as thiol
groups, and gelation time were rapidly decreased when the
CaO2 concentration was increased, resulting in insucient tis-
sue adhesive groups and time to interact with the tissue surface.
Accordingly, we selected CaO2 0.1 wt% as the optimal com-
position of biosealant for immobilizing a patch in vivo,
homogeneously.
Next, we determined the feasibility of biosealant on various
organ surfaces (Fig. 5F). Interestingly, our sealant bonded
be tween tissue surfaces and cut areas of various organs. e
tissue adhesion of our sealant targets the thiol and primary amine
groups present on most tissue surfaces. is result suggests that
our biosealant can be used as an adhesive to x patches on vari-
ous tissue surfaces.
Consequently, we developed gelatin-based tissue adhesive
biosealants with tunable elasticity and stronger tissue adhesive-
ness than commercial brin glue. is biosealant has great
potential as an advanced sutureless technique to immobilize
implantable devices in vivo.
Biocompatibility and biodegradability of biosealants
e biocompatibility and biodegradability of biosealant are
essential properties for successful clinical applications because
they ensure the safety of materials and their by-products and
suitable maintenance of the sealant at a rate compatible with
tissue healing [22]. To utilize our sealant for immobilizing the
S_3DP in vivo, we rst evaluated the cytocompatibility of the
polymers using HDFs and the WST-1 assay in vitro. GtnSH and
GtnMI showed high cell viability above 94% in TCPS (Fig. 6A).
is result indicates that our biosealant may be cytocompatible
Fig.6.Biocompatible and biodegradable biosealant. Cytocompatibility of (A) polymer and (B) biosealant on HDFs. (C) Fluorescence images of the HDFs after a day of hydrogel
eluate treatment. (D) Biodegradability and (E) tissue compatibility of biosealant invivo. The results in (A) and (B) are shown as the average values ± SD (n = 6). Scale bar: 100 μm.

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 13
with host cells. Additionally, the biosealant exhibited excellent
cell viability above 95.2% of TCPS and the predominant viable
cell population (Fig. 6B and C).
We next subcutaneously injected the biosealant in mice to
evaluate its biodegradability and tissue compatibility in vivo.
Figure 6D shows that the sealant was maintained for 8 weeks and
was sustainably degraded in vivo. is suggests that our sealant
is biodegradable and can be maintained at a specic volume
within the body for 8 weeks in vivo. Furthermore, we evaluated
the tissue compatibility of the biosealant using various major
organs harvested 8 weeks aer subcutaneous injection. Compared
with normal organs, our biosealant and their by-products revealed
no pathological changes in the major organs in vivo (Fig. 6E).
Consequently, our biosealant proved to be biocompatible
and biodegradable with the potential to x patches in vivo.
Based on these results, we investigated the therapeutic eects
of the S_3DP on cardiac infarction.
The therapeutic effect of S_3DP on
cardiac infarction
We investigated the therapeutic ecacy of S_3DP using a rat
MI model. To fabricate the MI model, we tied the LAD to induce
permanent vascular occlusion and observed changes in myo-
cardial color [7]. Aer fabricating the MI model, the S_3DP
was immediately immobilized on the heart surface using our
biosealant and stabilized for 3 min. As a result, we successfully
attached the patch to the heart surface without slippage under
a dynamic heartbeat, and it was stably maintained for 3 days
(Fig. 7A). An additional movie le shows this in more detail
(Movie S1). To compare the eect of patches or spheroids on
MI treatment, we selected the sealant-only (SO) and open/
closed pocket patch-only (PO) groups for comparison. Aer
4 weeks post-transplantation, we examined LVEF and LVFS,
which are the most commonly used parameters to quantify
heart function using echocardiography [3,11,49] (Fig. 7B). We
rst measured LVIDd and LVIDs to calculate the LVEF and
LVFS using the photographed ultrasound results [49]. LVIDd
and LVIDs in the controls (MI, SO, and PO) and S_OP group
were notably higher than those in the normal group. However,
there was no signicant dierence in the results of LVIDd and
LVIDs between the S_OPCL and normal group. Moreover, the
S_OPCL group showed signicantly decreased LVIDs com-
pared to the control groups (Fig. 7C and D). LVIDs is the result
of myocardial contractility, meaning that the narrower the
Fig.7.Transplantation of S_3DP in the rat MI model and improved cardiac function. (A) Representative photographs of the heart after acute MI modeling and transplantation
of S_3DP with biosealants. (B) Representative echocardiographic images of each group 4 weeks after S_3DP transplantation. The (C) LVIDd, (D) LVIDs, (E) LVEF, and (F) LVFS
were assessed using 2D echocardiography. The results in (C) to (F) are shown as the average values ± SD (n = 5). * and # indicate a significant difference compared to the
normal group and each group (*P and #P < 0.05, **P and ##P < 0.005).

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 14
LVIDs, the better the heart function. is result demonstrated
that the delivery of hADSC spheroids eectively improved myo-
cardial function through their paracrine eect, and open and
closed pockets were more suitable than open pockets.
LVEF and LVFS were calculated using established formulas,
with the internal diameter values measured during systole and
diastole. LVEF is the fraction of blood ejected by the LV during
the cardiac cycle, and LVFS indicates the percentage change in
the LV internal dimensions between systole and diastole [3].
All groups showed lower heart function than the normal group
in terms of both LVEF and LVFS. However, we conrmed that
the S_3DP groups (S_OP and S_OPCL) had signicantly
improved cardiac functions with increased LVEF and LVFS
compared to the control groups (Fig. 7E and F). Consequently,
we successfully immobilized our S_3DP system using a bioseal-
ant without physical damage, which occurs by suturing or injec-
tion. Our system was stably maintained for 4 weeks even in
the presence of pericardial uid. Moreover, we demonstrated
its therapeutic eects in the restoration of cardiac function.
Notably, the combination of open and closed pockets was the
most eective through the paracrine eect and protection of
the spheroids against harsh infarcted conditions.
Reduced cardiac fibrosis of S_3DP on
cardiac infarction
Following MI, cardiomyocytes undergo cell death while cardiac
broblasts are activated, replacing the damaged myocardium
with collagen-based scar tissue [11,50]. However, this brotic
tissue cannot perform normal functions, such as contraction
Table1. Sample codes and experimental condition of S_3DP
Sample
codes
Pocket
type
Patch swell-
ing time
Cell number
(cells/
patch)
Incuba-
tion time
2D - At least 5
min
3 × 10524 h
S_OP Open
S_OPCL Open/
closed
Fig.8.Therapeutic effect of the S_3DP on reducing cardiac fibrosis. Schematic illustration showing (A) the location of areas 1 to 3 in the heart and (B) the cross-section of the
heart. (C) Representative images of the whole heart and each area. Representative (D) H&E and (E) Masson’s trichrome stained images of areas 1 to 3 after 4 weeks of surgery
and treatment. Scale bar: 4 mm. Quantification of (F) the wall thickness, (G) infarction size, and (H) fibrosis area of the LV in each area based on Masson’s trichrome staining.
The results in (F) to (H) are shown as the average values ± SD (n = 5). * indicates statistical significance compared to each group (*P < 0.05, **P < 0.01, ***P < 0.005).

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 15
and relaxation. us, reducing the brotic area in the LV is
important for maintaining or restoring cardiac function. MI is
accompanied by ECM degradation, LV wall thinning, and ven-
tricular enlargement [10,50]. Based on these, we histologically
analyzed LV wall thickness, infarction size, and brosis area.
We sectioned the whole heart into 3 areas at 2-mm intervals
from the LAD-tied part, which are indicated in areas 1 to 3
(Fig. 8A and C). We then divided the cross-sectional tissue of
the heart into 3 zones: an infarcted zone (IZ), a border zone
(BZ), and a remote zone (RZ). We focused our analysis on the
IZ and BZ [7] (Fig. 8B). To determine the brotic area of the
heart, we performed H&E and Masson’s trichrome staining
on sectioned tissues (Fig. 8D and E). e blue area indicates
brotic tissue (collagen), and the red area indicates the myo-
cardium (Fig. 8E).
During post-infarction repair, brotic tissue formation causes
progressive thinning of LV wall thickness [50]. erefore, we
rst measured the LV wall thickness according to the zone. In
area 1, the LV wall thicknesses in the S_3DP groups (S_OP and
S_OPCL) were signicantly greater than those in the control
groups (MI, SO, and PO) in both the BZ and IZ. In area 2, only
the MI and S_OPCL groups showed a signicant dierence in
the BZ. Interestingly, in the IZ of area 2, the LV wall thickness
in the S_OPCL group was considerably greater than those in the
control groups. However, in area 3, no signicant dierences
were observed between all groups (Fig. 8F). Based on these
results, we conrmed that the LV wall thickness was increased
by the spheroid treatment through S_3DP, except in area 3, and
the S_OPCL group was more eective than in the S_OP group.
Furthermore, we analyzed the infarct size and brotic area
of LV. e following formula: [(a + b)/(c + d) × 100%] was used
to calculate the infarction size by measuring the length of the
letter, as shown in Fig. 8B. ere were no dierences among
the control groups in any area. The infarction size in the
S_OPCL group was significantly reduced in areas 1 and 2
compared with those in the control groups, but the infarction
size in the S_OP group did not decrease in any area (Fig. 8G).
To determine the percentage of brosis, we measured the total
LV and brotic areas in Fig. 8E. ere were no signicant dif-
ferences between the control groups in any area. In all areas,
the brosis area of the S_OP group was decreased compared
to those of the control groups, although the dierences were
not statistically signicant. In contrast, the brosis area of the
S_OPCL group was notably reduced compared to those of all
groups in all areas and revealed only signicant dierences in
areas 1 and 2 (Fig. 8H).
In conclusion, we conrmed that S_3DP transplantation
eectively reduces brotic scar tissue and increases LV wall
thickness. Notably, the S_OPCL group showed greater poten-
tial therapeutic eects than the S_OP group, which is consis-
tent with the cardiac function results.
Conclusion
In this study, we developed S_3DP and biosealant as a new type
of stem cell spheroid therapy to treat MI. We designed S_3DP
with dual pore modules (open and closed pockets) to improve
the engrament rate and paracrine eect of the spheroids. e
3D patches provided uniform conditions for stem cell culture
and spheroid fabrication with high cell viability. Moreover,
these spheroids produced secretomes related to angiogenesis
and exhibited improved paracrine eects in vitro. Next, we fab-
ricated a tissue adhesive biosealant as a sutureless method for
immobilizing S_3DP to the heart tissue, maintaining its posi-
tion for 4 weeks in vivo. When we transplanted our system into
a rat MI model, S_3DP facilitated the restoration of cardiac
function and decreased the infarcted area. Notably, the combi-
nation of open and closed pocket, S_OPCL, was more eective
than the open pocket, S_OP, demonstrating that the protective
eect of the closed pocket acts synergistically with direct spher-
oid delivery and paracrine eects. Collectively, our system has
great potential as a new type of spheroid delivery system with
a sutureless approach for MI treatment and can overcome the
challenges of heart disease therapy.
Acknowledgments
Funding: is work was supported by the National Research
Foundation of Korea (NRF) funded by the Ministry of Science
and ICT [NRF-2018M3A9E2023255], the Alchemist Project
of the Korea Evaluation Institute of Industrial Technology
[KEIT 20018560 and NTIS 1415184668] funded by the Ministry
of Trade, Industry & Energy (MOTIE, Korea), the Korea
Fund for Regenerative Medicine [KFRM21A0102L1-11],
and the NRF grant funded by the Korea government [NRF-
2021R1A4A5032185].
Author contributions: H.R.J. and J.I.K. conducted the experi-
ments, analyzed data, designed the gures, and wrote the original
manuscript. S.H.B. conceptualized the project and contributed
to interpretation of data. K.M.P. and D.-I.K. supervised and
conceptualized the project, edited the manuscript, and con-
tributed to interpretation of data. All authors have read and
approved the manuscript.
Competing interests: e authors declare that they have no
competing interests.
Data Availability
The data are available from the authors upon a reasonable
request.
Table2. Sample codes and compositions
Sample
codes
MI
induction
Treatment
of bioseal-
ant
3D patch
trans-
plantation
(pocket
type) Spheroid
Normal X X X X
MI O X X X
SO O O X X
PO O O O (Open/
closed
pocket)
X
S_OP O O O (Open
pocket)
O
S_OPCL O O O (Open/
closed
pocket)
O

Jeon et al. 2024 | https://doi.org/10.34133/bmr.0007 16
Supplementary Materials
Tables S1 and S2
Figs. S1 and S2
Movie S1
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