Transplantation of stem cell spheroid-laden three-dimensional patches with bioadhesives for the treatment of myocardial infarction.

doi:10.21203/rs.3.rs-3386370/v1

Download PDF

Research Article

Transplantation of stem cell spheroid-laden three-dimensional patches with bioadhesives for the treatment of myocardial infarction.

https://doi.org/10.21203/rs.3.rs-3386370/v1

This work is licensed under a CC BY 4.0 License

You are reading this latest preprint version

Background

Myocardial infarction (MI) is treated with stem cell transplantation using various methods. However, current methods, such as stem cell/spheroids injection, cell sheets, and cardiac patches have some limitations, such as low stem cell engraftment and poor therapeutic effects. Furthermore, these methods cause secondary damage due to injection and suturing to immobilize them in the heart, inducing side effects.

Methods

We fabricated human adipose-derived stem cell spheroids-laden three-dimensional (3D)-printed patches. The morphology, viability, and paracrine angiogenic effect of spheroids formed within 3D patches were analyzed in in vitro experiments. We synthesized thiolated gelatin and maleimide-conjugated gelatin as the polymers and fabricated a tissue adhesive biosealant using the polymers. The biocompatibility and biodegradability of the biosealant were evaluated using human dermal fibroblasts in vitro and the mouse subcutaneous models in vivo. In addition, the therapeutic effects of stem cell spheroid-laden 3D patches (S_3DP) with biosealant were evaluated using a rat MI model in vivo.

Results

The spheroids showed increased viability and expression of angiogenic factors compared to two-dimensional cultured cells. Our gelatin-based tissue adhesive biosealants were rapidly formed via a thiol-ene reaction and disulfide bond formation and revealed 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 fibrosis of the heart.

Conclusion

We developed S_3DP with gelatin-based tissue adhesive 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. In summary, S_3DP and biosealant have excellent potential as advanced stem cell therapies with a sutureless approach to MI treatment.

Stem cell

spheroid

polymeric hydrogel

myocardial infarction

Cardiovascular diseases, responsible for 17.7 million deaths annually, have emerged as the primary global cause of mortality. This number is expected to increase to 23.6 million by 2030, surpassing that of cancer and other ailments [1, 2]. Among cardiovascular diseases, myocardial infarction (MI) is a representative cardiac ischemic disease caused by coronary artery occlusion [3–7]. MI induces a lack of oxygen and nutrient supply, causing myocardial necrosis, loss of cardiac function, and eventual heart failure [3–5, 8–10]. Furthermore, unlike other organs, the regenerative ability of the damaged heart is limited because adult cardiomyocytes have low proliferative and self-renewal capacities [3, 8, 11, 12]. 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 [5, 11, 13–17]. 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 [7, 18, 19]. Therefore, it is necessary to develop treatments that restore the underlying damaged tissue 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 multipotency [12]. Mesenchymal stem cells (MSC) have been extensively used as a promising therapeutic agent for various diseases, extending beyond MI [20]. Although stem cells have been transplanted to treat MI, the low retention of stem cell delivery (5–10%) in the myocardium limits their applications [3, 7, 12, 21–30]. Therefore, various technologies, including spheroid formation, injectable biomaterials, cell sheets, and cardiac patches, have emerged as more effective approaches to deliver stem cells with improved retention rates to improve the restoration of infarcted myocardial functions [5–7, 9, 14, 25, 31–42].

Stem cell spheroids are dense three-dimensional (3D) cell clusters formed by cell-to-cell aggregation. Fabricating spheroids promotes stem cell survival and retention, and physiological and metabolic functions compared to single stem cells [43, 44]. Injectable biomaterials, such as polymeric hydrogels, have attracted considerable attention as delivery carriers for cells and spheroids. They can promote retention rates and biological functions by providing a 3D artificial extracellular matrix (ECM) that can recapitulate cell-to-matrix interactions and physically support cells and spheroids [22, 44–46]. 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 [45, 47]. Furthermore, physical injury caused by the injection can cause secondary damage, bleeding, and myocardial perforation, leading to possible adverse cardiac events [5, 7, 14, 25, 31–34, 38, 42, 48]. As with other strategies, cell sheets and cardiac patches have been widely utilized owing to their improved cell-to-cell communication and more powerful paracrine effects, which improve prognosis compared to stem cell solution therapy [3, 20, 49, 50]. 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, 35–37, 40, 41]. Therefore, it is necessary to develop advanced stem cell therapies to improve the stem cell engraftment rate and paracrine effects using a sutureless approach to maximize therapeutic efficacy.

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 adhesion [51–53]. Therefore, they have been used to prevent gas/liquid leakage and bleeding and immobilize implantable medical devices on target tissues [54]. Currently, various tissue adhesive biosealants, such as Tisseel®, Coseal®, Duraseal®, Progel™, and others, have been successfully approved by Food and Drug Administration (FDA) and are widely used in clinics [54]. However, these commercialized biosealants have relatively lower tissue adhesiveness when applied to myocardial tissue owing to the presence of the body or pericardial fluid and the dynamic pulsation of the heart [36, 55–58]. Therefore, there is a need to develop advanced functional biosealants with strong tissue adhesion. Toward this, biosealant to applicate for heart tissue have to meet some requirements, such as ⅰ) rapid and easy application for clinical compliance, ⅱ) biocompatibility of polymer and chemistry, ⅲ) mechanical similarity to surrounding tissue and elasticity to heartbeats, ⅳ) strong tissue adhesion on the heart surface under the pericardial fluid and the dynamic heartbeats, and ⅴ) biodegradability.

In this study, we developed two types of stem cell spheroid-laden 3D patches (S_3DP): i) open pockets to directly deliver stem cell spheroids with an increased paracrine effect, and ii) open/closed pockets to enhance the engraft rates of spheroids by protecting them from harsh microenvironments (Scheme 1a). The stem cell spheroids formed within these patches revealed high cell viability with improved bioactivities, such as the expression of genes and proteins and paracrine ability related to angiogenesis, in vitro. We also developed a tissue adhesive biosealant to immobilize functional S_3DP on the heart tissue using sutureless approaches. Finally, we transplanted S_3DP into a rat MI model using a biosealant and demonstrated its effect on cardiac function and fibrosis in vivo (Scheme 1b). The entire S_3DP preparation and transplantation process for MI treatment is depicted in Scheme 1.

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-buffered saline (PBS) was purchased from Biosesang (Seongnam, Korea). Human adipose-derived stem cell (hADSC) was obtained from Lonza (Walkersville, MD, USA). The 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 buffer were obtained from Bio-Rad (Hercules, CA, USA). 5% skim milk was purchased from BD Difco™ (Detroit, MD, USA). Antibodies of glyceraldehyde 3-phosphate dehydrogenase (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 factor (VEGF) antibody was purchased from Santa Cruz Biotechnology (Dallas, TX, USA). Goat anti-mouse IgG-HRP and goat anti-rabbit IgG-HRP antibodies were purchased from Bethyl Laboratories (Montgomery, TX, USA). The ECL reagent WESTSAVE UP was purchased from AbFrontier (Seoul, Korea). The X-ray films were obtained from AGFA HealthCare NV (Mortsel, Belgium). The Proteome Profiler Human Angiogenesis Array Kit was purchased from R&D Systems (Minneapolis, MN, USA). For in vivo animal study, 8–10-week-old male Sprague-Dawley (SD) rats were supplied by Orient Bio Inc. (Seongnam, Korea). 6 − 0 polypropylene, 4 − 0 surgifit, and 6 − 0 black silk sutures were obtained from Ailee (Busan, Korea). 10% formalin and paraffin were purchased from Sigma-Aldrich (St. Louis, MO, USA). Xylene was obtained from Daejung (Siheung, Korea). The 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 purification 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 dihydrochloride (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-buffered saline (DPBS) was obtained from Gibco (Grand Island, NY, USA). Dialysis membranes (molecular cut-off = 3.5 kDa) were supplied 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 Thermo Scientific (Rockford, IL, USA). For the in vitro cell study, human dermal fibroblast (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). The Live/Dead kit was obtained from Invitrogen (Grand Island, NY, USA). The 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). Isoflurane (Forane Solution) was purchased 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 bioscaffold (TPU-CEC363) with tunable biodegradability and elasticity for noninvasive stem cell-based therapy [59]. In this study, developed elastomer was used as a bioink to fabricate stem cell spheroid-encapsulating patches. Briefly, 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 copolymer, toluene, and hexamethylene diisocyanate. As a result, TPU-CEC363 containing 50% each of PEG and polycaprolactone (PCL) was prepared. The synthesized TPU-CEC363 was extruded to a diameter of 1.75 mm, and the 3D patch was 3D printed using FDM Vis Power Plus (Vision Technology Korea, Daejeon, Korea) [60].

Characterization of open and open/closed pocket patches

The 3D-printed patches with a size of 25 × 25 mm were supplied by the Korea Institute of Science and Technology and Vision Technology Korea. The 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). The swelling ability was performed with 1× PBS, and photographs of the pockets were captured using a light microscope (CKX41, Olympus) and Infinity analyze software (Lumenera Corporation, Ottawa, Canada). The pocket size was analyzed using ImageJ software (National Institutes of Health, Bethesda, MD, USA).

Preparation of 3D patch and spheroid formation in 3D patch

The 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 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 patch was swollen by adding 1× PBS to the wells of a 96-well plate. The 3D patch in PBS was centrifuged at 1,500 rpm for 5 min, and the PBS was replaced with serum-free DMEM. The 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. The medium was changed every 2 days, and hADSCs from passages 4 to 7 were used in the experiments. 50 µL of serum-free DMEM containing 3 × 10⁵ cells were dispensed over the 3D patch. The hADSCs entered the pockets of the 3D patch after centrifugation at 1,500 rpm for 5 min, and spheroids were formed in the open and open/closed pockets after 24 h.

Table 1

Sample codes and experimental condition of S_3DP.
Sample codes	Pocket type	Patch swelling time	Cell number (cells/patch)	Incubation time
2D	-	At least 5 min	3 × 10⁵	24 h
S_OP	Open
S_OPCL	Open/closed

Characterization of stem cell spheroid

We observed the spheroid morphology under a light microscope (CKX41, Olympus) after 24 h. Images were captured to confirm the diameter of the spheroids using a light microscope (CKX41, Olympus) and Infinity analyze software (Lumenera Corporation), and the diameter was analyzed using ImageJ software.

Quantitative reverse transcription-polymerase chain reaction (qRT-PCR)

Total RNA was extracted using 1 mL of TRIzol and 200 µL of chloroform. After 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. The samples were then dissolved in RNase-free water. To synthesize complementary DNA, 1 µg of total RNA and the AccuPower® RocketScript™ Cycle RT PreMix were used. The AccuPower® 2X Greenstar™ qPCR MasterMix and QuantStudio™ 6 Flex Real-Time PCR System (Applied Biosystems, Waltham, MA, USA) were used for qRT-PCR. The relative gene expression was analyzed using the 2^(-ΔΔCT) method and GAPDH served as the internal control.

Table 2

Sequences for primers used in qRT-PCR.
Species	Gene	Primer	Sequence (5’─3’)
Human	GAPDH	Forward	GTC GGA GTC AAC GGA TTT GG
	GAPDH	Reverse	GGG TGG AAT CAA TTG GAA CAT
	BAX	Forward	GCT ACA GGG TTT CAT CCA GGA TC
	BAX	Reverse	CCG TGT CCA CGT CAG CAA TC
	BAK	Forward	CTC AGA GTT CCA GAC CAT GTT G
	BAK	Reverse	CAT GCT GGT AGA CGT GTA GGG
	CASPASE-3	Forward	CCT GGT TAT TAT TCT TGG CGA AA
	CASPASE-3	Reverse	GCA CAA AGC GAC TGG ATG AA
	CASPASE-9	Forward	CTC AGA CCA GAG ATT CGC AAA C
	CASPASE-9	Reverse	GCA TTT CCC CTC AAA CTC TCA A
	BCL-2	Forward	TCC CTC GCT GCA CAA ATA CTC
	BCL-2	Reverse	ACG ACC CGA TGG CCA TAG A
	BCL-xL	Forward	AAC ATC CCA GCT TCA CAT AAC CCC
	BCL-xL	Reverse	GCG ACC CCA GTT TAC TCC ATC C
	VEGF	Forward	GAG GGC AGA ATC ATC ACG AAG T
	VEGF	Reverse	CAC CAG GGT CTC GAT TGG AT
	FGF-2	Forward	AGC GGC TGT ACT GCA AAA AC
	FGF-2	Reverse	GTA GCT TGA TGT GAG GGT CG
	HGF	Forward	TCA AAT GCC AGC CTT GGA ATT CC
	HGF	Reverse	TCA AGA GTG TAG CAC CAT GGC
	HIF-1α	Forward	CAG TTA CGT TCC TTC GAT CAG TTG
	HIF-1α	Reverse	TTT GAG GAC TTG CGC TTT CA

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 buffer containing β-mercaptoethanol, and samples were loaded onto a sodium dodecyl sulfate-polyacrylamide gel electrophoresis (SDS-PAGE) gel. After gel electrophoresis, proteins were transferred 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). The 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). The 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. After 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 film. To analyze protein expression, GAPDH and ImageJ software were used as a housekeeping control and an analysis tool.

Angiogenesis antibody array

To profile angiogenesis-related proteins in the conditioned medium (CM) of the spheroids, we used the Proteome Profiler Human Angiogenesis Array. After spheroid formation in the 3D patch, the cells were carefully washed with 1× PBS, and the spheroids in the 3D patch were incubated with fresh serum-free DMEM for 24 h to harvest the CM. The protein concentration of the harvested CM was measured by Bradford assay. The array membranes were blocked according to the manufacturer’s instructions and the buffer was aspirated. The sample/antibody mixtures were added to the membranes and incubated overnight at 4°C. The following day, we washed each membrane with 1× wash buffer and membranes were incubated with streptavidin-HRP for 30 min at RT. After washing, the membranes were covered with Chemi reagent mix and exposed to X-rays. Pixel densities on the developed X-ray film were quantified at each spot using ImageJ software.

Synthesis and characterization of biosealant polymers

To fabricate the biosealant for fixing S_3DP on the heart surface, we synthesized thiolated gelatin (GtnSH) and maleimide-conjugated gelatin (GtnMI) via EDC/NHS-mediated conjugation of thiol and maleimide groups to Gtn backbone. To synthesize GtnSH, we dissolved 250 mg gelatin in 125 mL of deionized water (DIW) at 37°C. After Gtn dissolution, 450.5 mg CYS (2.0 mmol) was dissolved in 2.5 mL DIW and added into the Gtn solution. We homogenously mixed them for 10 min at 37°C. Next, 192.0 mg EDC (1.0 mmol) and 116.0 mg NHS (1.0 mmol) were dissolved in DIW (2.5 mL) in a 10 mL vial. EDC and NHS solutions were sequentially added to Gtn/CYS and allowed to react for 2 h at 500 rpm. After the conjugation reaction, we injected DTT solution (617.0 mg, 4.0 mmol) dissolved in 5 mL DIW into the reactant and incubated for 24 h at 500 rpm. This reactant was serially dialyzed using a dialysis membrane bag (molecular weight cut-off = 3.5 kDa) against a 5 mM HCl solution for 36 h, followed by a 1 mM HCl solution for 24 h. The solution was changed twice to remove unconjugated residual chemicals such as unconjugated CYS, byproducts of the EDC/NHS reaction, and unreacted DTT molecules. To obtain the GtnSH polymer, we froze dialyzed the GtnSH solution at the − 80°C and lyophilized it for 5–6 days. The obtained polymers were stored in a desiccator until use.

To synthesize GtnMI, we dissolved 250 mg Gtn and 211.2 mg MHA (1.0 mmol) in 50 mL and 120 mL DPBS solution at 37°C, respectively. Each solution was mixed for 10 min at 37°C. Then, 230.1 mg EDC (1.2 mmol) and 161.1 mg NHS (1.4 mmol) were dissolved in 5 mL DPBS within a 10 mL vial, respectively. The EDC solution was mixed with the NHS solution and added to the Gtn/MHA mixture. After 2 h of reaction, the reactant was dialyzed using a dialysis membrane bag (molecular weight cut-off = 3.5 kDa) against DIW for 72 h to remove the unconjugated MHA molecules and byproducts of the EDC/NHS reaction. After dialysis, the GtnMI solution was frozen at -80°C and lyophilized for 5–6 days. Lyophilized GtnMI were stored in a deep freezer at -80°C without light.

A proton nuclear magnetic resonance (¹H-NMR) spectrometer (Agilent 400–MR, Agilent Technologies, CA) was used to characterize the chemical structures of GtnSH and GtnMI. We prepared 600 µL GtnSH and GtnMI solutions (10 mg/mL) using D₂O and analyzed each chemical structure.

Next, Ellman's assay was performed to determine the functional groups of the biosealant polymers. To measure the thiol content of GtnSH, we prepared 0–1 mg/mL L-cysteine solutions using DIW to obtain a standard curve. Next, 100 µL GtnSH solution (1 mg/mL) dissolved in DIW and standard solution were mixed with 100 µL Ellman’s reagent solution. After 20 min of reaction in the dark, the absorbance of the samples was measured at 405 nm using an ultraviolet (UV) spectrophotometry (Multiskan EX, Thermo Scientific, Rockford, IL, USA). Thiol content was calculated from a thiol standard curve at known L-cysteine concentrations (0.001–0.04 mg/mL). To measure the maleimide content of GtnMI, we performed a modified 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. After mixing 50 µL 0.04 mg/mL L-cysteine solution with 50 µL 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 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 spectrophotometry (Multiskan EX, Thermo Scientific, Rockford, IL, USA). The maleimide content was calculated from the maleimide standard curve at known MHA concentrations (0.005–0.04 mg/mL).

Fabrication of tissue adhesive biosealant

We fabricated a tissue adhesive biosealant by simply mixing GtnSH, GtnMI, and calcium peroxide (CaO₂) solutions (volume 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 first dissolved the 20 mg GtnSH and GtnMI in 200 µL and 180 µL prewarmed DPBS, respectively. We loaded the GtnSH solution into a 1 mL commercial syringe (Korea Vaccine Co., Ltd., Gyeonggi-do, Korea). Next, a 20 µL CaO₂ solution (0–2wt%) prepared using 1M Tri-HCl was mixed with 180 µL GtnMI solution and loaded into another 1 mL commercial syringe. Finally, both syringes were equipped with a dual syringe kit, and injected the solution to fabricate the tissue adhesive biosealant.

Rheological analysis of biosealant

We determined the elastic modulus (G′) of the biosealant using a rheometric fluid 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 strain of 0.1% at 37°C. To prevent solvent evaporation, a solvent trap was filled with DIW.

Tissue adhesive test of biosealant

The 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 modified 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. After swelling of decellularized porcine skin, we injected a biosealant solution using a dual syringe filled with 100 µL GtnSH and GtnMI/CaO₂ solutions between two pieces of decellularized porcine skin. We stabilized samples for 30 min under a force of 100 g at 37°C within a humid chamber. The 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 liver, kidney, heart, spleen, lungs, and skin. Various organs and skin tissues were harvested and rinsed with DPBS to remove blood and body fluids. Next, we cut the tissues in half and treated the tissue surfaces or cut edges with biosealant. We stabilized them for 5 min at 37°C within an incubator and lifted them to confirm the tissue adhesion of the biosealant. The 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 biosealants using HDFs. All reagents and solutions used in this experiment were sterilized by UV irradiation for 15 min or syringe filtration 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 [61, 62]. Briefly, 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 (0.5–2.0 mg/mL) of GtnSH and GtnMI media using DMEM-HG. After 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. The absorbance of the suspension was measured at 450 nm using a microplate reader. The 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% CO₂. Next, we prepared the biosealant eluates by incubating 40 µL biosealant droplets with 193 µL media for 24 h. Next, we treated 100 µL serially diluted eluates to HDFs and incubated for 24 h. Then, HDFs were incubated with 10% WST-1 solution for 20 min, and the absorbance was measured at 450 nm. The cell viability of the biosealant was calculated as a percentage of the control cells (TCPS).

In vivo biodegradability and tissue compatibility of biosealant

To evaluate the biodegradability of the biosealant, we subcutaneously 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 irradiation for 15 min or filtered through a syringe filter (pore size: 0.2 µm). After a week of stabilization, the mice were anesthetized with isoflurane in balanced oxygen. Next, we shaved dorsal hair and cleaned their skin with 70% ethanol and povidone. We prepared 200 µL biosealant solution and subcutaneously injected it into mice (6-week-old female C57BL/6N mice) using a dual syringe kit in vivo. For eight weeks, we monitored the change in biosealant volume and harvested the biosealant with the surrounding tissues from the mice. The tissue compatibility of the biosealant was evaluated using subcutaneous implantation and H&E staining. After eight weeks of subcutaneous injection, we harvested various organs and fixed them using 10% formalin. The animal study was performed 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 SD rats (8–10-week-old, 200–250 g) were anesthetized with isoflurane gas, endotracheally intubated, and ventilated with air using a small animal ventilator (Harvard Apparatus, Hopkinton, MA, USA). After opening the chest, MI was induced by permanent ligation of the left anterior descending coronary artery (LAD) using 6 − 0 polypropylene. Ischemia of the anterior wall of the left ventricle (LV) was confirmed by the myocardium turning red to pale pink, and the animals were randomly divided into five groups (n = 5 per group). After LAD ligation, S_3DP was immobilized on the ischemic heart surface using the biosealant. The biosealant was applied along the outside of the patch and allowed to stabilize for 3 min. After confirming that the biosealant and the 3D patch were fixed to the surface of the heart, the chest cavity was closed using 4 − 0 surgifit sutures. The chest muscles were then closed using 6 − 0 polypropylene, and the skin was sutured using 6 − 0 black silk sutures.

Table 3

Sample codes and compositions.
Sample codes	MI induction	Treatment of biosealant	3D patch transplantation (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

Echocardiography

The rats were anesthetized after 28 days of MI and S_3DP transplantation, and echocardiography was performed for cardiac function evaluation using a VisualSonics Vevo 2100 (VisualSonics Inc., Toronto, ON, Canada). All echocardiography measurements were performed by a single-blind investigator. The left ventricular internal dimension at end-diastole (LVIDd) and left ventricular internal dimension at end-systole (LVIDs) were measured using the two-dimensional (2D) targeted M-mode to calculate the left ventricular ejection fraction (LVEF) and left ventricular fractional shortening (LVFS). LVEF and LVFS were calculated using the following equations [7, 9, 31, 35]:

LVEF = [(LVIDd³ - LVIDs³)/LVIDd³ ×100%]

LVFS = [(LVIDd - LVIDs)/LVIDd ×100%]

Histological analysis

After echocardiography, whole hearts were harvested from the sacrificed rats and fixed with 10% formalin. The hearts were sliced into three sections along the transverse axis below the ligature to the apex at 2 mm intervals. The samples were dehydrated in a series of ethanol, embedded in paraffin, and cut into 4 µm-thick sections. The sections were deparaffinized in xylene, ethanol, and distilled water and stained with H&E and Masson’s trichrome to visualize cardiac fibrosis. Stained slides were scanned using the Aperio ScanScope AT Slide Scanner (Leica Biosystems Inc., Buffalo Grove, Illinois, USA) at × 20 magnification, and all images were captured using ImageScope software (Leica Biosystems Inc.). The infarction size and fibrosis area of the LV were determined using the following formula [40]:

Infarction size = [(epicardial circumference of the LV infarcted area + endocardial circumference of the LV infarcted area)/(epicardial circumference of the total LV area + endocardial circumference of the total LV area) × 100%]

Fibrosis area = [(fibrosis area of LV/total area of LV) × 100%]

The LV wall thickness was measured in the infarcted and border zones. Each zone was divided into six equal segments, and LV wall thickness was calculated by averaging the thicknesses of the six segments [5]. All histological analyses were quantified using ImageJ software.

Statistical analysis

The GraphPad Prism 5 and 7 software (GraphPad Software, 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 (S.D.). Statistical significance was set at p < 0.05.

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 3D-printable patch with dual pockets [59]. TPU-CEC363 has elasticity, flexibility, thermal property, water uptake, and biodegradability, enabling the patch to withstand repeated contractions and relaxations of the heart. We punched a patch with a diameter of 4 mm to obtain a uniform shape (Fig. 1a). In the swelling test, this patch swelled owing to the hydrophilic properties of the TPU-CEC363. We confirmed that the diameter of the patch increased slightly from 4 to 5 mm after swelling (Fig. 1b). We randomly measured the side lengths of the open and closed pockets to confirm differences in pocket size after swelling. As a result, the side length of the pockets was similarly increased by approximately 70 µm after swelling, and there was no significant difference between open and closed designs (Fig. 1c). These results indicate that we successfully fabricated a 3D-printed patch and constructed uniform dual pocket structures under moist conditions. Furthermore, the 3D-printed patch provided a uniform and sufficient cavity after swelling for spheroid formation. Fig. 1 Characteristics and preparation of the 3D patch for hADSC seeding. (a) The photograph of the 25 × 25 mm 3D patch and an illustration of 3D patch preparation for the experiments. (b-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 ± S.D. (n = 24). Scale bar: 200 µm.

Fabrication and characterization of hADSC spheroid in open and open/closed pocket patches

To fabricate the stem cell spheroid using the 3D printed patch, we loaded hADSCs suspension (3 × 10⁵ cells/patch) onto the 3D printed patch and centrifugated them to enter the hADSCs within pockets (Fig. 2a). After 24 h of cell seeding, spheroids with an uniform morphology were formed (Fig. 2b). The size of the spheroids did not significantly differ depending on the pocket type (Fig. 2c). 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 an important role in spheroid formation, and cadherin-to-cadherin binding compacts cell aggregation to form spheroids [63, 64]. Therefore, we evaluated the effect of the pocket system on spheroid formation by analyzing E-cadherin expression. The 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. 2d).

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 [63, 65]. Therefore, 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 significantly lower in these groups (Fig. 2e). In contrast, the gene expression of anti-apoptotic markers, such as BCL-2and 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. 2f). Regarding protein expression, we confirmed that BCL-2 expression was significantly increased in the S_OP and S_OPCL groups, whereas BAX expression was not significantly different from that in 2D cells (Fig. 2g). Consequently, our 3D patch with a dual module pocket demonstrated its 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 therapeutic effects compared with conventional monolayer stem cells because of their upregulated paracrine secretion [66, 67]. Previous studies reported that the expression of hypoxia-inducible factor 1-alpha (HIF-1α) increases in the mildly hypoxic inner zone of spheroids [63, 67]. HIF-1α acts as a transcription factor and regulates various target genes related to angiogenesis [67, 68]. With this rationale in mind, we compared the angiogenesis-related factor expression and paracrine effects between our system and 2D cultured stem cells in vitro.

We first evaluated gene expression of HIF-1α, resulting that the S_OP and S_OPCL groups upregulated gene expression of HIF-1α compared to the 2D group (Fig. 3a). Next, we analyzed the expression of VEGF, fibroblast growth factor (FGF)-2, and HGF as representative pro-angiogenic markers [69]. The S_OP and S_OPCL groups showed increased gene expression of each factor compared to that in the 2D group (Fig. 3a). Notably, the gene expression of VEGF was significantly increased by more than three times. Following gene expression analysis, we performed western blotting to assess the protein expression of angiogenesis-related growth factors, such as HGF and VEGF (Fig. 3b). The 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 angiogenesis, reduce apoptosis, and regulate inflammatory responses [70]. The paracrine effects of hADSCs have been validated in numerous studies of MI treatment [3, 21]. Consequently, the secretome released from hADSCs is considered a promising source for cytokine-based therapy, and its secretion and angiogenic potential have been reported to be promoted under 3D culture conditions [69, 70]. To investigate whether our 3D culture system could induce a stronger paracrine effect by releasing increased angiogenesis-related factors compared to 2D culture, we analyzed the collected CM using an angiogenesis array kit. The 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 (Figs. 3c and e). 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 activator (uPA) were modestly increased in the S_OP and S_OPCL groups compared to those in the 2D group (Fig. 3d). Interestingly, endocrine gland-derived vascular endothelial growth factor (EG-VEGF), IGFBP-2, and interleukin-1 beta (IL-1β), which are known to induce angiogenesis, were exclusively detected in the S_OP and S_OPCL groups (Fig. 3e). EG-VEGF is upregulated by hypoxia and is known to be involved in physiological angiogenesis [71, 72]. Additionally, IGFBP-2 promotes angiogenesis by upregulating VEGF expression [73], and IL-1β also plays an important role in angiogenesis as an angiogenic molecule [74]. These cytokines expressed only in the CM of the S_OP and S_OPCL groups showed greater therapeutic potential than conventional 2D cell transplantation. Consequently, we demonstrated that our system could improve the secretion of various angiogenic factors and their paracrine effects 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, which involve the thiol-ene reaction and disulfide bond formation (Figs. 4a and b). In this system, the biosealant is first 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 stiffness of the biosealant. CaO₂ and its derivative, calcium hydroxide (Ca(OH)₂), have been extensively used in food, cosmetic, and biomedical applications [61]. Notably, it is well known that CaO₂ produced hydrogen peroxide (H₂O₂) during its decomposition in aqueous media, inducing disulfide bond formation via H₂O₂-mediated thiol oxidation. Furthermore, these chemical reactions of the hydrogel enabled attachment to the heart via the interaction of maleimide and thiol groups with the primary amine and thiol groups on the heart surface. Therefore, thiolated gelatin (GtnSH) and maleimide-conjugated gelatin (GtnMI) were synthesized as polymers of biosealant using the EDC/NHS chemistry (Fig. S1a). Gtn was selected as the polymer backbone owing to its biocompatibility, biodegradability, bioactivity, and easy modification [61]. Next, we characterized the chemical structures of GtnSH and GtnMI by ¹H-NMR spectroscopy. As a result, GtnSH exhibited new peaks appearance at 2.8 and 3.5 ppm, indicating the conjugation of cystamine to the carboxyl group of Gtn [75] (Fig. S1b). For GtnMI, we confirmed a sharp peak appearance at 6.8 ppm and a decreased peak area at 3.0 ppm, which indicates the introduction of the maleimide group to a primary amine group of Gtn [62, 76] (Fig. S1c). 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 successfully synthesized polymers for fabricating tissue adhesive biosealant.

The viscoelasticity of the biosealant is crucial for patch fixation and sealing, particularly in resisting external forces such as the heartbeat. Accordingly, we fabricated tissue adhesive biosealants with various CaO₂ concentrations and measured their elastic modulus (G′) using a rheometer. As increasing CaO₂ concentration, the G′ of hydrogels was reinforced (0 wt%, 230 Pa; 0.01 wt%, 360 Pa; 0.05 wt%, 420 Pa; 0.1wt%, 920 Pa) (Fig. 4c). Typically, the stiffness of the rat myocardium is known to range from 0.1 to 140 kPa [77]. Additionally, it has been reported that the optimal elastic modulus of the hydrogel matrix for cardiac repair ranges from 380 to 600 Pa, representing a relatively soft matrix. This soft matrix facilitates the transmission of mechanical signals and the coordination of myocardial tissue, improving the reconstruction of cardiac function [78]. Consequently, we controlled the stiffness of our biosealant using CaO₂ concentrations, providing a suitable matrix in the dynamic cardiac environment for effective cardiac repair.

The tissue adhesive force of the biosealant should withstand external forces such as heartbeat and blood pressure, and stably immobilize the patch in place. Therefore, we measured the tissue adhesive force of our biosealant at various CaO₂ concentrations using decellularized porcine skin and a UTM (Fig. 4d). We observed that CaO₂ addition reinforced the tissue adhesive strength of the biosealant compared to that of the control (CaO₂ 0 wt%, 11.3 kPa; 0.01 wt%, 15.7 kPa; 0.05 wt%, 15.4 kPa; 0.1 wt%, 14.0 kPa) (Fig. 4e). These tissue adhesive strengths were higher than commercially available fibrin glue (5 to 10 kPa) [79, 80]. Our biosealant could strongly attach to tissue surfaces through a thiol-ene reaction and disulfide bond formation, in contrast to fibrin glue, which relies on hydrogen bond formation during coagulation [81] (Fig. 4b). Interestingly, we confirmed that the increasing tissue adhesive strength was saturated between CaO₂ 0.01 and 0.05 wt% and decreased at CaO₂ 0.1 wt%. The number of tissue adhesive functional groups, such as thiol groups, and gelation time were rapidly decreased when the CaO₂ concentration was increased, resulting in insufficient tissue adhesive groups and time to homogeneously interact with the tissue surface. Accordingly, we selected CaO₂ 0.1 wt% as the optimal composition of biosealant for immobilizing a patch in vivo.

Next, we determined the feasibility of biosealant on various organ surfaces (Fig. 4f). Interestingly, our sealant bonded between tissue surfaces and cut area of various organs. The tissue adhesion of our sealant targets the thiol and primary amine groups present on most tissue surfaces. This suggests that our biosealant can be used as an adhesive to fix patches on various tissue surfaces.

Biocompatible oxygen-releasing biosealant

The biocompatibility and biodegradability of biosealant are essential properties for successful clinical applications because they ensure the safety of materials and their byproducts, and suitable maintenance of the sealant at a rate compatible with tissue healing [51]. To utilize our sealant for immobilizing a S_3DP in vivo, we first 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. 5a). This result indicates that our biosealant may be cytocompatible with host cells. Additionally, the biosealant exhibited excellent cell viability above 95.2% of TCPS (Fig. 5b). This was confirmed by live and dead staining (Fig. 5c).

We next subcutaneously injected the biosealant in mice to evaluate its biodegradability and tissue compatibility in vivo. Figure 5d shows that the sealant was maintained for 8 weeks and was sustainably degraded in vivo. This suggests that our sealant is biodegradable and can be maintained at a specific 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 after subcutaneous injection. Compared with normal organs, no specific lesions were observed in any organs in the sealant group (Fig. 5e).

Consequently, our biosealant proved to be biocompatible and biodegradable with the potential to fix patches in vivo. Based on these results, we investigated the therapeutic effects of the S_3DP on cardiac infarction.

Therapeutic effect of S_3DP on cardiac infarction

We investigated the therapeutic efficacy of S_3DP using a rat model of MI. To fabricate the MI model, we tied the LAD to induce permanent vascular occlusion and observed changes in myocardial color [82]. After fabricating the MI model, 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 under a dynamic heartbeat, and it was stably maintained for 3 days (Fig. 6a). To compare the effect of patches or spheroids on MI treatment, we selected the sealant-only (SO) and open/closed patch-only (PO) groups for comparison. After 4 weeks after transplantation, we examined LVEF and LVFS, which are the most common parameters used to quantify heart function using echocardiography [8, 83, 84] (Fig. 6b). We first measured LVIDd and LVIDs to calculate the LVEF and LVFS using the photographed ultrasound results [84]. LVIDd and LVIDs in the controls (MI, SO, and PO) and S_OP group were significantly higher than those in the normal group. However, there was no significant difference in the results of LVIDd and LVIDs between the S_OPCL and normal group. Moreover, the S_OPCL group showed significantly decreased LVIDs compared to the control groups (Figs. 6c and d). LVIDs is the result of myocardial contractility, meaning that the narrower the LVIDs indicates the better heart function. This result demonstrated that the delivery of hADSC spheroids was effective in improving myocardial function through their paracrine effect, 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 [85], and LVFS indicates the percentage change in the LV internal dimensions between systole and diastole [8]. All groups showed lower heart function than the normal group in terms of both LVEF and LVFS. However, we confirmed that the S_3DP groups (S_OP and S_OPCL) had significantly improved cardiac functions with increased LVEF and LVFS compared to the control groups (Figs. 6e and f). Consequently, we successfully immobilized our S_3DP system using a biosealant without physical damage, which occurs by suturing or injection. Our system was stably maintained for 4 weeks even in the presence of pericardial fluid. Moreover, we demonstrated its therapeutic effects in the restoration of cardiac function. Notably, the combination of open and closed pockets was the most effective through the paracrine effect 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 fibroblasts are activated, replacing the damaged myocardium with collagen-based scar tissue [10, 86]. However, this fibrotic tissue cannot perform normal functions, such as contraction and relaxation. Thus, reducing the fibrotic area in the LV is important for maintaining or restoring cardiac function. MI is accompanied by ECM degradation, LV wall thinning, and ventricular enlargement [14, 87]. Based on this, we histologically analyzed LV wall thickness, infarction size, and fibrosis area. We sectioned the whole heart into three areas at 2 mm intervals from the LAD-tied part, which are indicated in areas 1 to 3 (Figs. 7a and c). We then divided the cross-sectional tissue of the heart into three zones: an infarcted zone (IZ), a border zone (BZ), and a remote zone (RZ). We focused our analysis on the IZ and BZ [88] (Fig. 7b). To determine the fibrotic area of the heart, we performed H&E and Masson’s trichrome staining on sectioned tissues (Figs. 7d and e) The blue area indicates fibrotic tissue (collagen) and the red area indicates the myocardium (Fig. 7e).

During post-infarction repair, fibrotic tissue formation causes progressive thinning of LV wall thickness [87, 89]. Therefore, we first measured the LV wall thickness according to the zone. In area 1, the LV wall thickness in the S_3DP groups (S_OP and S_OPCL) was significantly 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 significant difference 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 significant differences were observed between all groups (Fig. 7f). Based on these results, we confirmed 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 effective than in the S_OP group.

Furthermore, we analyzed the infarct size and fibrotic area of LV. The 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. 7b. There were no differences among the MI, SO, and PO 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. 7g). To determine the percentage of fibrosis, we measured the total LV and fibrotic areas in Fig. 7e. There were no significant differences between the control groups in any area. In all areas, the fibrosis area of the S_OP group was decreased compared to those of the control groups, although the differences were not statistically significant. In contrast, the fibrosis area of the S_OPCL group was notably reduced compared to those of all groups in all areas and revealed only significant differences in areas 1 and 2 (Fig. 7h).

In conclusion, we confirmed that S_3DP transplantation effectively reduces fibrotic scar tissue and increases LV wall thickness. Notably, the S_OPCL group showed greater potential therapeutic effects than the S_OP group, which is consistent with the cardiac function results.

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 engraftment rate and paracrine effect of the spheroids. This patch 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 effects in vitro. Next, we fabricated a tissue adhesive biosealant as a sutureless method for immobilizing S_3DP to the heart tissue, maintaining its position 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 combination of open and closed pocket, S_OPCL, was more effective than the open pocket, S_OP, demonstrating that the protective effect of the closed pocket acts synergistically with direct spheroid delivery and paracrine effects. 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.

¹H-NMR	Proton nuclear magnetic resonance
2D	Two-dimensional
3D	Three-dimensional
BAK	Bcl-2 homologous antagonist killer
BAX	Bcl-2-associated X
BCL-2	B-cell lymphoma 2
BCL-xL	B-cell lymphoma extra-large
BZ	Border zone
Ca(OH)₂	Calcium hydroxide
CaO₂	Calcium peroxide
CM	Conditioned medium
CO₂	Carbon dioxide
CYS	Cystamine dihydrochloride
D₂O	Deuterium oxide
DIW	Deionized water
DMEM	Dulbecco’s modified Eagle’s medium
DPBS	Dulbecco's phosphate-buffered saline
DTT ECM	DL-dithiothreitol Extracellular matrix
EDC	1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride
EG-VEGF	Endocrine gland-derived vascular endothelial growth factor
FBS	Fetal bovine serum
FDA	Food and Drug Administration
FGF	Fibroblast growth factor
G′	Elastic modulus
GAPDH	Glyceraldehyde 3-phosphate dehydrogenase
Gtn	Gelatin
GtnMI	Maleimide-conjugated gelatin
GtnSH	Thiolated gelatin
hADSC	Human adipose-derived stem cell
H&E	Hematoxylin and eosin
H₂O₂	Hydrogen peroxide
HCl	Hydrochloric acid
HDF	Human dermal fibroblast
HGF	Hepatocyte growth factor
HIF-1α	Hypoxia-inducible factor 1-alpha
IACUC	Institutional Animal Care and Use Committee
IGFBP	Insulin-like growth factor-binding protein
IL-1β	Interleukin-1β
IZ	Infarcted zone
LAD	Left anterior descending coronary artery
LV	Left ventricle
LVEF	Left ventricular ejection fraction
LVFS	Left ventricular fractional shortening
LVIDd	Left ventricular internal dimension at end-diastole
LVIDs	Left ventricular internal dimension at end-systole
MHA	6-maleimidohexanoic acid
MI	Myocardial infarction
MSC	Mesenchymal stem cell
NBCS	Newborn calf serum
NHS	N-hydroxysuccinimide
N.S	No significance
PBS	Phosphate-buffered saline
PCL	Polycaprolactone
PEG	Poly(ethylene glycol)
PO	Open/closed patch-only group
P/S	Penicillin/streptomycin antibiotics
PU	Polyurethanes
qRT-PCR	Quantitative reverse transcription-polymerase chain reaction
RT	Room temperature
RZ	Remote zone
SD	Sprague-Dawley
SDS-PAGE	Sodium dodecyl sulfate-polyacrylamide gel electrophoresis
S.D.	Standard deviation
S_3DP	Stem cell spheroid-laden 3D patches
S_OP	Spheroid-laden open pocket patch
S_OPCL	Spheroid-laden open/closed pocket patch
SO	Sealant-only group
TCPS	Tissue culture polystyrene
TPU-CEC363	Elastomeric bioscaffold
uPA	Urokinase-type plasminogen activator
UV UTM	Ultraviolet Universal testing machine
VEGF	Vascular endothelial growth factor

Acknowledgments

Not applicable.

Authors’ contributions

H.R.J. and J.I.K. conducted the experiments, analyzed data, designed the figures, 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 contributed to interpretation of data. All authors have read and approved the manuscript.

Funding

This 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, 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).

Availability of data and materials

The datasets used and/or analyzed during the current study are available from the corresponding author on reasonable request.

Ethics approval and consent to participate

The animal experiments in this study were approved by the Institutional Animal Care and Use Committee of the Incheon National University (INU-ANIM-2022-02) and Samsung Medical Center (SMCIACUC2021-03-23-001).

Consent for publication

Not applicable.

Competing interests

The authors declare no conflict of interest.

Author details

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

Benjamin EJ, Muntner P, Alonso A, Bittencourt MS, Callaway CW, Carson AP, et al. Heart disease and stroke statistics-2019 update: a report from the american heart association. Circulation. 2019;139(10):e56–28.
Grundy SM, Stone NJ, Bailey AL, Beam C, Birtcher KK, Blumenthal RS, AHA/ACC/AACVPR/AAPA/ABC/ACPM/ADA/AGS/APhA et al. /ASPC/NLA/PCNA guideline on the management of blood cholesterol: executive summary: a report of the american college of cardiology/american heart association task force on clinical practice guidelines. J Am Coll Cardiol. 2019;73(24):3168–3209.
Li L, Chen X, Wang WE, Zeng C. How to improve the survival of transplanted mesenchymal stem cell in ischemic heart? Stem Cells Int. 2016;2016:9682757.
Huang L, Ma W, Ma Y, Feng D, Chen H, Cai B. Exosomes in mesenchymal stem cells, a new therapeutic strategy for cardiovascular diseases? Int J Biol Sci. 2015;11(2):238–45.
Song Y, Zhang C, Zhang J, Sun N, Huang K, Li H, et al. An injectable silk sericin hydrogel promotes cardiac functional recovery after ischemic myocardial infarction. Acta Biomater. 2016;41:210–23.
Tang J, Wang J, Huang K, Ye Y, Su T, Qiao L, et al. Cardiac cell-integrated microneedle patch for treating myocardial infarction. Sci Adv. 2018;4(11):eaat9365.
Zhao C, Tian S, Liu Q, Xiu K, Lei I, Wang Z, et al. Biodegradable nanofibrous temperature-responsive gelling microspheres for heart regeneration. Adv Funct Mater. 2020;30(21):2000776.
Portillo Esquivel LE, Zhang B. Application of cell, tissue, and biomaterial delivery in cardiac regenerative therapy. ACS Biomater Sci Eng. 2021;7(3):1000–21.
Jin J, Jeong SI, Shin YM, Lim KS, Shin HS, Lee YM, et al. Transplantation of mesenchymal stem cells within a poly(lactide-co-epsilon-caprolactone) scaffold improves cardiac function in a rat myocardial infarction model. Eur J Heart Fail. 2009;11(2):147–53.
Yao Y, Ding J, Wang Z, Zhang H, Xie J, Wang Y, et al. ROS-responsive polyurethane fibrous patches loaded with methylprednisolone (MP) for restoring structures and functions of infarcted myocardium in vivo. Biomaterials. 2020;232:119726.
Das S, Nam H, Jang J. 3D bioprinting of stem cell-laden cardiac patch: a promising alternative for myocardial repair. APL Bioeng. 2021;5(3):031508.
Wee JH, Yoo K-D, Sim SB, Kim HJ, Kim HJ, Park KN, et al. Stem cell laden nano and micro collagen/PLGA bimodal fibrous patches for myocardial regeneration. Biomater Res. 2022;26(1):79.
Lee SH, Hong YJ, Ahn Y, Jeong MH. Past, present, and future of management of acute myocardial infarction. J Cardiovasc Interv. 2023;2(2):51–65.
Feng J, Xing M, Qian W, Qiu J, Liu X. An injectable hydrogel combining medicine and matrix with anti-inflammatory and pro-angiogenic properties for potential treatment of myocardial infarction. Regen Biomater. 2023;10:rbad036.
Cho MS, Park D-W. Stent thrombosis and optimal duration of dual antiplatelet therapy after coronary stenting in contemporary practice. Korean J Intern Med. 2017;32(5):769–79.
Hu H, Shen L. Drug-coated balloons in the treatment of acute myocardial infarction (Review). Exp Ther Med. 2021;21(5):464.
Jackson D, Tong D, Layland J. A review of the coronary applications of the drug coated balloon. Int J Cardiol. 2017;226:77–86.
Schafer A, Konig T, Bauersachs J, Akin M. Novel therapeutic strategies to reduce reperfusion injury after acute myocardial infarction. Curr Probl Cardiol. 2022;47(12):101398.
Vidal-Cales P, Cepas-Guillen PL, Brugaletta S, Sabate M. New interventional therapies beyond stenting to treat ST-segment elevation acute myocardial infarction. J Cardiovasc Dev Dis. 2021;8(9):100.
Li M, Wu H, Yuan Y, Hu B, Gu N. Recent fabrications and applications of cardiac patch in myocardial infarction treatment. View. 2022;3(2):20200153.
Wu R, Hu X, Wang J. Concise review: optimized strategies for stem cell-based therapy in myocardial repair: clinical translatability and potential limitation. Stem Cells. 2018;36(4):482–500.
Li J, Hu S, Zhu D, Huang K, Mei X, Lopez de Juan Abad B, et al. All roads lead to rome (the heart): cell retention and outcomes from various delivery routes of cell therapy products to the heart. J Am Heart Assoc. 2021;10(8):e020402.
Sheng CC, Zhou L, Hao J. Current stem cell delivery methods for myocardial repair. Biomed Res Int. 2013;2013:547902.
Sun R, Li X, Liu M, Zeng Y, Chen S, Zhang P. Advances in stem cell therapy for cardiovascular disease (Review). Int J Mol Med. 2016;38(1):23–9.
Li J, Lv Y, Zhu D, Mei X, Huang K, Wang X, et al. Intrapericardial hydrogel injection generates high cell retention and augments therapeutic effects of mesenchymal stem cells in myocardial infarction. Chem Eng J. 2022;427:131581.
Williams AR, Hare JM. Mesenchymal stem cells: biology, pathophysiology, translational findings, and therapeutic implications for cardiac disease. Circ Res. 2011;109(8):923–40.
Carvalho E, Verma P, Hourigan K, Banerjee R. Myocardial infarction: stem cell transplantation for cardiac regeneration. Regen Med. 2015;10(8):1025–43.
Li S, Wang X, Li J, Zhang J, Zhang F, Hu J, et al. Advances in the treatment of ischemic diseases by mesenchymal stem cells. Stem Cells Int. 2016;2016:5896061.
Behbahan IS, Keating A, Gale RP. Bone marrow therapies for chronic heart disease. Stem Cells. 2015;33(11):3212–27.
Chow A, Stuckey DJ, Kidher E, Rocco M, Jabbour RJ, Mansfield CA, et al. Human induced pluripotent stem cell-derived cardiomyocyte encapsulating bioactive hydrogels improve rat heart function post myocardial infarction. Stem Cell Rep. 2017;9(5):1415–22.
Waters R, Alam P, Pacelli S, Chakravarti AR, Ahmed RPH, Paul A. Stem cell-inspired secretome-rich injectable hydrogel to repair injured cardiac tissue. Acta Biomater. 2018;69:95–106.
Suhar RA, Doulames VM, Liu Y, Hefferon ME, Figueroa O, Buabbas H, et al. Hyaluronan and elastin-like protein (HELP) gels significantly improve microsphere retention in the myocardium. Biomater Sci. 2022;10(10):2590–608.
Park TY, Oh J-M, Cho JS, Sim SB, Lee J, Cha HJ. Stem cell-loaded adhesive immiscible liquid for regeneration of myocardial infarction. J Control Release. 2020;321:602–15.
Kim CW, Kim CJ, Park E-H, Ryu S, Lee Y, Kim E, et al. MSC-encapsulating in situ cross-linkable gelatin hydrogels to promote myocardial repair. ACS Appl Bio Mater. 2020;3(3):1646–55.
Choi A, Kim H, Han H, Park J-H, Kim J-J, Sim W-S, et al. Sutureless transplantation of in vivo priming human mesenchymal stem cell sheet promotes the therapeutic potential for cardiac repair. Biofabrication. 2022;15(1):015009.
Yao Y, Li A, Wang S, Lu Y, Xie J, Zhang H, et al. Multifunctional elastomer cardiac patches for preventing left ventricle remodeling after myocardial infarction in vivo. Biomaterials. 2022;282:121382.
Liang S, Zhang Y, Wang H, Xu Z, Chen J, Bao R, et al. Paintable and rapidly bondable conductive hydrogels as therapeutic cardiac patches. Adv Mater. 2018;30(23):e1704235.
Xie J, Yao Y, Wang S, Fan L, Ding J, Gao Y, et al. Alleviating oxidative injury of myocardial infarction by a fibrous polyurethane patch with condensed ROS-scavenging backbone units. Adv Healthc Mater. 2022;11(4):2101855.
Su T, Huang K, Mathews KG, Scharf VF, Hu S, Li Z, et al. Cardiac stromal cell patch integrated with engineered microvessels improves recovery from myocardial infarction in rats and pigs. ACS Biomater Sci Eng. 2020;6(11):6309–20.
Jang J, Park H-J, Kim S-W, Kim H, Park JY, Na SJ, et al. 3D printed complex tissue construct using stem cell-laden decellularized extracellular matrix bioinks for cardiac repair. Biomaterials. 2017;112:264–74.
Park B-W, Jung S-H, Das S, Lee SM, Park J-H, Kim H, et al. In vivo priming of human mesenchymal stem cells with hepatocyte growth factor-engineered mesenchymal stem cells promotes therapeutic potential for cardiac repair. Sci Adv. 2020;6(13):eaay6994.
Portillo-Esquivel LE, Nanduri V, Zhang F, Liang W, Zhang B. z-Wire: a microscaffold that supports guided tissue assembly and intramyocardium delivery for cardiac repair. Adv Healthc Mater. 2020;9(14):2000358.
Chang C, Yan J, Yao Z, Zhang C, Li X, Mao HQ. Effects of mesenchymal stem cell-derived paracrine signals and their delivery strategies. Adv Healthc Mater. 2021;10(7):e2001689.
Kim W, Gwon Y, Park S, Kim H, Kim J. Therapeutic strategies of three-dimensional stem cell spheroids and organoids for tissue repair and regeneration. Bioact Mater. 2023;19:50–74.
Jeong G-J, Im G-B, Lee T-J, Kim S-W, Jeon HR, Lee D-H, et al. Development of a stem cell spheroid-laden patch with high retention at skin wound site. Bioeng Transl Med. 2021;7(2):e10279.
Wang Y, Li G, Yang L, Luo R, Guo G. Development of innovative biomaterials and devices for the treatment of cardiovascular diseases. Adv Mater. 2022;34(46):e2201971.
Griffin KH, Fok SW, Kent Leach J. Strategies to capitalize on cell spheroid therapeutic potential for tissue repair and disease modeling. npj Regen Med. 2022;7(1):70.
Tariq U, Gupta M, Pathak S, Patil R, Dohare A, Misra SK. Role of biomaterials in cardiac repair and regeneration: therapeutic intervention for myocardial infarction. ACS Biomater Sci Eng. 2022;8(8):3271–98.
Chang D, Fan T, Gao S, Jin Y, Zhang M, Ono M. Application of mesenchymal stem cell sheet to treatment of ischemic heart disease. Stem Cell Res Ther. 2021;12(1):384.
Wei X, Wang L, Duan C, Chen K, Li X, Guo X, et al. Cardiac patches made of brown adipose-derived stem cell sheets and conductive electrospun nanofibers restore infarcted heart for ischemic myocardial infarction. Bioact Mater. 2023;27:271–87.
Nam S, Mooney D. Polymeric tissue adhesives. Chem Rev. 2021;121(18):11336–84.
Bao Z, Gao M, Sun Y, Nian R, Xian M. The recent progress of tissue adhesives in design strategies, adhesive mechanism and applications. Mater Sci Eng C Mater Biol Appl. 2020;111:110796.
Duan W, Bian X, Bu Y. Applications of bioadhesives: a mini review. Front Bioeng Biotechnol. 2021;9:716035.
Bu Y, Pandit A. Cohesion mechanisms for bioadhesives. Bioact Mater. 2022;13:105–18.
Yuk H, Varela CE, Nabzdyk CS, Mao X, Padera RF, Roche ET, et al. Dry double-sided tape for adhesion of wet tissues and devices. Nature. 2019;575:169–74.
Deng J, Yuk H, Wu J, Varela CE, Chen X, Roche ET, et al. Electrical bioadhesive interface for bioelectronics. Nat Mater. 2021;20:229–36.
Wu T, Zhang X, Liu Y, Cui C, Sun Y, Liu W. Wet adhesive hydrogel cardiac patch loaded with anti-oxidative, autophagy-regulating molecule capsules and MSCs for restoring infarcted myocardium. Bioact Mater. 2023;21:20–31.
Yao Y, Li A, Wang S, Lu Y, Xie J, Zhang H, et al. Multifunctional elastomer cardiac patches for preventing left ventricle remodeling after myocardial infarction in vivo. Biomaterials. 2022;282:121382.
Shin J, Choi S, Kim JH, Cho JH, Jin Y, Kim S, et al. Tissue tapes-phenolic hyaluronic acid hydrogel patches for off-the-shelf therapy. Adv Funct Mater. 2019;29(49):1903863.
Kim RI, Lee G, Lee J-H, Park JJ, Lee AS, Hwang SS. Structure–property relationships of 3D-printable chain-extended block copolymers with tunable elasticity and biodegradability. ACS Appl Polym Mater. 2021;3(9):4708–16.
Park S, Park KM. Hyperbaric oxygen-generating hydrogels. Biomaterials. 2018;182:234–44.
Kang JI, Park KM. Oxygen-supplying syringe to create hyperoxia-inducible hydrogels for in situ tissue regeneration. Biomaterials. 2023;293:121943.
Kouroupis D, Correa D. Increased mesenchymal stem cell functionalization in three-dimensional manufacturing settings for enhanced therapeutic applications. Front Bioeng Biotechnol. 2021;9:621748.
Cui X, Hartanto Y, Zhang H. Advances in multicellular spheroids formation. J R Soc Interface. 2017;14(127):20160877.
Egger D, Tripisciano C, Weber V, Dominici M, Kasper C. Dynamic cultivation of mesenchymal stem cell aggregates. Bioeng (Basel). 2018;5(2):48.
Han MA, Jeon JH, Shin JY, Kim HJ, Lee JS, Seo CW, et al. Intramyocardial delivery of human cardiac stem cell spheroids with enhanced cell engraftment ability and cardiomyogenic potential for myocardial infarct repair. J Control Release. 2021;336:499–509.
Jaukovic A, Abadjieva D, Trivanovic D, Stoyanova E, Kostadinova M, Pashova S, et al. Specificity of 3D MSC spheroids microenvironment: impact on MSC behavior and properties. Stem Cell Rev Rep. 2020;16(5):853–75.
Ziello JE, Jovin IS, Huang Y. Hypoxia-Inducible Factor (HIF)-1 regulatory pathway and its potential for therapeutic intervention in malignancy and ischemia. Yale J Biol Med. 2007;80(2):51–60.
Zhao L, Johnson T, Liu D. Therapeutic angiogenesis of adipose-derived stem cells for ischemic diseases. Stem Cell Res Ther. 2017;8(1):125.
Cunningham CJ, Redondo-Castro E, Allan SM. The therapeutic potential of the mesenchymal stem cell secretome in ischaemic stroke. J Cereb Blood Flow Metab. 2018;38(8):1276–92.
Brouillet S, Hoffmann P, Benharouga M, Salomon A, Schaal J-P, Feige J-J, et al. Molecular characterization of EG-VEGF-mediated angiogenesis: differential effects on microvascular and macrovascular endothelial cells. Mol Biol Cell. 2010;21(16):2832–43.
Alfaidy N, Brouilet S, Rajaraman G, Kalionis B, Hoffmann P, Barjat T, et al. The emerging role of the prokineticins and homeobox genes in the vascularization of the placenta: physiological and pathological aspects. Front Physiol. 2020;11:591850.
Azar WJ, Azar SHX, Higgins S, Hu J-F, Hoffman AR, Newgreen EF, et al. IGFBP-2 enhances VEGF gene promoter activity and consequent promotion of angiogenesis by neuroblastoma cells. Endocrinology. 2011;152(9):3332–42.
Chen Y-C, Fu Y-S, Tsai S-W, Wu P-K, Chen C-M, Chen W-M, et al. IL-1b in the secretomes of MSCs seeded on human decellularized allogeneic bone promotes angiogenesis. Int J Mol Sci. 2022;23(23):15301.
Shirani S, Varshosaz J, Rostami M, Mirian M. Redox responsive polymeric micelles of gellan gum/abietic acid for targeted delivery of ribociclib. Int J Biol Macromol. 2022;215(31):334–45.
Gilchrist AE, Serrano JF, Ngo MT, Hrnjak Z, Kim S, Harley BAC. Encapsulation of murine hematopoietic stem and progenitor cells in a thiol-crosslinked maleimide-functionalized gelatin hydrogel. Acta Biomater. 2021;131:138–48.
Zheng Z, Tan Y, Li Y, Liu Y, Yi G, Yu C-Y, et al. Biotherapeutic-loaded injectable hydrogels as a synergistic strategy to support myocardial repair after myocardial infarction. J Control Release. 2021;335:216–36.
Zhou J, Liu W, Zhao X, Xian Y, Wu W, Zhang X, et al. Natural melanin/alginate hydrogels achieve cardiac repair through ROS scavenging and macrophage polarization. Adv Sci. 2021;8(20):e2100505.
Lih E, Lee JS, Park KM, Park KD. Rapidly curable chitosan-PEG hydrogels as tissue adhesives for hemostasis and wound healing. Acta Biomater. 2012;8(9):3261–9.
Mehdizadeh M, Weng H, Gyawali D, Tang L, Yang J. Injectable citrate-based mussel-inspired tissue bioadhesives with high wet strength for sutureless wound closure. Biomaterials. 2012;33(32):7972–83.
Bal-Ozturk A, Cecen B, Avci-Adali M, Topkaya SN, Alarcin E, Yasayan G, et al. Tissue adhesives: From research to clinical translation. Nano Today. 2021;36:101049.
Xie Y, Xing Z, Wei J, Sun X, Zhao B, Chen Y, et al. Levosimendan postconditioning attenuates cardiomyocyte apoptosis after myocardial infarction. J Healthc Eng. 2022;2022:2988756.
Tissot C, Singh Y, Sekarski N. Echocardiographic evaluation of ventricular function-for the neonatologist and pediatric intensivist. Front Pediatr. 2018;6:79.
Rottman JN, Ni G, Brown M. Echocardiographic evaluation of ventricular function in mice. Echocardiography. 2007;24(1):83–9.
Cikes M, Solomon SD. Beyond ejection fraction: an integrative approach for assessment of cardiac structure and function in heart failure. Eur Heart J. 2016;37(21):1642–50.
Bektik E, Fu J-D. Ameliorating the fibrotic remodeling of the heart through direct cardiac reprogramming. Cells. 2019;8(7):679.
Weil BR, Neelamegham S. Selectins and immune cells in acute myocardial infarction and post-infarction ventricular remodeling: pathophysiology and novel treatments. Front Immunol. 2019;10:300.
Driesen RB, Verheyen FK, Dijkstra P, Thone F, Cleutjens JP, Lenders M-H, et al. Structural remodelling of cardiomyocytes in the border zone of infarcted rabbit heart. Mol Cell Biochem. 2007;302(1–2):225–32.
Konstam MA, Kramer DG, Patel AR, Maron MS, Udelson JE. Left ventricular remodeling in heart failure: current concepts in clinical significance and assessment. JACC Cardiovasc Imaging. 2011;4(1):98–108.

Additionalfile1.docx
Supplementary Information Additional file 1: Fig. S1. Synthesis and characterization of GtnSH and GtnMI. (a) Schematic illustration of GtnSH and GtnMI synthesis using EDC/NHS chemistry. ¹H NMR spectra of (b) GtnSH and (c) GtnMI with the content of each functional group.

Download PDF

Editorial decision: Major revision
06 Nov, 2023
Reviewers agreed at journal
09 Oct, 2023
Reviewers invited by journal
08 Oct, 2023
Editor assigned by journal
27 Sep, 2023
First submitted to journal
25 Sep, 2023

You are reading this latest preprint version

Transplantation of stem cell spheroid-laden three-dimensional patches with bioadhesives for the treatment of myocardial infarction.

Status:

Version 1

Abstract

Background

Methods

Results

Conclusion

Figures

Background

Materials and methods

Materials

Fabrication of open and open/closed pocket patches for stem cell spheroid encapsulation

Characterization of open and open/closed pocket patches

Preparation of 3D patch and spheroid formation in 3D patch

Characterization of stem cell spheroid

Quantitative reverse transcription-polymerase chain reaction (qRT-PCR)

Western blot analysis

Angiogenesis antibody array

Synthesis and characterization of biosealant polymers

Fabrication of tissue adhesive biosealant

Rheological analysis of biosealant

Tissue adhesive test of biosealant

Cytocompatibility of biosealant

In vivo biodegradability and tissue compatibility of biosealant

Acute MI model of rats and transplantation of S_3DP in MI model in vivo

Echocardiography

Histological analysis

Statistical analysis

Results and discussion

Fabrication of open and open/closed pocket patches

Fabrication and characterization of hADSC spheroid in open and open/closed pocket patches

Paracrine angiogenic effect of S_3DP

Fabrication and characterization of tissue adhesive biosealant

Biocompatible oxygen-releasing biosealant

Therapeutic effect of S_3DP on cardiac infarction

Reduced cardiac fibrosis of S_3DP on cardiac infarction

Conclusion

Abbreviations

Declarations

References

Supplementary Files

Status:

Version 1