# Engineered biomimetic micro/nano-materials for tissue regeneration

Juan Wang, Tonghe Zhu, Yawei Du, Liang Chen, Régis Laurent

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Juan Wang, Tonghe Zhu, Yawei Du, Liang Chen, Régis Laurent. Engineered biomimetic micro/nano-materials for tissue regeneration. Frontiers Media SA., 2023, (Frontiers in Bioengineering and Biotechnology), 10.3389/978-2-8325-4749-6. hal-04988112

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Engineered biomimetic
micro/nano-materials for
tissue regeneration

Edited by

Juan Wang, Tonghe Zhu, Yawei Du,
Liang Chen and Régis Laurent

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Frontiers in Bioengineering and Biotechnology

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ISSN 1664-8714
ISBN 978-2-8325-4749-6
DOI 10.3389/978-2-8325-4749-6

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# Engineered biomimetic micro/nano-materials for tissue regeneration

**Topic editors**

Juan Wang – Shanghai Jiao Tong University, China

Tonghe Zhu – Shanghai University of Engineering Sciences, China

Yawei Du – Shanghai Jiao Tong University, China

Liang Chen – Shanghai Jiao Tong University, China

Régis Laurent – UPR8241 Laboratoire de Chimie de Coordination (LCC), France

**Citation**

Wang, J., Zhu, T., Du, Y., Chen, L., Laurent, R., eds. (2024). *Engineered biomimetic micro/nano-materials for tissue regeneration*. Lausanne: Frontiers Media SA.
doi: 10.3389/978-2-8325-4749-6
Table of
contents

05 Chemically crosslinked hyaluronic acid-chitosan hydrogel for application on cartilage regeneration

Sandra Escalante, Gustavo Rico, José Becerra, Julio San Román, Blanca Vázquez-Lasa, Maria Rosa Aguilar, Iván Durán and Luis García-Fernández

19 Application of thermosensitive-hydrogel combined with dental pulp stem cells on the injured fallopian tube mucosa in an animal model

Lihua Luo, Qunyan Zhu, Yejian Li, Fengting Hu, Jiangtao Yu, Xiangyan Liao, Zhenjie Xing, Yan He and Qingsong Ye

34 SiO₂ nanosphere coated tough catheter with superhydrophobic surface for improving the antibacteria and hemocompatibility

Weixing Zhang, Juan Du, Tonghe Zhu and Ruilan Wang

45 A poly (glycerol-sebacate-acrylate) nanosphere enhanced injectable hydrogel for wound treatment

Jiajia Luo, Fenglei Sun, Pinhua Rao, Tonghe Zhu, Yonghang Liu, Juan Du, Sihao Chen, Xiangyun Jin, Jiale Jin and Yi Chai

57 Adipose stem cells-derived exosomes modified gelatin sponge promotes bone regeneration

Gen Li, Yin Zhang, Jiezhou Wu, Renhao Yang, Qi Sun, Yidong Xu, Bo Wang, Ming Cai, Yang Xu, Chengyu Zhuang and Lei Wang

68 The function of immunomodulation and biomaterials for scaffold in the process of bone defect repair: A review

Changchao Dong, Gang Tan, Guangyan Zhang, Wei Lin and Guanglin Wang

82 Preparation and characterization of a novel triple composite scaffold containing silk fibroin, chitosan, extracellular matrix and the mechanism of Akt/FoxO signaling pathway in colonic cancer cells cultured in 3D

Zhipeng Cao, Liang Chen, Gengming Niu, Yan Li, Zhiqing Hu, Runqi Hong, Xiaotian Zhang, Liang Hong, Shanliang Han and Chongwei Ke

95 Physical processing for decellularized nerve xenograft in peripheral nerve regeneration

Ming-Wei Hsu, Szu-Han Chen, Wan-Ling Tseng, Kuo-Shu Hung, Tzu-Chun Chung, Sheng-Che Lin, Jahyun Koo and Yuan-Yu Hsueh

101 Development of tissue-engineered tracheal scaffold with refined mechanical properties and vascularisation for tracheal regeneration

Tehreem Khalid, Luis Soriano, Mark Lemoine, Sally-Ann Cryan, Fergal J. O'Brien and Cian O'Leary
120 Nanofibers with homogeneous heparin distribution and protracted release profile for vascular tissue engineering

Hongmei Zhang, Qilu Zhang, Juan Du, Tonghe Zhu, Dian Chen, Feiying Liu and Yang Dong

135 Engineered biomimetic micro/nano-materials for tissue regeneration

Feng Han, Qingchen Meng, En Xie, Kexin Li, Jie Hu, Qianglong Chen, Jiaying Li and Fengxuan Han
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OPEN ACCESS

EDITED BY
Yawei Du,
Shanghai Jiao Tong University, China

REVIEWED BY
Yiwei Li,
North Carolina State University,
United States
Wei Xiong,
Beijing Tongren Hospital, Capital
Medical University, China
Tuo Yang,
University of California, San Diego,
United States

*CORRESPONDENCE
Luis García-Fernández,
luis.garcia@csic.es

SPECIALTY SECTION
This article was submitted to
Biomaterials,
a section of the journal
Frontiers in Bioengineering and
Biotechnology

RECEIVED 30 September 2022

ACCEPTED 07 December 2022

PUBLISHED 19 December 2022

CITATION
Escalante S, Rico G, Becerra J, San Román J, Vázquez-Lasa B, Aguilar MR, Durán I and García-Fernández L (2022), Chemically crosslinked hyaluronic acid-chitosan hydrogel for application on cartilage regeneration.
*Front. Bioeng. Biotechnol.* 10:1058355.
doi: 10.3389/fbioe.2022.1058355

COPYRIGHT
© 2022 Escalante, Rico, Becerra, San Román, Vázquez-Lasa, Aguilar, Durán and García-Fernández. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.

# Chemically crosslinked hyaluronic acid-chitosan hydrogel for application on cartilage regeneration

Sandra Escalante¹,², Gustavo Rico¹,², José Becerra¹,², Julio San Román²,³, Blanca Vázquez-Lasa²,³, Maria Rosa Aguilar²,³, Iván Durán¹,² and Luis García-Fernández²,³*

¹Department of Cell Biology, Genetics and Physiology, Faculty of Science, University of Malaga, Malaga, Spain, ²Centro de Investigación Biomédica en Red de Bioingeniería, Biomateriales y Nanomedicina (CIBER-BBN), Instituto de Salud Carlos III, Madrid, Spain, ³Grupo de Biomateriales, Departamento de Nanomateriales Poliméricos y Biomateriales, Instituto de Ciencia y Tecnología de Polímeros (ICTP), CSIC, Madrid, Spain

Articular cartilage is an avascular tissue that lines the ends of bones in diarthrodial joints, serves as support, acts as a shock absorber, and facilitates joint's motion. It is formed by chondrocytes immersed in a dense extracellular matrix (principally composed of aggrecan linked to hyaluronic acid long chains). Damage to this tissue is usually associated with traumatic injuries or age-associated processes that often lead to discomfort, pain and disability in our aging society. Currently, there are few surgical alternatives to treat cartilage damage: the most commonly used is the microfracture procedure, but others include limited grafting or alternative chondrocyte implantation techniques, however, none of them completely restore a fully functional cartilage. Here we present the development of hydrogels based on hyaluronic acid and chitosan loaded with chondroitin sulfate by a new strategy of synthesis using biodegradable di-isocyanates to obtain an interpenetrated network of chitosan and hyaluronic acid for cartilage repair. These scaffolds act as delivery systems for the chondroitin sulfate and present mucoadhesive properties, which stabilizes the clot of microfracture procedures and promotes superficial chondrocyte differentiation favoring a true articular cellular colonization of the cartilage. This double feature potentially improves the microfracture technique and it will allow the development of next-generation therapies against articular cartilage damage.

KEYWORDS

hydrogel, cartilage regeneration, microfracture, hyaluronic acid, chitosan, chondroitin sulfate, diisocyanate
# 1 Introduction

Articular cartilage (AC) is an avascular connective tissue located in the diarthrodial joints covering the surface of bones, their main function is to provide a low friction surface for transmitting the mechanical load between bones (Adams et al., 2009).

AC is formed by chondrocytes distributed within an extracellular matrix (ECM). This ECM is composed of collagen, mainly of type II, glycosaminoglycans (GAGs) and proteoglycans (Mow et al., 1992). The main characteristics of this tissue, in addition to its peculiar structure and topography (Becerra et al., 2010), are its low metabolism and the absence of blood vessels, nerves and lymphatic systems. These features could be the reasons for its limited capability of repair and regeneration, especially for large defects (>3 mm²) (Doherty, 1992; Armiento et al., 2018).

Cartilage damage, one of the major skeletal morbidities in our society, is usually produced by repeating loads, sudden impact, foreign bodies or damage in other connective tissue, but also because of the natural degeneration due to the decrease of biological and biomechanical properties through aging. In both cases, the progressive degeneration of AC is the most common form of joint disease and represents one of the main causes of osteoarthritis (OA) (Lawrence et al., 2008).

Currently there are several options for the treatment of cartilage defects, but most of the current treatments are focused on symptom relief (Mollon et al., 2013):

* Non-surgical treatment: Oral analgesia, physiotherapy, intraarticular injection of GAGs, platelet rich plasma (PRP) and/or anti-inflammatory drugs (Pontes-Quero et al., 2019; García-Fernández et al., 2020).

* Arthroscopic chondroplasty: The damaged cartilage, loose bodies, and chondral fragments are removed (mechanical saver devices, radiofrequency ablation...) to produce a smooth chondral surface and decrease the mechanical joint irritation (Liptak and Theodoulou, 2017).

* Pridie Drilling or Microfracture (MF): Minimally invasive procedure that drills the subchondral bone at the defected site with the objective of releasing osteoprogenitor cells into the defect (Dehghani and Fathi, 2017).

* Autologous or allogenic osteochondral grafting: This technique uses allogenic or autologous osteochondral tissue to fill the defect (Dehghani and Fathi, 2017). Autologous Chondrocytes Implant (ACI): Implantation of chondrocytes coming from an enzymatically treated biopsy from patient cartilage (Mistry et al., 2017).

* Matrix Autologous Chondrocytes Implant (MACI): Similar to ACI, but the cells are growing in a matrix, and this matrix is implanted to the patient (Mistry et al., 2017).

A typical lesion of professional athletes and sport players exposed to great efforts consists on articular damage with an average size of around 2–3 mm² without affection of the subchondral bone. In this situation, the most used procedure is the microfracture (Sledge, 2001). Microfracture is an inexpensive technique, it presents a quick recovery time and reports positive result in pain relief and joint functionality (Mollon et al., 2013). This technique consists of the micro-perforation of the subchondral bone, allowing bleeding. The blood coming from the subchondral bone fills the damage and forms a clot with high levels of stem cells and growth factors (Frehner and Benthien, 2018). This clot acts as a scaffold for the development of the new cartilage. However, this technique often results in the formation of fibrocartilage with lower biomechanical properties than those of hyaline cartilage. Other problems with this procedure are the instability of the clot (resulting in a deficient integration with the surrounding tissue) and the fast rate of degradation (not allowing the complete regeneration of the cartilage) (Chung and Burdick, 2008).

Recent advances in cartilage repair focus on the use of scaffolds enhancing the biochemical and biomechanical properties of the cartilage regeneration process. In these studies, different biomaterials have been proposed as scaffolds for cartilage regeneration like gellan gum, polyglucosamine/glucosamine carbonate hydrogel or fibrin-hyaluronan matrix among others (Nehrer et al., 2008; Oliveira, 2010; Vilela et al., 2018; Pipino et al., 2019).

In this article, we present the development of an adhesive polymer scaffold based on hyaluronic acid (HA) and chitosan (Ch) to fit into a cartilage defect after a microfracture procedure with the capacity to stabilize the blood clot in the cartilage. HA is non-immunogenic natural polymer with a unique viscoelastic nature and one of the main components of the hyaline cartilage. (Chartrain et al., 2022). It is commonly used in intra-articular injections to stimulate and repair damaged cartilage because it promotes cell adhesion and proliferation and presents anti-inflammatory properties as well as non-immunogenicity (Shiroud Heidari et al., 2023). However, hydrogels made of HA present quick degradation rates, for this reason it is necessary to modify them to enhance their degradation properties. Ch is a cationic polysaccharide easy to obtain from crustacean shells. It is widely used as biomaterial due to its low toxicity and good biocompatibility and degradability (Shigemasa and Minami, 1996). Ch also present good adhesive properties to tissue and hemostasis, these properties allow a normal clot formation and impeding clot retraction (Shive et al., 2006; Chen, 2018; Pereira et al., 2018). Different studies demonstrate the efficacy of the marrow-derived mesenchymal stem cells (MSC) to migrate into the defect, proliferate, fill cavities and differentiate creating new cartilage (Shapiro et al., 1993; Anraku, 2009). Chondroitin sulfate (CS) is one of the principal GAGs presents in the cartilage ECM. CS present anti-inflammatory activity promotes cell differentiation and protect cartilage from ECM degradation due to regulation of the
metabolism of cartilage tissue (Lafuente-Merchan, 2022). This together with the fact that blood drained from subchondral bone contains MSC constitutes the hypothesis to study differentiation processes in our novel biomaterial, testing it on this stem cell population. Our results demonstrate that our biomaterial composed of HA, Ch and Chondroitin Sulfate (HAChCS) not only favors MSC settlement but also induces their differentiation into cartilage tissue highly resembling the articular cartilage in its surface.

# 2 Materials and methods

## 2.1 Materials and reagents

High molecular weight (800–100 kDa) sodium hyaluronate (HA, pharmaceutical grade) and sodium chondroitin sulfate from bovine (injectable grade) were kindly supplied by Bioiberica (Barcelona, Spain). Medical grade Ch of low molecular weight (260 kDa) and 90.5% of deacetylation degree was purchased from Altakitin (Lisbon, Portugal). L-lysine diisocyanate (97%) was purchased from BOC Sciences (NY, United States), acetic acid and Pluronic F-127 was purchased from Sigma-Aldrich (St. Louise, United States).

All products were used as received without further purification.

## 2.2 Preparation of hydrogel scaffolds

Hyaluronic acid-Chitosan scaffolds (HACh) were prepared by crosslinking with L-lysine diisocyanate (LDI). HA was dissolved in an aqueous solution of Pluronic F-127 (1% w/v) and acetic acid (1% v/v) at a concentration of 20 mg/ml. Pluronic F-127 was added as a surfactant to preserve the porosity until the hydrogel is formed. Ch was added in a 1:1 relationship and stirred at 37°C until complete dissolution. In the case of hydrogels containing chondroitin sulfate 4 mg/ml of CS were added at this point. The final solution was mixed with LDI in a NH₂/LDI relationship of 1:5, dispersed with an UltraTurrax at 15,000 rpm for a minute and placed on a Teflon mold. The dispersion was allowed to crosslink at 37°C in a humidified chamber for 1 h. The final scaffold was washed with distilled water and freeze-dried to obtain the final scaffold.

## 2.3 Physicochemical characterization

### 2.3.1 Scanning electron microscopy and energy dispersive spectroscopy analysis

The morphology of the scaffolds was analyzed by analytic scanning electron microscopy (SEM) using a JEOL JSM-6010LV (Tokyo, Japan). Prior analysis all samples were

sputter-coated with gold using a Leica EM ACE600 coater (Leica Microsystems, Wien, Austria). Elemental composition was performed by energy dispersive spectroscopy (EDS, Bruker XFlash model with detector 5030) installed in the SEM. Three independent areas were selected in the scaffolds.

### 2.3.2 Fourier transform infrared attenuated total reflectance (ATR-FTIR)

ATR-FTIR spectroscopy was performed in a Perkin Elmer Spectrum One FTIR spectrometer using 32 scans, and a resolution of 4 cm⁻¹.

### 2.3.3 Dynamic mechanical analysis

Rheology measurements were carried out in a stress-controlled oscillatory rheometer ARG2 (TA Instruments) using parallel plate geometry, using a 20 mm diameter steel plate. Oscillatory frequency sweeping test of scaffolds was conducted with a frequency scanning from 0.01 to 20 Hz at 0.1% strain and 25°C. Five replicates of each sample swollen in phosphate buffer were evaluated.

## 2.4 Swelling-degradation profile and chondroitin sulfate release

The gravimetric swelling ratio is defined as the fractional increase in the weight of the hydrogel due to water absorption. The gravimetric swelling capacity was measured in rounded scaffolds (d = 11 mm). The dried scaffolds were weighed (m₀) and then incubated in PBS at 37°C for 1 day. At regular intervals, the scaffolds were removed from the PBS solution, the excess of water was eliminated with filter paper and the samples were weighed (mₜ). The gravimetrically swelling percentage was calculated as:

$$ \text{Gravimetric Swelling (\%)} = 100 \times (m_t - m_0)/m_0 \quad (1) $$

The volumetric swelling ratio is defined as the volume increase of the hydrogel due to water absorption. The volumetric swelling capacity was measured in rounded scaffolds (d = 11 mm). The dimensions were measured (V₀) prior to incubation in PBS at 37°C for 1 day. At regular intervals, the scaffolds were removed from the PBS solution and the samples were measured (Vₜ). The volumetric swelling percentage was calculated as:

$$ \text{Volumetric Swelling (\%)} = 100 \times (V_t - V_0)/V_0 \quad (2) $$

The stability of HACh and HAChCS scaffolds was carried out by immersing the membranes in 5 ml of PBS solutions and incubating at 37°C. After different periods of time the samples were removed from de PBS, washed with distilled water to eliminate the remaining salts and freeze-dried to obtain the
weight of the sample at time t. The weight loss of the sample can be determined as:

$$Weight loss (\%) = 100 \times (m_i - m_t)/m_i \quad (3)$$

where $m_i$ is the initial weight of the sample and $m_t$ is the weight of the degraded sample at each time point.

For release experiments, scaffolds containing CS were incubated in PBS at 37°C. At regular intervals, the supernatant liquid was collected and analyzed by UV (Nanodrop One, Thermofisher). The absorbance value at 222 nm where compare with a calibration curve previously calculated to obtain the concentration of CS on the supernatant liquid.

## 2.5 Adhesion strength test

The bioadhesion properties of the scaffolds to bone tissue were tested using a modification of the ASTM F2258-05 method. Chicken keel bone were cut into homogeneous samples of 12 mm × 12 mm × 3 mm (length, width, thickness). To simulate the microfracture procedure five holes ($d = 0.1$ mm) were made on the surface of the bone. The corresponding scaffold were swollen in different conditions: PBS, fibrin glue (FG) or with chicken blood, and put in contact for 1 min with the bone. This time allow the coagulation of the fibrin glue or the formation of the clot. Adhesion strength was tested in a Universal Testing Machine (UTM, Instron model 3366) comparing drilled and non-drilled bones. The data were collected using a 100 N load cell and a loading rate of 5 mm/min.

## 2.6 In vitro toxicity and cell adhesion experiments

### 2.6.1 Cell cultures

Primary human osteoblast from femoral bone tissue (hOBs), and human articular chondrocytes from knee articular cartilage (hACs) (Innoprot, Derio, Spain) were used in this study. hOBs were grown in DMEM/F12 medium (Dulbecco's modified Eagle's medium/F12; Gibco®, Life Technologies, Carlsbad, United States) supplemented with 10% (v/v) fetal bovine serum (FBS; Life Technologies, Carlsbad, CA, United States) and 1% (v/v) penicillin/streptomycin solution (final concentration of penicillin 100 units/ml and streptomycin 100 mg/ml, Life Technologies, Carlsbad, CA, United States). Both cell lines were cultured until 90% of confluence at 37°C and 5% $CO_2$, changing the culture medium every 2-3 days to be ready for the different assays.

### 2.6.2 Cytotoxicity assay

An MTT test ((3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide)) was used to indirectly analyze

the cytotoxicity of the different scaffolds. Rounded scaffold ($d = 11$ mm) were immersed in 5 ml of FBS-free culture medium and incubated at 37°C. Thermanox® (TMX) discs (Nunc®, Thermofisher) were used as negative control. Aliquots of the supernatant medium were taken at 1, 2, 7, 14 and 21 days under sterile conditions and replaced with fresh medium. The samples were kept frozen until use. hObs and hACs were seeded separately in a 96-well plate at $9 \times 10^4$ cells/ml of density and incubated for 24 h in complete medium. Then, the medium was substituted by the extracts and incubated at 37°C. After 24 h the medium was replaced by a solution of MTT (0.5 mg/ml) in warm FBS-free medium and the cells were incubated at 37°C for 3-4 h. Mediua containing MTT were removed and DMSO was added to dissolve the formazan crystals formed in the cells. The absorbance of the medium was measured with a Biotek Synergy HT detector at 570 nm and a reference wavelength of 630 nm. Cell viability was calculated as:

$$Cell Viability y (%) = 100 \times (OD_S - OD_B)/(OD_C - OD_B) \quad (4)$$

where $OD_S$, $OD_B$ and $OD_C$ are the optical density (OD) of formazan production for the sample, blank and control, respectively.

### 2.6.3 Cell adhesion and proliferation assay

Cell adhesion and proliferation on the scaffolds were tested by the Alamar Blue assay following the manufacturer's instructions. Cells were seeded at $4 \times 10^5$ cell/ml on the previously swollen scaffolds and incubated at 37°C in a humidified atmosphere with 5% $CO_2$. Cell proliferation was measured at determinate times (1, 4, 7, 14 and 21 days) by replacing the culture medium with a 10% solution of AB in phenol red free DMEM medium. After 4 h the solution was transferred into a 96-well plate and the fluorescence was measured in a microplate reader (Synergy HT, BioTek, Instruments, United States). The excitation wavelength was 530 nm and the emission was recorded at 590 nm.

## 2.7 In vitro evaluation of cartilage formation on HACh scaffolds

### 2.7.1 Cell culture

C3H10T1/2 cell line (mouse: ATTC, CCL-226) was cultured in 75-cm² polystyrene culture flask in Dulbecco's Modified Eagle's Medium—low glucose (Sigma, D5546-1X500ML) supplemented with 10% fetal bovine serum (GIBSO, F7524-500ml), 1% Penicillin-Streptomycin (Sigma, P4333-100ML) and 1% Glutamax (GIBCO, 35050-038) in a humidified incubator at 37°C with 5% $CO_2$. Passages were performed at 80% confluence with Trypsin-EDTA 0.05% (Thermofisher, 25300096) and cells were always used in the 15th passage.
## 2.7.2 Scaffold seeding

Ch, HACH and HACHCS disks were rinsed in 70% ethanol for 20 min in agitation and punched into cylinders of 3 mm diameter. The pieces were sterilized in 70% ethanol for 2 h, rinsed in PBS for 30 min three times and maintained in culture medium at 4°C until use.

For the scaffold seeding, the cylinders obtained from each material were left at room temperature for 1 h, partially dried with a cell strainer (DICSA) and placed on a non-treated culture petri dish. C3H10 cells were trypsinized, resuspended in culture media and $1 \times 10^5$ cells were seeded in each 3 mm cylinders of biomaterial. The scaffolds were incubated at 37°C and 5% CO2 for 2 h to allow adhesion, adding 3 µl of culture media every 30 min. Then, the cylinders were placed in the center of a well in a standard 48-well plate (Jet BioFil) and incubated in culture media in a humidified incubator at 37°C with 5% CO2 for 1 and 2 weeks. Culture media were replaced every 2-3 days the first week and 3-4 days the second week. Micromass of C3H10 cells was also obtained as controls by centrifugation of $1 \times 10^5$ cells at 1.800 rpm at 5 min, removing the supernatant and adding 250 µl of culture media. Micromass was kept under the same conditions as Ch scaffolds. To stimulate differentiation, the biomaterials were incubated in culture media containing 100 ng/ml BMP-2 (R&D Systems, 355-BM-050) at 37°C in a humidified incubator with 5% CO2 for 1 and 3 weeks. Culture media were replaced every 2-3 days the first week and 3-4 days the second week.

## 2.7.3 Histology

After the differentiation assays, the scaffolds were rinsed with PBS and fixed with 4% paraformaldehyde at 4°C overnight. Then, they were washed in PBS, dehydrated with a graded series of ethanol, embedded in paraffin and sectioned at 15 µm thick. For histological analysis, sections were rehydrated and stained with Alcian Blue and Safranin-O.

## 2.7.4 Immunohistochemistry

For immunohistochemistry, sections were treated with citrate buffer for antigen retrieval and quenched by peroxidase solution. After blocking with goat serum, Aggrecan antibody (13880 Proteintech) was incubated overnight 1:100. Sections were developed with histostain plus kit with DAB as a chromogen (Invitrogen). Negative controls for primary antibody were included. Briefly, control slides were exposed to antigen retrieval and maximum DAB development to confirm no unspecific signal development (Supplementary Figure S3).

## 2.7.5 Real-time polymerase chain reaction (qPCR)

Total RNA from the cultured scaffolds was extracted using the Trizol reagent (ThermoFisher, 15596018) and RNA concentrations were determined using Nanodrop 2000 spectrophotometer (Thermo Fisher Scientific).

Complementary DNA (cDNA) was amplified with the PrimeScript™ RT Master Mix (Takara, RR036A) and used for qPCR, which was performed in triplicates by using the TB Green® Premix Ex Taq™ (Takara, RR420L) on the Bio-Rad C1000 Thermal Cycler (Bio-Rad). The housekeeping gene, β-actin, was used for gene expression normalization and the fold change expression was calculated using the $2^{-\Delta ACT}$ method. Primer sequences were purchased from Invitrogen and listed in Table 1.

## 2.8 Statistical analysis

Results are given as mean and standard deviation (minimum n = 4). Data from the different groups were compared in pairs with ANOVA test. All the statistical analyses were performed using the Origin 9 (Origin Lab corporation, Northampton, United States). The significance level was set at *p* < 0.05, **p** < 0.01 and ***p**<0.001.

# 3 Results and discussion

## 3.1 Scaffold synthesis and characterization

The use of lysine diisocyanates as crosslinker allows different reactions between Ch and HA as is shown in Figure 1A.

The presence of the different reactions could be elucidated by ATR-FTIR spectra of HA, Ch and HACH scaffolds (Figure 1B). HA spectrum showed the most characteristics bands: the broad band around $3330 \text{ cm}^{-1}$ is associated with the intra- and intermolecular stretching vibration of -OH group. The signal at $2,875 \text{ cm}^{-1}$ correspond to the stretching of -CH₂ group; the bands from 1,610 and $1,405 \text{ cm}^{-1}$ are correlated with symmetric and asymmetric vibration of COO⁻ group. Finally, the peak from $1,033 \text{ cm}^{-1}$ is related with the C-O-C from the saccharide units (Vasi, 2014).

Ch spectrum showed also their typical bands: The strong bands at 3369 and 3295 $\text{cm}^{-1}$ correspond to N-H and O-H stretching. The bands at 2,921 and $2,867 \text{ cm}^{-1}$ correspond to C-H stretching (symmetric and asymmetric). At $1,654 \text{ cm}^{-1}$ correspond to C=O stretching of amide I and at $1,589 \text{ cm}^{-1}$ correspond to the N-H bending of the primary amine. The absorption band at $1,153 \text{ cm}^{-1}$ can be attributed to asymmetric stretching of the C-O-C bridge.

Analyzing the HACH spectrum, the signal of HA at 1,610 and $1,405 \text{ cm}^{-1}$ (stretching of carboxylate group) was vanished, indicating that the conjugate was formed through the $COO^{-}$ group. In the case of the Ch, the signal corresponding to the primary amine ($1,589 \text{ cm}^{-1}$) also disappears in the HACH spectrum, indicating the formation of the crosslinking between the amine and isocyanate group. In the HACH spectrum, new signals appear at 1,550 and $1725 \text{ cm}^{-1}$, which
<table><caption>TABLE 1 Sequences of qRT-PCR primers.</caption><thead><tr><td>Gene</td><td>Forward</td><td>Reverse</td></tr></thead><tbody><tr><td>β-actin</td><td>5'-GGAGATTACTGCCCTGGCTCCTA-3'</td><td>5'-GACTCATCGTACTCCTGCTTGCTG-3'</td></tr><tr><td>Sox9</td><td>5'-GAGGCCACGGAACAGACTCA-3'</td><td>5'-CAGCGCCTTGAAGATAGCATT-3'</td></tr><tr><td>Acan1</td><td>5'-GGTCACTGTTACGCCACTT-3'</td><td>5'-CCCCTTCGATAGTCCTGTCA-3'</td></tr></tbody></table>

<figure><img src="image_4.png" /></figure>

are in the range of the secondary amine bands and C=O stretching in polyurethanes and polyureas. These results demonstrate the formation of a stable scaffold.

The use of the UltraTurrax in the synthesis process produced scaffolds with a high porosity due to the introduction of microbubbles in the system (Figure 2).

In HAcH scaffolds (Figures 2A1, A2), we can observe the presence of two different kinds of microstructures: a macroporosity (≈200 µm) and a microporosity (≈10 µm). The microporosity is partially occulted on CS loaded scaffolds (Figures 2B1, B2). In this case, CS is deposited on the microstructure partially covering it. The presence of CS on the surface was corroborated using EDS (Supplementary Table S1) that detected the presence of sulfur coming from the CS. This porosity pattern will allow the colonization of the scaffold by the cells contributing to the full regeneration of the damage.

Mechanical properties of HAcH and HAcHCS swollen scaffolds were analyzed by rheology (Supplementary Figure S1A). Storage and loss moduli exhibited a visco-elastic solid behavior close to a gel-like. HAcHCS scaffolds showed lower storage moduli values in comparison with not loaded scaffolds.

Previous studies define the behavior of articular cartilage using the complex modulus (G*) and define different ranges: immature cartilage: G* < 1 MPa and mature cartilage 5 < G*< 16 MPa at low frequencies (Perni and Prokopovich, 2020). The

values obtained for our scaffold (Supplementary Figure S1B) are in the range of an immature cartilage (0.5–0.8 MPa) (Perni and Prokopovich, 2020) being adequate for the regeneration of new tissue when it is applied in articular defects currently treated with microfracture procedure.

## 3.2 Swelling, degradation and CS release

Liquid adsorption capacity is an essential feature and determine the ability of the scaffold to interact with the blood coming from the microfracture process. This adsorption needs to be fast to stabilize the blood clot and adsorb the growth factors and stem cells coming from the blood. Figure 3A shows the swelling capacity of the HAcH and HAcHCS scaffolds. The maximum water-uptake equilibrium is reached after 4–6 h, reaching values of 225% for HAcH and 200% for HAcHCS. High initial swelling degrees are necessary for ensuring an appropriate source of nutrients to the whole scaffold, but too high swelling degrees can compromise scaffold integrity (Mora-Boza, 2020).

Volumetric swelling is an important scaffold characteristic to determine the ability of the gel to remain inside the microfracture holes. High volume changes could affect the stability of the gels once they are implanted in the defect. In our case (Figure 3B), the
<figure><img src="image_5.png" /><figcaption>FIGURE 2
SEM images of (A) HAcH and (B) HAcHCS at different magnifications: (A1,B1) ×100 and (A2,B2) ×500.</figcaption></figure>

volume increase values are only 27% for HAcH scaffolds and 17% for HAcHCS scaffolds, being the maximum variation in the diameter of the scaffold of 0.55 ± 0.08 mm which ensures the permanence of our hydrogels in the defect.

Figure 3C shows the degradation profile of the scaffolds submerged in PBS at 37°C. Both scaffolds showed an initial weight loss due to the release of not integrated polymer chains and then they were stable for more than 2 months. The HAcHCS presented a more sustained degradation may due to polyelectrolyte interactions between chondroitin sulfate and Ch (Rodrigues et al., 2016). Figure 3D shows the release of chondroitin sulfate from the HAcHCS scaffold. The release profile shows an initial fast release of CS (10%) that was progressively increasing with the time until reaching a value of 40% after 40 days. This behavior is typical for systems that are capable of swelling and whose release profile fit the Korsmeyer-Peppas model Eq. 5:

$$ \frac{M_t}{M_{\infty}} = k \cdot t^n \qquad (5) $$

were $\frac{M_t}{M_{\infty}}$ is the fraction of CS released at time t, k is the release rate constant and n is the diffusional exponent which is indicative of the transport mechanism (Ritger and Peppas, 1987). Assuming

our scaffold as a thin film we fit our data to Eq. 5 obtaining a value of n = 0.4982. Table 2 shows model values for n of the Peppas equation and that of the HAcHCS sample.

Then, our system fit to thin film were the CS is released by Fickian diffusion from the swollen matrix. The CS that is homogeneously distributed in the polymer matrix is released by diffusion through the pores of the matrix.

## 3.3 Adhesion strength test

Ch is a natural mucoadhesive and could be a good option to stabilize the scaffold in the microfracture procedure. There are several options to stabilize the scaffold as can it be the use of fibrin glue or surgical adhesives (Frehner and Benthien, 2018; Orth et al., 2020). To measure the mucoadhesive properties of our scaffolds we put in contact the scaffold with bone tissue in different conditions. Figure 4 displays the values of adhesion strength of the scaffolds to bone:

The adhesion of the scaffolds to the bone in absence of microfracture was practically negligible even when fibrin glue was applied. After the application of the microfracture to the bone, the adhesion strength noticeably increased by the
<figure><img src="image_6.png" /></figure>

<figure><img src="image_7.png" /></figure>

<figure><img src="image_8.png" /></figure>

<figure><img src="image_9.png" /><figcaption>**FIGURE 3**
(A,B) Swelling behavior of HAcH and HAcHCS scaffolds in PBS at 37°C. (C) Degradation of HAcH and HAcHCS scaffolds in PBS at 37°C. (D) Chondroitin sulfate released from HAcHCS scaffolds in PBS at 37°C.</figcaption></figure>

<table><caption>TABLE 2 Experimental and theoretical exponent n of the Peppas equation for drug release mechanism.</caption><thead><tr><th>HAcHCS</th><th>Thin film</th><th>Cylinder</th><th>Sphere</th><th>Drug release mechanism</th></tr><tr><td></td><td colspan="3">Exponent, n</td><td></td></tr></thead><tbody><tr><td>0.4982</td><td>0.5</td><td>0.45</td><td>0.43</td><td>Fickian diffusion</td></tr><tr><td></td><td>0.5 &lt; n &lt; 1.0</td><td>0.45 &lt; n &lt; 0.85</td><td>0.43 &lt; n &lt; 1.45</td><td>Anomalous transport</td></tr><tr><td></td><td>1.0</td><td>0.89</td><td>0.85</td><td>Case-II transport</td></tr></tbody></table>

augmentation of the contact surface due to the bone holes. In addition, there was also observed an important difference between the use of fibrin glue or blood to seal the scaffold.

The adhesion strength was significantly higher in the presence of blood. Moreover, the presence of Ch could accelerate blood coagulations, platelet adhesion and thrombin generation
<figure><img src="image_10.png" /><figcaption>**FIGURE 4**
Adhesion strength of the scaffolds to a bone in different conditions. FG: the scaffold was swollen with fibrin glue. MF: A microfracture procedure was applied to the bone. Blood: The scaffold was swollen with blood.</figcaption></figure>

resulting in an increase of the adhesion strength (Hu, 2018). This improvement of the adhesion strength makes this kind of systems good candidates to stabilize the clot coming from the microfracture process in the treatment of cartilage injuries.

## 3.4 In vitro toxicity and cell adhesion studies

The toxicity of lixiviates coming from the different scaffolds was evaluated on hACs and hOBs (Figures 5A, B). Cell viability of studied samples with hACs (Figure 5A) was not affected on the first 2 days in comparison with the control. After 7 days, significant differences in the viability were observed but in general, the cell viability in presence of any extract was higher than 90%. In the case of hOBs (Figure 5B) cell viability increased in contact with scaffold lixiviates. Ho (2014) observed that the presence of low molecular weight chains of Ch and/or HA in the culture medium improved the cell growth.

Cell adhesion and proliferation on the different scaffolds was evaluated by Alamar Blue assay (Figures 5C, D, Supplementary Figure S2). Cell proliferation increased over time for all the systems but was lower on the scaffolds that on the control. This difference can be ascribed to the surface of the scaffolds, the cells need more time to adhere on an irregular surface, but

<figure><img src="image_11.png" /></figure>

<figure><img src="image_12.png" /></figure>

<figure><img src="image_13.png" /></figure>

<figure><img src="image_14.png" /><figcaption>**FIGURE 5**
(A,B) Cell viability of hACs (A) and hOBs (B) cultures with HAcH and HAcHCS scaffolds lixiviates. Significant differences with the control at each time are marked with * (*p* < 0.05, **p** < 0.01, ***p** < 0.001) (C,D) Cell proliferation values of hACs (C) and hOBs (D) seeded on HAcH and HAcHCS scaffolds. Significant differences with the control (Thermanox® discs) at each time are marked with * (*p* < 0.05, **p** < 0.01, ***p** < 0.001) and between scaffolds are marked with + (*p* < 0.05, **p** < 0.01, ***p** < 0.001).</figcaption></figure>
A

Sox9

1 week

<figure><img src="image_15.png" /></figure>

B

Acan

<figure><img src="image_16.png" /></figure>

C

Sox9

2 weeks

<figure><img src="image_17.png" /></figure>

D

Acan

<figure><img src="image_18.png" /></figure>

E

<figure><img src="image_19.png" /><figcaption>**FIGURE 6**
Expression level of Sox9 (A,C) and Acan (B,D) genes in MSC seeded as spheroids micromass (S) or on scaffolds: Ch, HACH and HACHCS during 1 week (A,B) or 2 weeks (C,D). Results are expressed as mean ± standard deviation (*p* < 0.05, **p** < 0.01, ***p** < 0.001). (E): Alcian Blue stained sections of Ch, HACH and HACHCS scaffolds seeded with MSC during 1 week. Bars represent 500 µm.</figcaption></figure>
<figure><img src="image_20.png" /><figcaption>**FIGURE 7**
(A,B): Expression level of Sox9 and Acan genes in MSC seeded on Ch without BMP-2 stimulation (Ch-) and Ch, HAcH and HAcHCS scaffolds after 3 weeks of differentiation with BMP-2 stimulation. Results are expressed as mean ± standard deviation (\*p < 0.05, \*\*p < 0.01, \*\*\*p < 0.001). (C): Alcian Blue stained sections of Ch, HAcH and HAcHCS scaffolds with MSC and BMP-2 stimulation for 3 weeks. Bars represent 500 µm (D): High magnification showing well differentiated chondrocytes on the Ch, HAcH and HAcHCS scaffolds. Bars represent 50 µm. (E): Safranin-O stained sections of Ch, HAcH and HAcHCS scaffolds with MSC and BMP-2 stimulation. Bars represent 500 µm (F): High magnification showing safranin-O staining on the Ch, HAcH and HAcHCS scaffolds. Bars represent 50 µm. (G) ACAN (Aggrecan) immunohistochemical analysis of paraffin-embedded sections of Ch, HAcH and HAcHCS scaffolds seeded with MSC and BMP-2 stimulation. Bars represent 500 µm (H) High magnification showing ACAN (Aggrecan) immunohistochemical analysis on the Ch, HAcH and HAcHCS scaffolds. Bars represent 50 µm.</figcaption></figure>

after the first 7 days the cell growing on the scaffold surface significantly increase with the time. In SEM pictures (Supplementary Figure S2) we can observe some groups of cells on the surface of the scaffold after 2 days, but after 21 days, the scaffold is completely cover by both tupes of cells. Also in both kinds of cells, the presence of CS on the scaffold improve cell proliferation because CS is one of the main components in ECM and is involved on cell differentiation and proliferation processes (Izumikawa et al., 2014).

## 3.5 Cartilage differentiation on HA based scaffolds

To determine the true potential of our new scaffold in cartilage repair we performed different *in vitro* articular cartilage differentiation assays. The two most common articular cartilage repair interventions (drilling or microfracture surgical procedures), infuse the damaged zone with MSC. We decided to test the differentiation capabilities of our novel biomaterial using a MSC line with and without BMP2 stimulation. This is a reliable way to induce chondrocyte differentiation from MSC lineages (Shea, 2003). We determined the cartilage formation capacity on the three biomaterial combinations (Ch, HAcH and HAcHCS) and compared to that of a scaffold-free control generating MSC spheroids, a standard protocol for cartilage differentiation (Denker et al., 1995). Then, we measured two aspects of chondrogenesis by quantitative gene expression: 1) predifferentiation towards chondrogenesis measuring Sox9 expression; 2) actual cartilage ECM production measuring aggrecan (Acan) transcription, one of the major components of the cartilage ECM.

First, we tested the MSC reaction to the different scaffolds to determine if the biomaterials could be pro-chondrogenic activity by themselves without external differentiation factors. Short-term experiments for 1 week of differentiation showed that HAcHCS had the best performance giving the highest expression of Sox9, the gene marker for early stages of chondrogenic differentiation, compared even to the standard spheroid method of cartilage differentiation (Figure 6A). Regarding cartilage ECM production, HAcHCS also showed higher Acan expression levels than Ch or HAcH, with similar performance to standard spheroids (Figure 6B). Mid-term culture of MSC on tested scaffolds during 2 weeks showed similar levels for chondrogenesis marker Sox9 for all biomaterials (Figure 6C). However, at 2 weeks cartilaginous ECM expression significantly increased in HAcHCS scaffold (Figure 6D). Histological analysis by Alcian Blue revealed more production of proteoglycans, a major component of
cartilage extracellular matrix, and presence of more cells on HAChCS compared to Ch scaffold (Figure 6E). This was confirmed by histological analysis with Alcian Blue, a high affinity staining to cartilaginous proteoglycans, which preferentially binds to cartilaginous ECM (Figure 6E). These results suggest that HA based scaffolds, and more specifically, HAChCS biomaterial inherently induces chondrogenesis in MSC. These results could be due to the release of CS from the scaffolds. CS has shown capacity to activate MSC chondrogenesis and neocartilage deposition (Wang and Yang, 2017).

Even if scaffolds are prochondrogenic, to produce fully differentiated chondrocytes, cell 3D culture systems need differentiation factors such as BMP-2. To reproduce the microfracture procedure and the articular cartilage environment in the biomaterial, we stimulated MSC with BMP-2 (one of the factors provided by the blood coming from the microfracture of the subchondral bone) allowing cartilage differentiation during 3 weeks (Figure 7). Under these conditions, Sox9 chondrogenic marker showed stabilized levels similar to unstimulated culture after 2 weeks (Figures 6C, 7A), as expected for an early differentiation factor. On the other hand, Acan expression showed higher levels on HA based materials; especially the HAChCS scaffold (Figure 7B. Histological analysis by Alcian Blue (Figures 7C, D) and Safranin-O (Figures 7E, F) showed a production of proteoglycans own of mature articular cartilage (Figure 7C) correlated with well-differentiated chondrocytes with a preferential location on the scaffold surface (Figure 7). Immunostaining of aggrecan, a chondrocyte specific component of cartilaginous extracellular matrix, showed a tissue specific deposition of cells in contact with the HAcH and the HACHCS biomaterials, mainly in the surface region of the engineered tissue, confirming its similarities to articular cartilage (Figures 7G, H).

These results suggest that HAcH and HACHCS biomaterials promoted in vitro chondrogenic differentiation process, as shown by chondrogenic expression markers and cartilage-like tissue formation, and it is expected that they do it in the case of their use in a microfracture procedure.

## 4 Conclusion

Microfracture procedure is a common technique used in cartilage diseases but present several limitations. The use of HACHCS scaffolds on this process could improve the regeneration of articular cartilage. HACHCS presents physico-chemical properties similar to immature cartilage that could allow the regeneration and formation of articular cartilage. Also, the mucoadhesive properties could stabilize the clot coming from the microfracture process and improve the integration with the surrounding tissues. HACHCS did

not present toxicity and allowed the adhesion and proliferation of articular cartilage and mesenchymal stem cells. The capacity to release CS activated chondrogenesis process on MSC and the production of proteoglycans without the necessity of any other supplements. Finally, we simulated the environment of a microfracture procedure by addition of BMP-2, one of the common compounds present on the subchondral blood. In these conditions, both HAcH and HACHCS scaffolds promoted the chondrogenic differentiation process and cartilage-like tissue formation doing these systems good candidates for their use in the microfracture procedure as well as a scaffold for stem cell and chondro-lineages in next-generation therapies.

## Data availability statement

The raw data supporting the conclusion of this article will be made available by the authors, without undue reservation.

## Author contributions

SE: Data curation. Formal analysis. Investigation. Methodology. GR: Data curation. Formal analysis. Investigation. Methodology. Writing—review and editing. JB: Funding acquisition. Resources, Supervision, Writing—review and editing. JSR: Conceptualization, Funding acquisition. Resources, Supervision, Writing—review and editing. BV-L: Funding acquisition. Resources, Supervision, Writing—review and editing. MRA: Funding acquisition. Resources, Supervision, Writing—review and editing. ID: Conceptualization, Formal analysis, Investigation, Methodology, Supervision, Validation, Visualization, Writing—original draft. LG-F: Conceptualization, Data Curation, Formal analysis, Funding acquisition, Investigation, Methodology, Project administration, Supervision, Validation, Visualization, Writing—original draft.

## Funding

This work was supported by the Ministry of Science, Innovation and Universities (Spain) (PID 2020-114086RB-I00 and PID 2020-117255RB-I00-VASCUBONE) and by the Biomedical Research Networking Center in Bioengineering, Biomaterials and Nanomedicine (CIBER-BBN, project name: HYDROCARTREG). This research work was performed in the framework of the Nanomedicine CSIC HUB (ref202180E048) and BV-L, LG-F and MRA are members of the SusPlast + platform from the CSIC.
## Conflict of interest

The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

## Publisher's note

All claims expressed in this article are solely those of the authors and do not necessarily represent those of their affiliated

organizations, or those of the publisher, the editors and the reviewers. Any product that may be evaluated in this article, or claim that may be made by its manufacturer, is not guaranteed or endorsed by the publisher.

## Supplementary material

The Supplementary Material for this article can be found online at: https://www.frontiersin.org/articles/10.3389/fbioe.2022.1058355/full#supplementary-material

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<figure><img src="image_21.png" /></figure>

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OPEN ACCESS

EDITED BY
Yawei Du,
Shanghai Jiao Tong University, China

REVIEWED BY
Jinglei Wu,
Donghua University, China
Lichen Wang,
Stevens Institute of Technology,
United States

*CORRESPONDENCE

Lihua Luo,
luolihua81@126.com

Yan He,

helen-1101@hotmail.com

Qingsong Ye,

qingsongye@whu.edu.cn

SPECIALTY SECTION

This article was submitted to
Biomaterials,
a section of the journal
Frontiers in Bioengineering and
Biotechnology

RECEIVED 06 October 2022

ACCEPTED 19 December 2022

PUBLISHED 06 January 2023

CITATION

Luo L, Zhu Q, Li Y, Hu F, Yu J, Liao X, Xing Z, He Y and Ye Q (2023), Application of thermosensitive-hydrogel combined with dental pulp stem cells on the injured fallopian tube mucosa in an animal model.
*Front. Bioeng. Biotechnol.* 10:1062646.
doi: 10.3389/fbioe.2022.1062646

COPYRIGHT

© 2023 Luo, Zhu, Li, Hu, Yu, Liao, Xing, He and Ye. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.

# Application of thermosensitive-hydrogel combined with dental pulp stem cells on the injured fallopian tube mucosa in an animal model

Lihua Luo<fnref n="1" />*†, Qunyan Zhu<fnref n="1" />†, Yejian Li<fnref n="1" />†, Fengting Hu<fnref n="1" />,
Jiangtao Yu², Xiangyan Liao<fnref n="1" />, Zhenjie Xing<fnref n="1" />, Yan He³* and
Qingsong Ye<fnref n="1" />,⁴*

**Objectives:** Fallopian tube (FT) injury is an important factor that can lead to tubal infertility. Stem-cell-based therapy shows great potential for the treatment of injured fallopian tube. However, little research has shown that mesenchymal stem cells (MSCs) can be used to treat fallopian tube damage by *in situ* injection. In this study, we *in situ* transplanted PF127 hydrogel encapsulating dental pulp stem cells (DPSCs) into the injured sites to promote the repair and regeneration of fallopian tube injury.

**Materials and methods:** The properties of dental pulp stem cells were evaluated by flow cytometry, immunofluorescence analysis, and multi-differentiation detection. The immunomodulatory and angiogenic characteristics of dental pulp stem cells were analyzed on the basis of the detection of inflammatory factor expression and the formation of capillary-like structures, respectively. The biocompatibility of PF127 hydrogel was evaluated by using Live/Dead and CCK-8 assays. The effects of PF127 hydrogel containing dental pulp stem cells on the repair and regeneration of fallopian tube injury were evaluated by histological analysis [e.g., hematoxylin and eosin (H&E) and Masson's trichrome staining, TUNEL staining, immunofluorescence staining, and immunohistochemistry], Enzyme-linked immunosorbent assay (ELISA), and RT-PCR detections.

**Results:** Dental pulp stem cells had MSC-like characteristics and great immunomodulatory and angiogenic properties. PF127 hydrogel had a thermosensitive feature and great cytocompatibility with dental pulp stem cells. In addition, our results indicated that PF127 hydrogel containing dental pulp stem cells could promote the repair and regeneration of fallopian tube damage by inhibiting cell apoptosis, stimulating the secretion of angiogenic factors, promoting cell proliferation, modulating the secretion of inflammatory factors, and restoring the secretion of epithelial cells.

<footnote n="1">¹These authors have contributed equally to this work</footnote>

<footnote n="1">¹School and Hospital of Stomatology, Wenzhou Medical University, Wenzhou, Zhejiang, China, ²The First Affiliated Hospital of Wenzhou Medical University, Wenzhou, Zhejiang, China, ³Tianyou Hospital, Wuhan University of Science and Technology, Wuhan, China, ⁴Center of Regenerative Medicine, Renmin Hospital of Wuhan University, Wuhan, China</footnote>
**Conclusion:** In this study, our results reported that *in situ* injection of PF127 hydrogel encapsulating dental pulp stem cells into the injured sites could provide an attractive strategy for the future treatment of fallopian tube injury in clinical settings.

KEYWORDS

dental pulp stem cells, fallopian tube injury, hydrogel, immunoregulation, repair

# 1 Introduction

The World Health Organization has reported that the infertility rate is more than 15% among women of reproductive age worldwide, and infertility should be regarded as a public health problem (Adegbola and Akindele 2013). Tubal infertility is a common cause of female infertility, accounting for 30%–35% (Zou et al., 2014). Many influencing factors, including infection, fallopian tube (FT) surgery, ectopic pregnancy, and endometriosis, may cause salpingitis or pelvic inflammation with FT mucosa damage and structural collapse, which are important indicators of tubal infertility (Li et al., 2017). At present, tubal infertility shows a yearly rising trend because of the popularity of sexually transmitted diseases and the repetitive infection of FT (Luo et al., 2015; Almasry et al., 2018).

Recently, the principal curing methods for tubal infertility include laparoscopic surgery and antibiotic therapy in clinical settings. However, the repetitive attachment of FT after surgery and the occurrence of antibiotic-resistant strains of the long-term use of antibiotics can lead to failure treatment of tubal infertility (Liao et al., 2019). Therefore, new therapeutic strategies have been required to repair the injury of FT and restore the reproduction function of FT, thereby reducing the incidence of tubal infertility.

Mesenchymal stem cells (MSCs) have pluripotent characteristics, and they can be induced to differentiate into many types of cells, which could provide tissue-reconstructing cells in the damaged sites. In addition, MSCs have shown strong paracrine effects by secreting several growth factors and providing immunosuppressive and anti-inflammatory actions in stem-cell-based therapies (Liao et al., 2019). Thus, MSC-based therapy can provide a novel strategy for the repair and regeneration of injured FT.

Dental pulp stem cells (DPSCs), as a kind of MSCs, can be readily cultured from the dental pulps which are isolated from impacted third molars, exfoliated deciduous teeth, permanent and primary teeth, and supernumerary teeth, without extra harm and invasive surgical procedures (Luo et al., 2018; Luo et al., 2021). DPSCs possess all MSCs-like properties, e.g. multi-differentiation potential and immunomodulatory properties, and high proliferative and self-renewal capacities (Luo et al., 2018). In addition, the capacity of DPSCs to secrete growth factors is higher than other tissue-derived MSCs such as bone-marrow-derived mesenchymal stem cells (BM-MSCs).

Meanwhile, DPSCs have a higher proliferation rate and greater immunomodulatory properties than BM-MSCs (Lan et al., 2019). Therefore, DPSCs have become an attractive seeded cell in stem-cell-based therapies. Moreover, we have confirmed an angiogenic characteristics of DPSCs' in our previous work (Luo et al., 2018), where DPSCs participated in the formation of new blood vessels and enhanced the blood supply around injured sites. Angiogenesis is an important influencing factor for the treatment of FT injury.

Although DPSCs have outstanding MSC-like properties, using DPSCs alone can hardly receive ideal restoration of reproduction function after FT injury because of their poor cellular density and survival rate in damaged sites. As shown in previous studies, hydrogels have a 3D network structure and great porosity, which can be used to accommodate cells and provide a suitable microenvironment for cell survival and growth (Zhu and Marchant 2011). Our previous studies indicated that the thermosensitive Pluronic F-127 (PF127, also known as poloxamer 407) hydrogel, which has a solution state at 4°C and a hydrogel state at human body temperature, had a porous structure and great cytocompatibility, and it could be a great cell carrier in tissue engineering and regenerative medicine (Albashari et al., 2020). Furthermore, the controllable degradability of PF127 hydrogel could provide an appropriate microenvironment for cellular adhesion, growth and proliferation (Supplementary Figure S5). Thus, in this work, we used PF127 as the scaffold and transplanted PF127 hydrogel containing DPSCs into the damaged site to evaluate its effects on the repair and regeneration of FT injury.

# 2 Materials and methods

## 2.1 Ethics and animals

Twenty-four female New Zealand white rabbits (weighing 2,000 ± 250 g) were provided by the Animal Experimental Center of Wenzhou Medical University and housed with sufficient water and food in animal cages for 14 days before surgery. All experimental procedures were performed under the regulations of the committee of the Institutional Animal Care and the Ethics Committee of Wenzhou Medical University (wydw 2019-0949).
## 2.2 Isolation, culture, and identification of DPSCs

The isolation and culture procedures of DPSCs were performed as previously described (Luo et al., 2021). In brief, dental pulp tissues, which were collected from the tooth, were cut into small pieces and digested using the mixture of collagenase type I (Gibco, United States) and dispase (Sigma, Germany) for 30 min at 37°C in a 5% CO₂ cell incubator. Then, the pulp tissue was suspended and cultured using α-modified Eagle's medium (α-MEM, Gibco, United States) supplemented with 20% fetal bovine serum (FBS, Gibco, United States), 100 µg/ml of streptomycin, and 100 U/mL of penicillin (Gibco, United States) at 37°C in a 5% CO₂ cell incubator. The culture medium was replaced every 3 days. MSC-like characteristics of DPSCs were identified by flow cytometry, immunofluorescence analysis, and multi-differentiation detection. Flow cytometry was executed using human primary antibodies, including CD19, CD166, CD73, CD90 (BD Pharmingen™, United States), CD34, and HLA-DR (BioLegend, United States), and the data were evaluated using a CytoFLEX flow cytometer (Beckman Coulter, California, United States). The immunophenotype of DPSCs was analyzed using primary antibodies against CD44 (1:400, Abcam, Cambridge, United Kingdom), and the immunofluorescence of DPSCs was observed using a fluorescence microscope (Eclipse 80i, Nikon, Japan). Based on the procedures described in our previous work (Albashari et al., 2020; Luo et al., 2021), the multipotent property of DPSCs was analyzed by osteogenic, adipogenic, and chondrogenic differentiation with alizarin red S staining, Oil Red O staining, and Alcian blue staining, respectively. Images were observed using a light microscope (TS100, Nikon).

## 2.3 Effects of DPSCs on murine macrophage RAW264.7 cells

The effects of DPSCs on murine macrophage RAW264.7 cells were evaluated by performing a macrophage-DPSCs coculture experiment as previously described (Németh et al., 2009; Chen et al., 2019). First, the mouse monocyte-macrophage cells (RAW 264.7, ATCC, United States) were cultured in 24-well plates at a concentration of 2 × 10⁵ cells/well. After incubation for 24 h, the supernatants were removed, and the cells were treated with lipopolysaccharide (LPS, Sigma; 1 µg/ml) for 24 h. Then, DPSCs (2 × 10⁴ cells/well), which were cultured in insert membranes of the transwell system (0.4 µm pore size, Corning), were transferred to the wells. After incubation for another 24 h, the supernatants were collected and stored at -20°C for further use. The expression of tumor necrosis factor-alpha (TNF-α) and interleukin-10 (IL-10) of

challenged cells and supernatants was assessed by immunofluorescence staining and Enzyme-linked immunosorbent assay (ELISA) (R&D Systems), respectively.

## 2.4 Evaluation of angiogenic characteristics of DPSCs

To evaluate the angiogenic characteristics of DPSCs, following *in vitro* experiment was performed to assess the formation of capillary-like structures. 96-well plates were pre-coated with gelatin methacryloyl (GelMA) hydrogel as previously described (Chen et al., 2012). In brief, DPSCs and human umbilical vein endothelial cells (HUVECs, ATCC, United States) were plated onto hydrogel-treated 96-well plates and cultured with an endothelial cell growth medium-2™ (EGM-2™, Lonza Bioscience, Switzerland) supplemented with 2% FBS, 0.4% hFGF-B, 0.1% VEGF, 0.1% hEGF, 0.1% R3-IGF-1, 0.1% Heparin, 0.1% ascorbic acid, 0.1% gentamicin/amphotericin-B, and 0.04% hydrocortisone. The cultured medium was changed every 2 days. After differentiation for 7 days, the cells were fixed with 4% PFA and treated with Phalloidin-TRITC (1:200, Solarbio, China), and the nuclei were stained by 4'6-Diamidino-2-phenylindole-dihydrochloride (DAPI, Beyotime Institute of Biotechnology, Shanghai, China). The images were taken using a fluorescence microscope (Eclipse 80i, Nikon, Japan), and the vascular tube length and number were calculated using NIS-Elements AR 3.1 (Nikon).

## 2.5 Preparation of PF127 hydrogel

PF127 powder was slowly added to the complete α-MEM (containing 20% FBS, 100 µg/ml of penicillin, and 100 µg/ml of streptomycin) to prepare a 20% (w/v) solution. The solution was placed on a shaking table at 4°C for 24 h to ensure that the powder was fully dissolved. Then, the PF127 solution was filtered by using a .22 µm pore size bottle-top filter and stored at 4°C for future use. Our previous study confirmed that the PF127 solution could preserve a solution state at 4°C and transform a hydrogel state at body temperature (Albashari et al., 2020).

## 2.6 Cytocompatibility of PF127 hydrogel

The biocompatibility of PF127 hydrogel was evaluated using a Live/Dead Viability/Cytotoxicity Kit (Invitrogen, CA, United States) and Cell Counting Kit-8 (CCK-8, Dojindo Molecular Technologies) assay. In brief, the cultured DPSCs were trypsinized and resuspended in PF127 solution on ice. Then, 200 µl of cell-PF127 mixture solution containing 5 × 10⁴ cells/ml was added to a 48-well plate. After seeding, the
plate was kept at 37°C in 5% CO₂ for 5 min to induce gel
formation. The fresh medium was then added to each well,
and the plate was transferred to a cell incubator. After being
incubated for 1, 3, and 5 days, the Calcein-AM/pyridine iodide
reagent mixture solution was added to the plates and incubated
in the dark at 37°C for 10 min and observed using a fluorescence
microscope (Eclipse 80i, Nikon, Japan). 10% CCK-8 solution was
added to each well and transferred to a cell incubator for another
2 h to quantitatively analyze the cell proliferation activity.
Absorbance at 450 nm was measured using an absorbance
microplate reader (Varioskan LUX, Thermo Fisher Scientific).

## 2.7 Animal model establishment

All experiments were performed under sterile conditions.
Female rabbits were anesthetized by a mixture of 2% isoflurane in 70% nitrous oxide and 30% oxygen and maintained with 1% isoflurane via a face mask. Then the skin was prepared, and a 4-5 cm incision was made to expose the FT and ovaries. The junction between the FT and the uterus was clamped with an arterial clip. 1 ml of absolute ethanol was injected into the FT using a 1 ml syringe. After 2 min, the arterial clip was loosened, and the FT was wiped around using a moist gauze and sutured. After 4 h, the HE results indicated that the FT had evident damage, and the structure of mucosa folds collapsed (Supplementary Figure S1), indicating that the animal model of FT injury was successfully constructed.

## 2.8 Experimental animals and grouping

Twenty four female white rabbits were randomly assigned
to four groups: 1) sham group without any treatment; 2) FT
injury (FTI) group that directly injected saline into the FT
after the effect of anhydrous alcohol for 4 h; 3) DPSCs group
that received an intravenous injection of DPSCs at the ear
margin after the effect of anhydrous alcohol for 4 h; 4) PF127-
encapsulated DPSCs (DPSCs-PF127) group that injected the
mixture of DPSCs and PF127 into the FT after the effect of
anhydrous alcohol for 4 h. After treatment for 4 weeks, the
rabbits were euthanized by intraperitoneal injection of 3%
pentobarbital sodium and the samples were collected and used
for future analysis.

## 2.9 Histopathologic evaluation

The FT specimens were fixed in 4% formaldehyde, embedded
in paraffin, sectioned into slides, stained with hematoxylin and
eosin (H&E) and Masson's trichrome, and recorded using an
optical microscope (ECLIPSE 80i, Nikon, Tokyo, Japan).

## 2.10 TdT-mediated dUTP nick-end
labeling (TUNEL) staining

The TUNEL apoptosis assay kit (Beyotime, Nantong, Jiangsu, China) assay was performed as previously described (Zhu et al., 2020). In brief, the sections were dewaxed in xylene and rehydrated with gradient ethanol. Then, the sections were treated with TUNEL-FITC (1:200), and the nuclei were stained with DAPI for 10 min. Images were taken using a fluorescence microscope (Eclipse 80i, Nikon, Japan).

## 2.11 Immunofluorescence and
immunohistochemistry

Immunofluorescence was performed to assess the primary
antibody of VEGF-A and KI67 according to the manufacturer's
procedures. The sections were treated with primary antibodies,
namely, VEGF-A (1:200, Abcam, Cambridge, United Kingdom)
and KI67 (1:200, Abcam, Cambridge, United Kingdom),
overnight. Then, the sections were incubated with fluorogenic
secondary antibodies for 1 h at 37°C. Finally, the sections were
washed with PBS and incubated with DAPI for 10 min. For
immunohistochemistry, the sections were treated with a primary
antibody, namely, Vimentin (VIM; 1:200, Abcam, Cambridge,
United Kingdom), overnight at 4°C and then incubated with
horseradish-peroxidase-conjugated secondary antibody for
another 2 h at 37°C. Then, the sections were developed using
3,3'-diaminobenzidine. All sections were observed using a
fluorescence microscope (Eclipse 80i, Nikon, Japan).

## 2.12 ELISA

The concentrations of TNF-α and IL-10 in serum were
measured using TNF-α (CSB-E06998Rb, CUSABIO Biotech
Co., Ltd., Wuhan, China) and IL-10 (CSB-E06897Rb,
CUSABIO Biotech Co., Ltd., Wuhan, China) ELISA kits. The
blood of rabbits was collected; the serum was obtained after
centrifugation, and the serum level of inflammatory cytokines
was detected in accordance with the manufacturer's protocols.
All of the samples were repeated.

## 2.13 Gene expression analysis

After 4 weeks, the FT was collected. Total RNA was extracted
from 50 mg of FT using a TRIzol reagent (Invitrogen, CA,
United States). One microgram of RNA was reverse
transcribed into cDNA using a PrimeScript™ RT reagent kit
gDNA Eraser (RR047A, TaKaRa BIO Inc., Kusatsu, Shiga,
Japan). Real-time PCR was performed using the SYBR Premix
EX TaqTM II. cDNA was amplified using the primer sequences
<table><caption>TABLE 1 Forward and reverse primer sequences.</caption><thead><tr><th rowspan="2">Genes</th><th colspan="2">Primer sequence (5'-3')</th></tr><tr><th>Forward</th><th>Reverse</th></tr></thead><tbody><tr><td>18S</td><td>GAATTCACCAGTAAGTGCGGGTCATA</td><td>CGAGGGCCTCACTAAACCATC</td></tr><tr><td>OVGP</td><td>GGATGTCTGAAGCACCCAGAGGT</td><td>AGGTCATCGTCATCTTGCCAGGG</td></tr><tr><td>IL10</td><td>GCCAAGCCTTGTCGGAGATGATC</td><td>CTGCTCCACTGCCTTGCTCTTG</td></tr><tr><td>Ki67</td><td>ATGTTCCATTTCATCAGCCA</td><td>TTTAAATCGCTCCTCCATCC</td></tr><tr><td>Caspase-3</td><td>TGTAAATGCAGCAAACCTCG</td><td>GACTCCTTCATCACCGTGGC</td></tr></tbody></table>

shown in Table 1. The PCR cycling parameters used were as follows: 95°C for 10 s, annealed at 95°C for 5 s and extended at 60°C for 30 s for 40 cycles, followed by thermal denaturation. The mRNA expression was evaluated on the basis of SYBR green fluorescence for 18 s as the endogenous control.

## 2.14 Statistical analysis

Graph pad 7.0 statistics was used to process the data, and the results were presented as means ± standard errors. Statistical differences were obtained by Student's t-test and one-way analysis of variance, followed by Tukey's test or Dunnett post hoc test. *p* < .05 was considered as significant differences.

# 3 Results

## 3.1 DPSCs expressed MSC-like markers and possessed multipotency

The dental pulp is a soft connective tissue containing mesenchymal tissues, which is located in the medullary cavity of the tooth. DPSCs were extracted from the dental pulp, and the procedure for obtaining dental pulp tissues is shown in Supplementary Figure S2A. DPSCs had great cell proliferation capacity, and they can complete the whole primary culture for about 20 days (Supplementary Figure S2B). DPSCs possessed MSC-like properties and positively expressed MSC-like surface markers, including CD166, CD73, CD90, and CD44 (Supplementary Figures S3A, B), but they negatively expressed hematopoietic lineage markers, including CD19, CD34, and HLA-DR (Supplementary Figure S3A). Moreover, DPSCs had an excellent capacity for multilineage differentiation, and they could form mineralized nodules by positive staining of alizarin red S in the induced differentiation of osteoblasts. DPSCs were also induced to differentiate into adipogenesis with the formation of lipid droplets, which were stained with oil red O. In addition,

DPSCs could differentiate into chondrocytes and show the evident staining of Alcian Blue (Supplementary Figure S3C).

## 3.2 DPSCs inhibited the expression of a pro-inflammatory factor (TNF-α) and enhanced the expression of an anti-inflammatory cytokine (IL-10)

The effects of DPSCs on the secretion of inflammatory factors in LPS-induced RAW264.7 cells were evaluated by immunofluorescence staining and ELISA. The results indicated that the expression of TNF-α and IL-10 was upregulated in macrophages by LPS stimulation for 24 h. The expression level of TNF-α was reduced in the presence of DPSCs (Figures 1A, B), whereas the expression of IL-10 was significantly increased (Figures 2A, B). As shown in Figures 1C, 2C, the results indicated that the level of TNF-α and IL-10 in supernatants was also increased after LPS induction for 24 h. In addition, DPSCs could downregulate the expression of TNF-α and upregulate the expression of IL-10, and they showed a significant difference compared with the LPS-induced one.

## 3.3 DPSCs promoted the formation of capillary-like structures

The angiogenic characteristics of DPSCs were evaluated by tubular network formation. The directly induced DPSCs (d-DPSCs), which were 3D cultured with GelMA hydrogel in the completed EGM-2™ medium, could form capillary-like networks. In addition, the number and diameter of closed networks were higher in d-DPSCs than in DPSCs cultured with α-MEM medium, and the difference was statistically significant (Figures 3A, B). Moreover, our results indicated that DPSCs could enhance the capillary-like structure formation of HUVECs using the indirectly cocultured system. Meanwhile, the number and length of capillary-like structures were higher in HUVECs cocultured with DPSCs than in HUVECs cultured alone (Figures 3C, D).
<figure><img src="image_22.png" /><figcaption>**FIGURE 1**
The expression of TNF-α in DPSCs co-cultured with RAW264.7 cells *in vitro*. (A) Immunofluorescence staining of TNF-α. Scale bar: 50 µm. (B) Quantification of the fluorescence intensity of TNF-α. ***p < .001 versus the RAW group. $^{†}$p < .05 versus the LPS/RAW group. (C) Comparison of the supernatants TNF-α concentration among groups. **p < .001 versus the RAW group. $^{‡‡}$p < .001 versus the LPS/RAW group.</figcaption></figure>

<figure><img src="image_23.png" /><figcaption>**FIGURE 2**
The expression of IL-10 in DPSCs co-cultured with RAW264.7 cells *in vitro*. (A) Immunofluorescence staining of IL-10. Scale bar: 50 µm. (B) Quantification of the fluorescence intensity of IL-10. ***p < .001 versus the RAW group. $^{‡‡}$p < .01 versus the LPS/RAW group. (C) Comparison of the supernatants IL-10 concentration among the groups. **p < .01, ***p < .001 versus the RAW group. $^{††}$p < .05 versus the LPS/RAW group.</figcaption></figure>
<figure><img src="image_24.png" /></figure>

## 3.4 PF127 hydrogel had a thermosensitive property and great biocompatibility

As shown in Figure 4A, PF127 hydrogel demonstrated a thermosensitive characteristic. It was clear and liquid solution when it was cold and transited to a homogeneous semi-solid hydrogel status at 37°C. Its gelation time was 80 ± 5 s. The biocompatibility of PF127 hydrogel containing DPSCs was evaluated using Live/Dead and CCK-8 assays. As for the Live/Dead assay, the results indicated that PF127 hydrogel had great cytocompatibility. After being cocultured with DPSCs, most of the cells were alive and stained with green from day 1 to day 5. In addition, the ratio of live cells had no difference between day 3 and day 5, but all of them were higher than that of day 1 (Figures 4B, C). The proliferation of encapsulated DPSCs in PF127 hydrogel was detected for 5 days by CCK-8 assay. The results showed that DPSCs could continue to proliferate in PF127 hydrogel from day 1 to day 5, and the difference was statistically significant (Figure 4D).

## 3.5 DPSCs ameliorated the morphological features of the FT tissue

Four weeks after surgery, upon macroscopic view, injured part of the FT tissue in the FTI and DPSCs groups shrunk greatly

(red arrow). In contrast, injured FT tissue in the DPSCs-PF127 group appeared normal and was similar to that of in the sham group (Supplementary Figure S4). Meanwhile, the cross-section of the middle site of the FT tissue was stained with H&E and Masson's trichrome among the experimental groups. As shown in Figure 5A, the H&E-staining results showed that the normal FT tissue had large mucosa-fold-filled lumen and loosely thick muscularis wall in the sham group. Moreover, the mucosal fold was composed of a single columnar epithelium and a branched core of vascular tissues. In the FTI group, the mucosal fold disappeared, and several infiltrated inflammatory cells and densely thin muscle layers could be observed in the damaged tube tissue. After treatment with DPSCs alone or DPSCs-PF127 in situ injection, the FT gradually returned to its normal structure, and mucosal folds with multi-branching recovered, but inflammatory infiltration also occurred in the lumen and muscle wall in the DPSCs group. In the DPSCs-PF127 group, the structure of mucosal folds and muscle walls was almost similar to that of the sham control group. As for Masson's trichrome staining in each group (Figure 5B), the results showed thin and sparse collagen fibers, which were stained with blue dispersed in the submucosal and muscle wall in the sham group. However, massive collagen fiber infiltration occurred, which was stained with deep blue in the muscle wall and lumen in the FTI group. After treatment with DPSCs alone, a moderate amount of collagen fibers (blue) scattered in the muscle
<figure><img src="image_25.png" /><figcaption>**FIGURE 4**
The biocompatibility of PF127 hydrogel. (A) The thermosensitive property of PF127. (B) The images of Dead/Live assay of DPSCs encapsulated into the PF127 hydrogel after 1, 3 and 5 days, Scale bar: 200 µm. (C) Statistical analysis of the cell viability. The ratio of green to green plus red cells provided the percentage of cell viability. **p** < .01, ***p** < .001 versus the Day 1. (D) The proliferation activity of DPSCs encapsulated into the PF127 hydrogel after 1, 3 and 5 days were measured using Cell Counting Kit-8. **p** < .001 versus the Day 1, $$p$$ < .001 versus the Day 3.</figcaption></figure>

wall, some of which were infiltrated in the lumen. In the DPSCs-PF127 group, the FT preserved its normal structure, and the muscle wall and mucosal fold were filled with a small number of collagen fibers, which had been shown in blue.

## 3.6 Evaluation of TUNEL, immunofluorescence staining, and immunohistochemistry

The effect of DPSCs on FT cell apoptosis was analyzed by TUNEL staining. As shown in Figure 6A, TUNEL staining showed evident green fluorescence in the FTI group, indicating that FT cells underwent significant apoptosis. However, a small amount of TUNEL staining cells were found in the sham group. After treatment of DPSCs, TUNEL staining with green fluorescence gradually

reduced, but the expression of TUNEL in the DPSCs group was stronger than that in the DPSCs-PF127 group. The number of TUNEL staining cells was highest in the FTI group and lowest in the DPSCs-PF127 group (Figure 6B). The number of TUNEL staining cells had no significant difference between the sham group and DPSCs-PF127 group (p > .05).

For immunofluorescence staining and immunohistochemistry analysis, the expression of a cell proliferation marker (KI67) and endothelial cell markers (VEGF-A and VIM) was investigated to evaluate the effect of DPSCs on FT cell proliferation and vascularization *in vivo*. The fluorescence expression of KI67 and VEGF-1 in the DPSCs and DPSCs-PF127 groups was stronger than that in the FTI group. All of the markers were highly expressed in the DPSCs-PF127 group and moderately expressed in the DPSCs and sham groups but weakly expressed in the FTI group (Figures 7A, 8A). In the three experimental groups, the
<figure><img src="image_26.png" /><figcaption>**FIGURE 5**
The H&E (A) and Masson trichrome (B) staining of FT after 4 weeks surgery among the groups. Scale bar: 100 µm.</figcaption></figure>

<figure><img src="image_27.png" /><figcaption>**FIGURE 6**
DPSCs-PF127 reduces cell apoptosis of the FT at 4 weeks after injury. (A) Immunofluorescence staining for TUNEL (green) of sections from the FT in each group. Scale bar: 50 µm. (B) Quantitative estimation of apoptotic and TUNEL cells from three independent sections. All data represent Mean values ±SEM. \*p < .05, \*\*p < .01 versus the Sham group. \*\*\*p < .001 versus the FTI group. \#p < .05 versus the DPSCs group.</figcaption></figure>
<figure><img src="image_28.png" /><figcaption>**FIGURE 7**
DPSCs-PF127 promotes angiogenesis of the FT at 4 weeks after injury. (A) Immunofluorescence staining for VEGF-A (red) of sections from the FT in each group. Scale bar: 50 µm. (B) Quantitative estimation of VEGF-A-positive cells from three independent sections. All data represent Mean values ±SEM. **p** < .01 versus the Sham group. \*\*p < .01 versus the FTI group. \*\*\*p < .01 versus the DPSCs group.</figcaption></figure>

intensity of KI67 and VEGF-A-positive regions was presented in the following order: DPSCs-PF127 > DPSCs > FTI. Meanwhile, a statistically significant difference was found between the DPSCs-PF127 group and the sham group (*p* < .01, Figures 7B, 8B). For immunohistochemistry (Figure 9), the results showed that VIM was highly expressed in the DPSCs-PF127 group and moderately expressed in the sham group but weakly expressed in the DPSCs and the FTI groups (Figure 9A). The intensity of VIM-positive regions was highest in the DPSCs-PF127 group compared with that of the other groups, but no statistically significant difference was found between the DPSCs group and FTI group (*p* > .05, Figure 9B).

## 3.7 Detection of serum TNF-α and IL-10 expression levels and real-time PCR analysis

ELISA was used to evaluate the serum levels of TNF-α and IL-10 in each group. The results showed that the serum TNF-α levels in the FTI and DPSCs groups were significantly higher than those in the sham and DPSCs-PF127 groups (*p* < .001). Meanwhile, no statistically significant difference was observed between the DPSCs-PF127 and sham groups (*p* > .05). However, the serum IL-10 levels in the FTI group were lower than those in the other groups (*p* < .01), and no statistically significant difference was observed among these groups except for the FTI group (*p* > .05,

Figure 10A). For real-time PCR (RT-PCR) analysis (Figure 10B), the results indicated that OVGP, IL-10, and KI67 mRNA were highly expressed in the DPSCs-PF127 group and moderately expressed in the DPSCs and sham groups but weakly expressed in the FTI group. However, no significant difference was observed between the sham and DPSCs groups (*p* > .05). In addition, caspase 3 mRNA was weakly expressed in the DPSCs-PF127 group and highly expressed in the TFI group. Moreover, the expression level of caspase 3 showed a significant difference between the TFI group and other groups (*p* < .05).

## 4 Discussion

The FT is important for egg transportation, sperm activation, and early embryonic development in the reproductive process. Therefore, FT, which preserves the normal structure, secretion property, and peristalsis, plays an important role in reproductive functions of women (Lyons et al., 2006; Duran et al., 2014). In addition, the integrity of mucosal fold and the great secretion of epithelial cells in the FT directly affect the surface area of cell reception and cell activation (Hunter 2005). In our previous study, the H&E results confirmed that the structure of FT, particularly the mucosal fold and muscle wall, was destroyed by anhydrous ethanol for 4 h (Supplementary Figure S1C). After 4 weeks, the mucosal fold disappeared, and inflammatory macrophage
<figure><img src="image_29.png" /><figcaption>**FIGURE 8**
DPSCs-PF127 enhances cell proliferation of the FT at 4 weeks after injury. (A) Immunofluorescence staining for KI67 (green) of sections from the FT in each group. Scale bar: 50 µm. (B) Quantitative estimation of KI67-positive cells from three independent sections. All data represent Mean values ±SEM. \*p < .05 versus the Sham group. \textasciitilde;p < .05 versus the FTI group. \textasciitilde;p < .01 versus the DPSCs group.</figcaption></figure>

<figure><img src="image_30.png" /><figcaption>**FIGURE 9**
The immunohistochemistry assay of the FT at 4 weeks after injury. (A) Representative images of VIM from immunohistochemistry. Scale bar: 100 µm. (B) Quantification of the VIM positive staining ratio to normal in FT. The quantification results obtained by ImageJ. All data represent Mean values ±SEM. \textasciitilde;p < .001 versus the Sham group. \textasciitilde;p < .001 versus the FTI group. \textasciitilde;p < .001 versus the DPSCs group.</figcaption></figure>

infiltration, muscle wall fibrosis, and lumen stenosis were observed in the FTI group, which had not received any treatment (Figure 5). Moreover, such structural and functional destructions in the FT would increase the tubal

infertility rate. Therefore, the repair and regeneration of the damaged FT played important roles in reducing tubal infertility.

MSCs have shown a strong immunomodulatory property, which could be used as an immunosuppressive agent to modulate
<figure><img src="image_31.png" /></figure>

<figure><img src="image_32.png" /></figure>

<figure><img src="image_33.png" /></figure>

<figure><img src="image_34.png" /></figure>

<figure><img src="image_35.png" /></figure>

<figure><img src="image_36.png" /><figcaption>**FIGURE 10**

Concentration of inflammatory factors in serum and mRNA detection of the FT at 4 weeks after surgery. (A) Comparison of the serum TNF-α and IL-10 concentration among the groups. \*\*p < .01, \*\*\*p < .001 versus the Sham group. \*\*\*\*p < .001 versus the FTI group. \*\*\*\*p < .01 versus the DPSCs group. (B) The expression of OVGP, IL-10, Caspase 3, Ki67 mRNA of the FT tissues among the groups. \*p < .05, \*\*p < .01 versus the Sham group. \*\*p < .05, \*\*\*p < .01 versus the FTI group. \*\*\*\*p < .01, \*\*\*\*\*p < .001 versus the DPSCs group.</figcaption></figure>

the immune response in inflammatory or autoimmune diseases (Liu et al., 2020; Wu et al., 2020; Zhou et al., 2020). In addition, MSCs had a great angiogenic property, and they could promote vessel-like formation by secreting various growth factors or directly differentiating into endothelial cells (Takeuchi et al., 2019; Merckx et al., 2020; Santos et al., 2020). Thus, MSC-based cell therapy, such as Wharton's jelly-derived mesenchymal stem cells, BM-MSCs, and human umbilical cord mesenchymal stem cells, served as a suitable strategy for the treatment of FT injury. In these studies, most MSCs were used by intraperitoneal or intravenous injection, which had systemic effects on the repair and regeneration of FT injury (Luo et al., 2015; Li et al., 2017; Almasry et al., 2018; Liao et al., 2019). However, no study had reported that the therapeutic method of MSCs in the treatment of FT damage was currently used by *in situ* injection.

In this study, we used DPSCs, which were obtained from dental pulp without harm or invasive surgery, as seeded cells for the treatment of FT injury by PF127 hydrogel *in situ* transplantation. As previously described, PF-127 hydrogel had a thermosensitive property and great cytocompatibility with DPSCs, which was an ideal scaffold for cell survival and proliferation (Albashari et al., 2020). In *in vitro* cell culture

experiments, our results indicated that DPSCs, which were dispersed in PF-127 hydrogel, could proliferate from day 1 to day 5 (Figures 4B-D). Moreover, DPSCs showed great immunomodulatory and angiogenic characteristics, and they could promote the expression and secretion of anti-inflammatory cytokines such as IL-10, but they could inhibit the expression and secretion of pro-inflammatory factors such as TNF-α (Figures 1, 2) and enhance the formation of vessel-like structures (Figure 3). Therefore, PF127 hydrogel containing DPSCs could provide an effective treatment for FT injury by *in situ* injection.

For histological analysis, the H&E and Masson's trichrome staining results indicated that FT structural damage was improved after treatment with DPSCs, but the repair effects in the DPSCs-PF127 group were better than those in the DPSCs group (Figure 5). Previous studies revealed that 3D-network hydrogel could be used as a cell carrier, which could not only ensure the number of active cells in the injured site but also provide a suitable microenvironment for cell growth and proliferation (Long et al., 2020; Villiou et al., 2020). In this study, the results indicated that PF127 hydrogel encapsulating DPSCs by *in*
situ injection could be used to promote the repair and regeneration of FT injury.

In our study, TUNEL assay was performed to detect FT apoptotic cells in the injured site because of massive DNA fragmentation in the nuclei of apoptotic cells, which could be labeled with green FITC fluorescence (Zhu et al., 2020; He et al., 2021). Caspase-3, as an important marker of programmed cell death, was responsible for the cleavage of many proteins during apoptosis (Elmore 2007). Our results showed that the positive region of TUNEL staining in the DPSCs-PF127 group was lowest among the groups; meanwhile, caspase-3 mRNA expression was significantly lower in the DPSCs-PF127 group than in the FTI group. These results supported in situ transplantation of PF127 hydrogel containing DPSCs into the injured FT, which likely promoted the repair and regeneration of the injured FT by inhibiting cell apoptosis.

KI67, which is a representative proliferation marker, was widely used to evaluate cell proliferation (Willenbacher et al., 2020; Murata et al., 2021). VEGF-A and VIM, which served as vascular endothelial-associated markers contributing to angiogenesis, were commonly used to assess the angiogenic property (Almasry et al., 2018; Chen et al., 2021). In our research, the results revealed that the expression level of KI67, VEGF-A, and VIM was higher in the DPSCs-PF127 group than in the sham control group. Meanwhile, the difference in KI67 and VEGF-A showed no statistically significant difference between the DPSCs and sham control groups, whereas VIM expression showed a contrary result. These results indicated that PF127 hydrogel containing DPSCs could enhance the damage repair and the tissue regeneration of FT, by stimulating the secretion of angiogenic factors and promoting cell proliferation at the injury site.

Tumor necrosis factor-alpha (TNF-α) and interleukin-10 (IL-10) are two representative cytokines produced by the immune cells. Both play key roles in maintaining immune homeostasis and modulate the inflammatory reaction (Wang et al., 2022). And IL-10 can inhibit the production of TNF-α (Kuwata et al., 2003; Passoja et al., 2010). Thus, in this study, TNF-α and IL-10 were used to evaluate the immunomodulatory function of DPSCs in the repair of injured FT. TNF-α, which is a pro-inflammatory factor produced by mononuclear macrophages, was closely associated with the reproductive immunization of the female genital system (Morales et al., 2006; Oróstica et al., 2013). IL-10, which served as a cytokine exerting anti-inflammatory effects, was produced by activated macrophages. A previous study indicated that DPSCs could upregulate the secretion of IL-10 by coculturing with T cells (Demircan et al., 2011). Our results showed that the expression of IL-10 in serum had no significant difference between the DPSCs-PF127 and sham control groups. Moreover, the mRNA expression of IL-10 in the DPSCs-PF127 group was higher than that in the sham control group. However, the level of TNF-α showed a lower expression in the DPSCs-PF127 group than in other groups. These results indicated that PF127 hydrogel containing DPSCs promoted the repair and regeneration of FT damage, thereby stimulating the secretion of IL-10 and inhibiting the secretion of TNF-α.

Oviductal glycoprotein (OVGP), which is a multi-protein complex produced by the secretion of FT epithelium, could be used to evaluate the function of epithelial cells (Luo et al., 2015; Li et al., 2017). Our results indicated that OVGP was highly expressed in the DPSCs and DPSCs-PF127 groups, which had no significant difference compared with the sham control group. In addition, OVGP was weakly expressed in the FTI group, indicating a serious damage in the mucosal fold. These results indicated that DPSCs could promote the repair and regeneration of mucosal fold and restore its secretion functions.

## 5 Conclusion

In the present study, the effects of PF127 hydrogel containing DPSCs on the repair and regeneration of FT injury were evaluated by histological analysis (e.g., H&E and Masson's trichrome staining, TUNEL staining, immunofluorescence staining, and immunohistochemistry), ELISA, and RT-PCR detections. Our results indicated that PF127 hydrogel containing DPSCs promoted the repair and regeneration of FT damage, thereby inhibiting cell apoptosis, stimulating the secretion of angiogenic factors, promoting cell proliferation, modulating the secretion of inflammatory factors, and restoring the secretion of epithelial cells. Therefore, this study was the first to report that *in situ* injection of PF127 hydrogel encapsulating DPSCs could be used to treat FT injury in clinical settings.

## Data availability statement

The original contributions presented in the study are included in the article/Supplementary Material, further inquiries can be directed to the corresponding authors.

## Ethics statement

The animal study was reviewed and approved by the Animal Experimental Center of Wenzhou Medical University. Written informed consent was obtained from the owners for the participation of their animals in this study.

## Author contributions

QY, LL, and YL performed the majority of experiments and analyzed the data; QZ collected the data and performed statistical analysis; JY cultured DPSCs; ZX, XL, and FH performed multi-differentiation of DPSCs; LL, QY, and YH designed and coordinated the research; LL and QZ wrote the paper.
## Funding

This research was partially supported by the National Natural Science Funding of China under Grant No. 81871503, the Wenzhou Major Scientific and Technological Innovation Key Medical and Health Project under Grant No. ZY2019010, and Zhejiang Provincial Natural Science Foundation of China under Grant No. LGF21H140007, and Zhejiang Province Program of the Medical and Health Science and Technology under Grant No. 2022KY218.

## Conflict of interest

The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

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## Supplementary material

The Supplementary Material for this article can be found online at: https://www.frontiersin.org/articles/10.3389/fbioe.2022.1062646/full#supplementary-material

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<figure><img src="image_37.png" /></figure>

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EDITED BY

Hengchong Shi,
Changchun Institute of Applied Chemistry
(CAS), China

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Kai Wang,
Nankai University, China
Wei Nie,
Wake Forest Baptist Medical Center,
United States

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Ruilan Wang,
 wangyusun@hotmail.com

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This article was submitted to Biomaterials,
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RECEIVED 11 October 2022

ACCEPTED 19 December 2022

PUBLISHED 10 January 2023

CITATION

Zhang W, Du J, Zhu T and Wang R (2023), SiO₂ nanosphere coated tough catheter with superhydrophobic surface for improving the antibacteria and hemocompatibility.
*Front. Bioeng. Biotechnol.* 10:1067139.
doi: 10.3389/fbioe.2022.1067139

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SiO₂ nanosphere coated tough
catheter with superhydrophobic
surface for improving the
antibacteria and
hemocompatibility

Weixing Zhang¹, Juan Du², Tonghe Zhu² and Ruilan Wang¹*

¹Department of Critical Care Medicine, School of Medicine, Shanghai General Hospital, Shanghai Jiao Tong University, Shanghai, China, ²School of Chemistry and Chemical Engineering, Shanghai Engineering Research Center of Pharmaceutical Intelligent Equipment, Shanghai Frontiers Science Research Center for Druggability of Cardiovascular Non-coding RNA, Institute for Frontier Medical Technology, Shanghai University of Engineering Science, Shanghai, China

Catheter infection is the most common complication after vascular catheter placement, which seriously threatens the survival of critically ill patients. Although catheters with antibacterial drug coatings have been used, catheter infections have not been effectively resolved. In this research, a SiO₂ nanosphere-coated PTFE catheter (PTFE-SiO₂) with enhanced antibacterial and excellent mechanical properties was prepared *via* dopamine as a graft bridge. The microscopic morphology results show that the nanospheres are uniformly dispersed on the surface of the catheter. The physicochemical characterization confirmed that PTFE-SiO₂ had reliable bending resistance properties, superhydrophobicity, and cytocompatibility and could inhibit thrombosis. Antibacterial results revealed that PTFE-SiO₂ could hinder the reproduction of *E. coli* and *S. aureus*. This research demonstrates the hydroxyl-rich materials obtained by hydroboration oxidation have the advantages of better dispersion of functional coatings, indicating their potential for helpful modification of catheters.

KEYWORDS

superhydrophobic, SiO₂ nanosphere, catheter, antibacterial, blood compatibility

1 Introduction

Long-term intravenous fluid rehydration, central venous pressure and arterial monitoring, hemodialysis, and ECMO are important treatment methods in the critical care department through the placement of central venous catheters, hemodialysis catheters, and radial artery catheters (Gilbert et al., 2013; Shimizu et al., 2020; Schalk et al., 2022). However, catheter-related infection is a common and severe complication after implantation, which often requires removal of the catheter, performing catheter end and blood culture to determine the source of infection and bacterial species (Ferreira et al., 2020; Brescia et al., 2021). Finally, antibiotics are administered, and the catheter is reinserted at another site. The infectious bacteria associated with the catheter is MASA. Due to the poor antibacterial effect of most antibiotics, the emergence of super bacteria has caused great harm to human health. This causes pain for the patient, increases the cost of treatment, and increases the risk of other complications (Akhrass et al., 2011; Jana et al., 2018).

Previous studies have shown that bacteria's rapid, nonspecific adhesion to the hydrophilic and lipophilic groups on the catheter surface is the initial cause of catheter infection (Liu et al.,
2017; Alvarenga et al., 2020; Li et al., 2020). The effect of bacterial adhesion depends on the microscopic topography and chemical properties of the catheter surface, as well as the characteristics of the bacteria (Monzillo et al., 2012; Liu et al., 2017; Yu et al., 2017; Faustino et al., 2020; Yu et al., 2021). How to reduce the colonization of bacteria in the catheter and avoid the formation of biofilm, thereby reducing the occurrence of catheter infection and the difficulty of treatment, has increasingly become a fundamental problem in clinical medicine, and more and more studies are constantly proposing new methods to solve this problem (Goncalves et al., 2014; Radu Mihu et al., 2017; Zhang et al., 2019; Gunaratnam et al., 2020). The current research on catheter-related infection globally mainly focuses on the antibacterial treatment after catheter infection. Once infected, high-grade antibiotics are required for treatment. In terms of prevention, the primary method is to coat the catheter with antibiotics. There are currently dialysis catheters that use coatings such as bismuth, which reduce bacterial colonization and thus film formation (Schindler et al., 2010; Balikci et al., 2021; Pant et al., 2022).

In addition to coating antibacterial drugs on catheters, finding the best material for antibacterial is also one of the directions we need to work on. Inspired by the origin of bacterial adhesion, we expect to develop a new type of coated vascular catheter that can effectively isolate bacterial adhesion factors. Therefore, we plan to design a novel anti-infective superhydrophobic nanocomposite-coated vascular catheter. The catheter coating and base material will not use antibacterial drugs, and the anti-infection purpose is mainly achieved by the superhydrophobicity of the surfaces of material.

# 2 Materials and methods

## 2.1 Materials

Commercial venous catheters were purchased from Arrow International, Inc. (Cleveland, United States). Both PTFE membranes (FP301350, with 0.25 mm thickness) and PTFE catheters (with 0.6 mm wall thickness and 1.0 mm inner diameter) were purchased from Goodfellow Tading Co., Ltd. (Shanghai, China). SiO₂ nanosphere powder (average particle size is 20 nm) were purchased from Suzhou Vmicronano New Materials Co., Ltd. (Suzhou, China). Benzoin (purity ≥ 99.5%), potassium tert-butoxide (purity ≥ 99.9%), and borane tetrahydrofuran (purity ≥ 99.99%) were purchased from Adamas Reagent Co., Ltd. (Shanghai, China). Dimethyl sulfoxide (DMSO), tetrahydrofuran (THF), anhydrous ethanol (EtOH, purity ≥ 99.99%) were obtained from Sinopharm Chemical Reagent Co., Ltd. (Shanghai, China). Other reagents and materials are not treated before use unless otherwise specified.

## 2.2 Surface modification of PTFE membrane and catheter

### 2.2.1 Preparation of PTFE-OH membrane and catheter

The circular PEFE membranes with a diameter of 14 mm and PTFE catheters with 0.6 mm thickness and 1.0 mm inner diameter (uniformly named PTFE) were ultrasonically cleaned in 75% medical

alcohol for 30 min, and then freeze-dried in vacuum for subsequent experiments, respectively.

Firstly, 1.5 g of benzoin and 5.1 g of potassium tert-butoxide were dissolved in 10 and 50 ml of dimethyl sulfoxide, respectively. Then, the above two harvested solutions were mixed until present clear and homogeneous (Costello and McCarthy, 1984). Secondly, The cleaned PTFE was added to the mixed solution, and the reaction was continued for 36 h at 54°C in a constant temperature water bath. After the reaction was over, the modified PTFE were washed twice with tetrahydrofuran reagent, following washed with deionized water for 5 times. Finally, the modified PTFE were freeze-dried for 48 h, namely PTFE-M.

Finally, hydroxyl groups were introduced on the surface of PTFE according to the hydroboration oxidation reaction (Meng et al., 2015; Dan et al., 2020): PTFE-M was added to 10 ml borane tetrahydrofuran solution in three-necked flask. Then the three-necked flask was sealed, and the reaction was carried out at room temperature for 12 h. The borane tetrahydrofuran solution was then transferred out, and then 10 ml of deionized water, 10 ml of 3 mol/L NaOH solution and 10 ml of 30% H₂O₂ were added in sequence. Finally, place the three-necked flask into the prepared ice-water mixture and ice-bath for 3 h under a magnetic stirrer. After the reaction is completed, the modified PTFE is washed with dilute hydrochloric acid solution, alkene sodium hydroxide solution and deionized water in sequence. Finally, the hydroxylated PTFE were vacuumed in a desiccator for 48 h to remove redistill solvent and namely PTFE-OH.

### 2.2.2 Preparation of PTFE-SiO₂ membrane and catheter

First, 200 mg of dopamine and 2 g of nanospheres were added to a Tris (pH = 8.5) solution, which were prepared using 200 ml deionized water. Then PTFE-OH was added to the above solution which was mixed well and soaked for 4 h. Finally, the modified PTFE were washed thoroughly with deionized water, following freeze-dried in a vacuum freeze dryer for 48 h to remove redistill solvent and namely PTFE-SiO₂.

## 2.3 Characterization and tests of prepared membrane or catheter

The morphology and surface structure of modified PTFE were carried out using a scanning electron microscope (SEM, Phenom XL, Netherlands) operating with sputter gold plating for 35 s at 5 mA at an accelerating voltage of 10 kV. FTIR spectra of modified PTFE were analyzed on a Nicolet 6700 spectrometer (Thermo Fisher, United States) in the spectrum range of 500 and 4000 cm⁻¹ with a resolution of 1 cm⁻¹, setting 30 scans for a single analysis. A contact angle measuring device (JC 2000D 2A, Shanghai Zhongchen Digital Technology Equipment Co., Ltd., China) was used to testing the wettability of modified PTFE. In this testing, 0.02 ml deionized water was added to the sample, and three groups of parallel samples were taken from each group of PTFE to test the water contact angle and calculate the average value.

High-precision mechanical property testing machine (HY-025CS, Shanghai Hengyu Instrument Co., Ltd., China) with a transducer with a load range of 0–1000 N was employed to testing the compressive performance of prepared catheter in dry conditions at room temperature. Catheter with a size of 2 mm inner diameter, 50 mm
<figure><img src="image_38.png" /><figcaption>FIGURE 1</figcaption></figure>

(A) The digital photos of commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter, respectively; (B) FTIR spectra of commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter, respectively; (C) and (D) Water contact angle of the lumen surface of the inner surfaces of PTFE catheter, and PTFE-SiO₂ catheter at the 5 s time point, respectively.

length, and approximately 0.5 mm in wall thickness were used for radial compression testing. The cyclic compression tests at 50% of the outer diameter of catheters were repeatedly performed 1000 times at a fixed rate of 0.2 mm/min. Each test was repeated five times during mechanical analysis.

## 2.4 Cell culture and cytocompatibility evaluation

Mice fibroblast cells (L929) were cultured with growth medium consisting of dulbecco's modified eagle medium (DMEM), 10% fetal bovine serum and 1% penicillin/streptomycin. Before cell seeding, the commercial membranes, PTFE, and PTFE-SiO₂ membranes were cut into discs with a diameter of 14 mm, then placed into 24-well plates one by one, following covered with sterilized stainless rings, respectively. Each sample was sterilized in 75% ethanol for 12 h and then washed with PBS three times, finally irradiated with ultraviolet for 2 h and socked in growth medium for incubation overnight. HUVECs were seeded at a density of 1.5 × 10⁴ cells/well, and the culture medium was maintained to be replaced every 2 days.

The cell viability of L929 was tested by using the Cell Counting Kit-8 (CCK-8). The cells were cultured in commercial membranes, PTFE, and PTFE-SiO₂ membranes for 1 day, 3 days, and 5 days, respectively. When the set culture time point, the well plate was taken from the CO₂ incubator, and then a series of operations as removal medium, wash samples in well with PBS, adding 360 µL DMEM medium and 40 µL CCK-8 solution into the lucifugal well plate in sequence. After 1 h incubation in CO₂ incubator, The optical

density (OD) value was measured with a microplate reader at a wavelength of 450 nm.

Before the cells SEM images test, the cultured L929 were fixed with 4% paraformaldehyde and dehydrated by gradient ethanol (30%, 50%, 70%, 80%, 90%, 95%, and 100%) on the 3rd day and the morphology carried out using a scanning electron microscope (SEM, Phenom XL, Netherlands) operating with sputter gold plating for 35 s at 5 mA at an accelerating voltage of 10 kV, respectively.

Additionally, the live cells (green) and dead cells (red) were stained with Calcein-AM/PI after 3 days' culture, respectively. The dyed samples were observed immediately under the TS100 fluorescence microscope (Nikon, Japan).

## 2.5 Blood compatibility test in vitro

Fresh blood, which was drawn from the marginal vein, and 3.2% sodium citrate solution in a volume ratio of 9:1 (v/v) were collected using a plastic vacuum blood collection tube (2.7 ml, Becton Dickinson, United States), containing 3.2% sodium citrate solution. All animal experimental protocols are in accordance with the policy of the Institutional Review Board for Human Investigations at School of Medicine of Shanghai Jiao Tong University. Detailed procedures are available in the Supplementary Material.

## 2.6 Testing of antibacterial activity in vitro

*Staphylococcus aureus* (*S. aureus*, BCRC 10451) and *Escherichia coli* (*E. coli*, BCRC 11634) were used to assess the antibacterial ability
<figure><img src="image_39.png" /><figcaption>FIGURE 2
The surfaces SEM images of commercial catheter, PTFE catheter, and PTFE-SiO₂ catheter, respectively.</figcaption></figure>

of commercial membranes, PTFE, and PTFE-SiO₂ membranes. Antibacterial activities were investigated by agar disc diffusion assay as described previously.

Firstly, commercial membranes, PTFE, and PTFE-SiO₂ membranes were pruned into rounded membranes with a diameter of 14 mm. Then, circular membranes were irradiated under an ultraviolet lamp for 12 h, following fumigated with ethanol overnight for thorough sterilization and stored in the bechtop for further testing or evaluation. For non-quantitative synthesis, the bacterial inhibition ring of the *E. coli* and *S. aureus* cultured on the LB agar plates was assessed. In short, *E. coli* and *S. aureus* infection fluid (200 µL) was the first diffusion into the agar plates and the experiment mats (commercial, PTFE, and PTFE-SiO₂ circular membranes) with a diameter of 14 mm and a thickness of about 300 µm were attached to the agar plate. After incubation at 37°C for 24 h, the bacterial inhibition ring of each specimen on the plate was visually measured. The inner and outer diameters and calculated diameter distinction were analyzed for evaluating the antimicrobial activity of prepared circular nanofibrous mats.

## 2.7 Biofilm experiment of catheter in vitro

First, all the catheter (commercial catheter, PTFE catheter, and PTFE-SiO₂ catheter) were cut into multiple segments with 5 mm length, following ultrasonic washed with distilled water and ethanol for 5 min, respectively. Then, all the samples were immersed into 200 µL of selected medium (BHI for VRE and TSB supplemented with 1% glucose for *S. aureus*) in a 96-well plate after vacuum freeze drying

for 48 h. 1.0 × 10⁶ CFU/ml bacteria suspension was made by diluting in 10 mM PBS with pH = 7.4 from overnight culture. 8 µL of the above prepared diluted bacteria suspension was added to each well, following the plate was incubated at 37°C for 24 h (Zhou et al., 2017). All the above samples were then removed and washed gently thrice with PBS to remove unattached bacteria. Then, all the samples were fixed with 4% paraformaldehyde and dehydrated by gradient ethanol (10%, 30%, 50%, 60%, 75%, 80%, 90%, and 100%) and the morphology carried out using a scanning electron microscope (SEM, Phenom XL, Netherlands) operating with sputter gold plating for 35 s at 5 mA at an accelerating voltage of 10 kV, respectively.

## 2.8 Statistical analysis

Origin 9.0 statistical software (Origin Lab Inc., United States), one-way ANOVA and Tukey's test were applied to statistical analysis of the data. All data are expressed as mean ± standard deviation (Mean ± SD). *p*-values < 0.05 (*) were deemed to be statistically significant. * indicates *p* < 0.05, ** indicates *p* < 0.01, *** indicates *p* < 0.001.

# 3 Results and discussion

## 3.1 Microstructure and surface phychemical properties of catheter

As a conduit material, PTFE has advantages in mechanical properties. However, PTFE is an inert material, which limits its
<figure><img src="image_40.png" /><figcaption>**FIGURE 3**
(A) Maximum compressive force and (B) compressive modulus after 1000 cycle compressive fatigue tests; (C) Images of three-point bending tests; (D) Typical the first and one thousandth loading-unloading compression cycle curves of commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter.</figcaption></figure>

application in medical catheters. In order to change the chemical inertness of the surface of PTFE catheter and improve the biocompatibility of PTFE, a very active unsaturated carbon chain layer was formed on the surface of PTFE by treating the mixed solution of benzoin and potassium tert butoxide. Then, a large number of hydroxyl groups were reactive grafted on the surface of PTFE membrane by borohydride oxidation reaction, following grafted SiO₂ nanosphere on the surface of PTFE matrix through dopamine bridging, which presented strong adhesion between the hydroxyl groups of polydopamine and the SiO₂ nanosphere surface. Macroscopically, the catheter could maintain its tubular shape after being treated with organic solution (Figure 1A). We obtained catheters with a length of 25 mm and an inner diameter of 1.5 mm (Figure 1A). The presence of as coated SiO₂ nanosphere was verified via FTIR spectra, by which we demonstrated that PTFE-SiO₂ catheter received more characteristic peak of stretching vibration of Si-O bond at 958 cm⁻¹ after coating (Figure 1B).

With the increasing attention of superhydrophobic surfaces, the reduction of bacterial adhesion by superhydrophobic surfaces has also been considered as an effective antibacterial method (Ozkan et al., 2021; Li et al., 2022). The superhydrophobic surface interface of PTFE catheter play an

important role in the adsorption and the adhesion of bacteria. As shown in Figures 1C, D, the water contact angle values of the lumen surface of the inner surfaces of PTFE catheter and PTFE-SiO₂ catheter were 110.9° ± 1.8° and 151.2° ± 1.5° at 5 s, respectively. The SiO₂ nanosphere-coated PTFE catheter exhibits stronger hydrophobicity. The antibacterial mechanism of superhydrophobic surfaces is that bacteria are more likely to adhere to surfaces with high surface energy, while superhydrophobic surfaces have lower surface energy, which is not conducive to the adhesion of bacteria (Boinovich et al., 2018; Huang et al., 2018; Zhan et al., 2021). Moreover, the air layer trapped by the micro-nanostructure will isolate the bacteria from the SiO₂ coating. When the superhydrophobic catheter is in contact with bacteria, the contact area between the coating and the colony will reducing due to the effect of the air layer, which can avoid the formation of biofilm on the surface of the catheter, and finally achieve the effect of bacteriostasis.

The dispersion of the SiO₂ coating is also one of the main factors affecting the antibacterial effect of the catheter surface. The scanning electron microscope images in Figure 2 showed that the coating consists of many micro- and nano-scale protrusions and pits. These micro-nano structures are mainly formed by the accumulation of nano-SiO₂ particles. The nano-SiO₂ is tightly fixed
<figure><img src="image_41.png" /><figcaption>FIGURE 4</figcaption></figure>

CCK-8 assay of (A) the adhesion viability after 24 h culture and (B) proliferation viability after 1, 3, and 5 days culture of HUVECs onto cover slips, commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter, respectively; (C) Quantitative number measurements of L929 cells on the inner surfaces of cover slips, commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter after culturing for 3 days; (D) Live/dead staining viability assay of L929 cells cultured on the inner surfaces of cover slips, commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter after culturing for 3 days.

on the surface of the PTFE substrate by the dopamine matrix, which improves the mechanical stability of the surface of the superhydrophobic coating. Comparing the commerical group and the PTFE group, it is not difficult to find that the immobilized SiO₂ using dopamine after hydroxylation treatment not only shows a higher graft density, but also disperses more uniformly on the PTFE surface.

## 3.2 Mechanical properties

In order to clarify whether the original mechanical properties of PTFE catheters were changed after the grafting of SiO₂ nanospheres, mechanical properties experiments were carried out. The results of radial cyclic compression at 50% of the outer diameter of catheters after 1000 cycle compressive fatigue tests are shown in Figure 3. As shown in Figures 3A,B, the compressive and compressive modulus after 1000 cycle compressive fatigue tests of PTFE-SiO₂ catheter were significantly lower than those of PTFE group, but consistently superior to the commerical group. According to the standard guidelines of three-point bending test of ASTM F2606-09 as present in Figures 3C, D, all the samples were bendable and returned to their original shape without any visible damage and permanent deformation. Therefore, the modified catheter still maintains reliable bending resistance, which will be sufficient to support the catheter against any possible external force during surgical operation.

## 3.3 Cytocompatibility and hemocompatibility *in vitro*

To test the biocompatibility of the samples before and after modification, the same amount of L929 was inoculated on each sample. As shown in Figures 4A,B, the adhesion ability of the PTFE-SiO₂ group to L929 was higher than that of the commercial group and the PTFE group but still lower than that of the cover slips group. Figures 4C,D show that by day 3, the number of viable cells in the cover slips group and the commerical group was less than that in the other two groups. The PTFE-SiO₂ group had the smallest viable cells, mainly due to the hydrophobic surface and chemical inertness caused by the fluorocarbon structure. The SiO₂ nanocluster structure is unsuitable for cell adhesion and proliferation.

After vascular puncture, the catheter stays in the body for a long time, which can easily lead to nonspecific protein adsorption, microbial infection, and blood coagulation (Hitz et al., 2012; Erathodiyil et al., 2020). These phenomena eventually lead to deep vein thrombosis (PICC-DVT), pulmonary embolism (PE) due to thrombus shedding, and catheter-related bloodstream infections (CRBSIs) due to microbial infections. Currently, drugs such as heparin and antibiotics are injected into patients to reduce catheter-related complications. However, neither heparin nor antibiotics maintain long-term efficacy in the body. Repeated injections of chemicals will not only reduce the patient's immunity and cause suffering from drug side effects. Therefore, optimizing and modifying existing materials is an effective measure to reduce
<figure><img src="image_42.png" /></figure>

<figure><img src="image_43.png" /></figure>

<figure><img src="image_44.png" /></figure>

<figure><img src="image_45.png" /><figcaption>**FIGURE 5**
(A) Whole blood clotting time (B) Plasma recalcification time; (C) Quantification of relative hemolysis rate; (D) Platelet deposition determined by lactate dehydrogenase assay. (For plasma recalcification time test, TCPs exposed to PPP with and without CaCl₂ were used as positive control and negative control, respectively; For hemolysis test, water and 0.9% normal saline (NS) serve as positive and negative groups, respectively; *n* = 5, \*p* < 0.05, \*\*p* < 0.01).</figcaption></figure>

catheter-related complications and improve the anticoagulant and antibacterial properties of catheters.

The anticoagulant property of all samples could be evaluated directly by whole blood clotting time, plasma recalcification time, quantification of relative hemolysis rate, and quantification of lactate dehydrogenase activity. We assessed the whole blood clotting of catheter by the whole blood clotting time test. After the whole blood incubation, the absorbance of the supernatant was measured at 540 nm, as shown in Figure 5A. Compared with the other two groups, the PTFE catheter coated with SiO₂ nanospheres had the maximum absorbance at any time point. The higher the absorbance, the higher the hemoglobin content in the supernatant, which means that there are more blood cells in the free state, less blood cells that produce coagulation, and the slower the coagulation speed. All in all, under the same experimental conditions, the clotting time of the control group was significantly shorter, indicating that the coated-modified PTFE catheter has highly hydrophobic properties and can resist the adhesion of proteins to the material surface. Therefore, PTFE-SiO₂ catheter can reduce the adhesion of blood cells and has outstanding anticoagulant properties.

As we all know, the essence of blood coagulation is the process of fibrinogen in plasma from soluble to insoluble, and blood coagulation is divided into two types: intrinsic coagulation and extrinsic coagulation. All coagulation factors involved in intrinsic coagulation are provided by plasma. When the plasma comes into contact with the material, the electrical charge on the surface activates

the coagulation factors in the plasma. The plasma recalcification curve is a method used to characterize the endogenous coagulation system. The recalcification time is the time it takes for the plasma to coagulate after removing the calcium source and then adding Ca²⁺. The kinetics of plasma recalcification can be seen in Figure 5B for the three water conduits, the positive control group and the negative control group. The positive control has the fastest clot formation time at 15 min, while the clot formation time of the commercial and PTFE group were 20 and 24 min, respectively. Notably, the PTFE-SiO₂ group consistently kept a low absorbance and closed to the negative control without emerging the inflection point. The above results indicate that the PTFE-SiO₂ group has a superhydrophobic surface interface effect, so the activation of endogenous coagulation is less, and the anticoagulation performance of the catheter is improved.

The hemolysis rate of a material refers to the degree to which hemoglobin is lysed by erythrocytes after the material is in contact with erythrocytes. According to international standards, the hemolysis rate of biological materials should be less than 5% (Zhu et al., 2021). Figure 5C is the hemolysis rate of PTFE catheters before and after coating. As shown in the figure, the hemolysis rate of the uncoated PTFE catheter was 1.756%, while after the SiO₂ nanosphere coating, the anti-hemolysis ability was significantly reduced. In conclusion, PTFE-SiO₂ had the best blood compatibility among all samples, and the hemolysis rate was only 0.521%, which was almost close to that of the negative control group.
<figure><img src="image_46.png" /><figcaption>FIGURE 6</figcaption></figure>

(A) Visual images of the antibacterial zone of catheter to *Escherichia coli* (*E. coli*) and *Staphylococcus aureus* (*S. aureus*) after incubation for 6 h. Spot ①, ②, and ③ represents commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter, respectively; of commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter, respectively; (B) and (C) Quantified antibacterial zone of commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter after 6 h incubation with *E. coli* and *S. aureus*, respectively. (n = 5, **p** < 0.01)

The lactate dehydrogenase (LDH) activity of platelet adhesion on the inner surface of catheter after 2 h incubation with rabbit platelet-rich plasma are shown in Figure 5D. The deposition of adhered platelets on the inner surface of PTFE-SiO₂ catheter was 85 platelets/mm², which was significantly reduced to 622 platelets/mm² for the PTFE catheter and to 2136 platelets/mm² for the commerical catheter. SiO₂ nanosphere coated-modified PTFE catheter adsorbed fewer platelets compared with the PTFE catheter, indicating that coated SiO₂ nanosphere not only significantly reduced platelet deposition but also efficiently suppressed the activation and transmutation of platelets.

## 3.4 Antibacterial effect in vitro

After the catheter enters the blood vessel, some nonspecific proteins (polyglycan matrix, etc.) will form a thin film on the surface of the catheter. Bacteria will then stick to the film and

increase, following a biofilm formed. Because the resulting biofilm is not recognized by tissues and the bacteria are not killed by the immune system, it can eventually lead to bacterial infections. The superhydrophobic surface is an effective method to limit the formation of biofilm, which can inhibit the adhesion of bacteria to the material surface.

The bactericidal effect on Gram-negative bacteria (*Escherichia coli*, *E. coli*) and Gram-positive bacteria (*Staphylococcus aureus*, *S. aureus*) was evaluated to determine whether the superhydrophobic SiO₂ nanosphere coating had a pronounced antibacterial effect. As shown in Figure 6, after co-culture of PTFE-SiO₂ catheters on the solid medium of two bacteria, bacteriostatic circles appeared, and the areas of bacteriostatic processes were 4.05 and 8.75 mm², respectively. Compared the antibacterial effects of the three groups of samples, it can be found that the commercial group samples did not show evident antibacterial rings after co-culture on the two kinds of colony media, and the PTFE catheter only showed small antibacterial
<figure><img src="image_47.png" /><figcaption>**FIGURE 7**
Quantitative analysis of the antibacterial properties of catheter: (A) Conoly-counting assay of catheter against *E. coli* and *S. aureus*; (B) Antibacterial rate of catheter after 6 h incubation with *E. coli* and *S. aureus* (quantitative analysis of colony statistics); (C) Quantitative analysis of photometric method after 6 h incubation with *E. coli* and *S. aureus*. (n = 5, \*p < 0.05, \*\*p < 0.01)</figcaption></figure>

rings on the surface of the medium with *S. aureus*. These results indicate that the superhydrophobic conduit coated with SiO₂ nanospheres has specific antibacterial properties.

We also compared the bacteriostatic rate of the three groups of samples by colony statistical test. As shown in Figure 7A, *E. coli* and *S. aureus* were cultured for 6 h, respectively. Then, a certain amount of bacterial liquid was drawn and placed in a sterile centrifuge. After adding the three samples to the corresponding centrifuge tubes, the centrifuge tubes were placed in a constant temperature shaker for co-cultivation for a certain period of time, and the bacterial solution was diluted 10,000 times. Then take 10 µL of the diluted bacterial culture solution, spread it evenly on the surface of the petri dish containing the solid medium in one direction, place it in a constant temperature shaker for 6 h, and count the number of colonies. After superhydrophobic modification using SiO₂ nanospheres, the antibacterial rate of the catheter against *E. coli* was enhanced from 39.7% to 52.1%, and the antibacterial rate against *S. aureus* was increased from 50.5% to 89.3% (as shown in Figure 7B). It shows that the coating using SiO₂ nanospheres enhances the antibacterial activity of the PTFE

catheter, which may due to the superhydrophobic coating of SiO₂ nanospheres not being conducive to the formation of the bacterial film.

To illustrate that the modified PTFE catheter has prominent antibacterial properties, the antibacterial properties were quantitatively analyzed by photometry. As shown in Figure 7C, after co-culturing the catheter with the bacterial solution for 6 h, the absorbance of the commercial group and the PTFE group was significantly higher than that of the PTFE-SiO₂ group, regardless of whether it was *E. coli* and *S. aureus*. It shows that the catheter of the commercial group does not have antibacterial properties, which is consistent with the inhibition zone and colony statistics results. In conclusion, the antibacterial ability of the PTFE catheter was significantly improved after the layer modification.

Surface functionalization is an effective strategy to improve the application potential of medical catheters. Proper surface wettability is the key to improve the antibacterial ability of the catheter (Chauhan et al., 2014). The mechanism of antibacterial activity of SiO₂ nanosphere coated tough catheter against *S. aureus* was examined by SEM images. As in Figure 8, SEM images showed large numbers of
<figure><img src="image_48.png" /><figcaption>**FIGURE 8**
SEM images of *S. aureus* on the surfaces of commerical catheter, PTFE catheter, and PTFE-SiO₂ catheter, respectively. (Yellow solid represent bacteria in biofilm).</figcaption></figure>

*S. aureus* on the surfaces of commerical catheter. On the surface of PTFE catheter, SEM images show the number of viable *S. aureus* to be greatly reduced, confirming the anti-*S. aureus* nature of the surface of electronegativity of PTFE. It is worth noting that although PTFE-SiO₂ catheter has a large contact angle, it also prevents the formation of *S. aureus* biofilm, possibly because of the large surface anionic charge of SiO₂, which were uniformly distributed on the PTFE surface.

## 4 Conclusion

In the present study, we developed a SiO₂ nanosphere-coated PTFE catheter (PTFE-SiO₂) based on hydroxyl-rich materials obtained by hydroboration oxidation. SEM, FTIR, water contact angle, mechanical tests, cell and blood compatibility test in vitro were used to analyze the chemical structure, microstructure, surface wettability, cytocompatibility, and hemocompatibility of PTFE-SiO₂ catheter. The results show that the SiO₂ nanospheres can be uniformly dispersed on the surface of the PTFE catheter, and the coating will not significantly change the bending resistance of the catheter. Moreover, the catheter demonstrated a safe hemolysis rate of less than 5% while also not causing coagulation. Finally, antibacterial experiments showed that PTFE-SiO₂ catheter was the group with the highest antibacterial rate among the three groups. In conclusion, the SiO₂ nanospheres-coated PTFE catheter with regular morphology and high bending resistance property has a potential application prospect in anticoagulant and anti-infective catheters.

## Data availability statement

The original contributions presented in the study are included in the article/Supplementary Material, further inquiries can be directed to the corresponding author.

## Author contributions

WZ was responsible for the design of the experimental scheme and the writing of the paper; JD was responsible for the processing of experimental data; TZ was responsible for the revision of the paper; RW was responsible for the data review and revision of the paper.

## Acknowledgments

The authors sincerely appreciate projects sponsored by Technology Support Project of Science and Technology Commission of Shanghai Municipality (19441906700).

## Conflict of interest

The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

## Publisher's note

All claims expressed in this article are solely those of the authors and do not necessarily represent those of their affiliated organizations, or those of the publisher, the editors and the reviewers. Any product that may be evaluated in this article, or claim that may be made by its manufacturer, is not guaranteed or endorsed by the publisher.

## Supplementary material

The Supplementary Material for this article can be found online at:
https://www.frontiersin.org/articles/10.3389/fbioe.2022.1067139/full#supplementary-material
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<figure><img src="image_49.png" /></figure>

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OPEN ACCESS

EDITED BY

Muhammad Wajid Ullah,
Jiangsu University, China

REVIEWED BY

Meifeng Zhu,
Nankai University, China
Arunkumar Palaniappan,
Vellore Institute of Technology (VIT), India

*CORRESPONDENCE

Pinhua Rao,
raopinhua@sues.edu.cn
Jiale Jin,
22018143@zju.edu.cn
Yi Chai,
chaiyi@renji.com

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This article was submitted to Biomaterials,
a section of the journal
Frontiers in Bioengineering and
Biotechnology

RECEIVED 06 November 2022

ACCEPTED 22 December 2022

PUBLISHED 12 January 2023

CITATION

Luo J, Sun F, Rao P, Zhu T, Liu Y, Du J, Chen S, Jin X, Jin J and Chai Y (2023), A poly (glycerol-sebacate-acrylate) nanosphere enhanced injectable hydrogel for wound treatment.
*Front. Bioeng. Biotechnol.* 10:1091122.
doi: 10.3389/fbioe.2022.1091122

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© 2023 Luo, Sun, Rao, Zhu, Liu, Du, Chen, Jin, Jin and Chai. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.

A poly (glycerol-sebacate-acrylate)
nanosphere enhanced injectable
hydrogel for wound treatment

Jiajia Luo¹, Fenglei Sun², Pinhua Rao¹*, Tonghe Zhu¹,
Yonghang Liu¹, Juan Du¹, Sihao Chen¹, Xiangyun Jin³, Jiale Jin⁴*
and Yi Chai⁵*

¹School of Chemistry and Chemical Engineering, Institute for Frontier Medical Technology, Shanghai Engineering Research Center of Pharmaceutical Intelligent Equipment, Shanghai Frontiers Science Research Center for Druggability of Cardiovascular Non-coding RNA, Shanghai University of Engineering Science, Shanghai, China, ²Department of Neurosurgery, Weifang People's Hospital, Weifang, Shandong, China, ³Department of Orthopaedics, Renji Hospital, School of Medicine, Shanghai Jiao Tong University, Shanghai, China, ⁴Spine Lab, Department of Orthopaedic Surgery, The First Affiliated Hospital, Zhejiang University, Hangzhou, China, ⁵Department of Neurosurgery, Renji Hospital, School of Medicine, Shanghai Jiao Tong University, Shanghai, China

Wound repair remains a huge clinical challenge, which can cause bleeding, infection, and patient death. In our current research, a bioactive, injectable, multifunctional composite hydrogel doped with nanospheres was prepared with antibacterial and angiogenesis-promoting functions for the treatment of wounds. Amino groups in ε-polylysine (ε-EPL) undergo dynamic Schiff base reaction cross-linking with oxidized hyaluronic acid (OHA), and F127 exhibits unique temperature sensitivity to form an injectable thermosensitive hydrogel (FHE10), which can form a hydrogel to cover the wound at body temperature. Nanospheres (PNs) prepared using poly (glyceryl-sebacate-acrylate) (PGSA) were loaded into hydrogels (FHE10) for promoting wound repair. The prepared FHE10 exhibited rapid gelation, good injectable abilities, and showed resistance to the flourish of *Escherichia coli* (*E. coli*) and *Staphylococcus aureus* (*S. aureus*). *In vitro* investigations showed that FHE10 had good hemocompatibility and cytocompatibility. FHE10@PNs exhibited good proliferation, migration, and tube formation of human umbilical vein endothelial cells (HUVECs) and human foreskin fibroblasts (HFF-1). Furthermore, FHE10@PNs significantly promoted reepithelialization and collagen deposition as well as micro-vascularization compared with the use of FHE10 or PNs alone, thereby accelerating the repair of wounds. In general, this study demonstrated that the multifunctional injectable composite hydrogel showed great potential in wound treatment.

KEYWORDS

poly (glycerol-sebacate-acrylate) nanosphere, injectable hydrogel, antibacterial, vascularization, wound treatment

1 Introduction

Poly (glycerol-sebacate) (PGS) elastomer (Wang Y D et al., 2002) is obtained by polycondensation of two natural metabolic intermediates, glycerol, and sebacic acid, and have good biocompatibility and biodegradability. PGS can be used as scaffolds for cardiovascular (Rai et al., 2015), nerve (Saudi et al., 2019), cartilage (Souza et al., 2017), and bone (Piszko et al., 2021) tissue engineering, and has good applications in drug delivery systems (Hsieh et al., 2018) and wound repair (Keirouz et al., 2020). However, the traditional thermal crosslinking curing method requires long-term high-temperature reaction conditions
(Wang Y D et al., 2002), and the cured product cannot be further processed, which leads to its limited application in biomedicine. PGS is modified in the presence of acrylate groups to form polyglyceryl sebacate acrylate (PGSA) (Nijst et al., 2007). PGSA has the ability of fast photocuring in the presence of photoinitiators, which not only solves the high-temperature problem of the reaction but also greatly reduces response time. PGSA also has good biocompatibility and biodegradability and can be used as a scaffold for tissue engineering such as heart and nerve, as well as the ability to enhance tissue vascular reconstruction and promote wound repair (Jiang et al., 2021).

Hyaluronic acid (HA) is a polysaccharide derived from the extracellular matrix in tissues, which is vital in angiogenesis (Kang et al., 2019) and wound healing (Wu et al., 2017). A variety of hydrogel scaffolds were prepared from hyaluronic acid for tissue regeneration and proved to be able to absorb body fluids rapidly, with good water retention, biodegradability, and biocompatibility (Rezaeeyazdi et al., 2018). Studies have shown that some peptides have obvious antibacterial properties (Sathoff et al., 2018). Shime and Sakai (Shima and Sakai, 1977) first discovered in the 1970s that *Streptomyces* Albicans can produce ε-polylysine, a natural polypeptide composed of L-lysine monomers, which has biodegradable properties, good water solubility, non-toxic, and other advantages. ε-EPL can be used in the food industry as an antibacterial agent (Ge et al., 2022), and in recent years, its application in biomedicine has gradually increased (Shukla et al., 2012), such as drug carrier (de la Torre et al., 2018), interferon inducer (Wakamoto et al., 2007), dental adhesive (Xie et al., 2019), wound dressings (Sun et al., 2020), etc. F127 has received more and more attention as a drug delivery system, biomaterial, and tissue engineering hydrogel thanks to its marvelous biocompatibility and temperature sensitivity (Park et al., 2006), but owing to its insufficient mechanical properties and rapid degradation, limiting its appliance in regenerative medicine. By modifying it or combining it with other substances, it can be widely used after improving the above shortcomings (Li et al., 2020). The Schiff base reaction is a chemical reaction that forms dynamic covalent imine bonds through the cross-linking of amine groups and aldehyde groups (McKay and Finn, 2014). The dynamic cross-linked network are important to form self-healing hydrogels in a mild way, which can make the self-healing hydrogels automatically restore their original structure and function after damage (Xu et al., 2019).

Whether it is disease or physical stimulation, it may cause skin damage, and even affect the life and health of patients in severe cases. Therefore, the development of effective wound dressings is crucial for clinical treatment. The wound dressing should have hemostatic function, keep the tissue isolated from the external environment, prevent bacterial infection, and promote the repair of wound. The wound is easy got infected by bacteria in the procedure of wound healing which leads to the long-term failure of the wound to heal, which may lead to an inflammatory response and affect the normal function of the infected tissue (Holubová et al., 2021). Therefore, the preparation of materials with antibacterial properties that can promote wound healing is crucial. The important courses of wound healing include migration, proliferation, deposition of cell, and formation of extracellular matrix (Telgenhoff and Shroot, 2005; Jie et al., 2007), all of which require the transport of oxygen, nutrients, and cytokines with the help of blood vessels, so the building of new blood vessels in wound is crucial (Michalczyk et al., 2019). Injectable

hydrogels (Xuan et al., 2021) are promising candidate materials with structures like the extracellular matrix, which can be injected to cover irregular wounds, absorb body fluids at the wound surface, and isolate the external environment.

In this study, we used HA to form oxidized hyaluronic acid (OHA) by modifying it with sodium periodate. At room temperature or physiological conditions, OHA interacts with ε-EPL through Schiff base reaction to form hydrogels. F127 is temperature-sensitive and automatically form a gel at room temperature or under physiological conditions. After mixing the above three substances uniformly, an injectable multifunctional hydrogel with temperature-sensitive properties can be formed at room temperature or under physiological conditions. PGSA was synthesized by modifying PGS with acryloyl chloride. Under the action of the photoinitiator, PGSA undergoes free radical polymerization to form nanospheres. The nanospheres were doped into the hydrogel to prepare an injectable and multifunctional wound repair hydrogel dressing for promoting angiogenesis.

# 2 Materials and methods

## 2.1 Materials

Glycerol was provided by China National Pharmaceutical Group Co. (Shanghai, China). Sebacic acid (SA), acryloyl chloride, ethyl acetate (EA), ethanol, ethylene glycol, and triethylamine were purchased from Adamas Reagent, Ltd. (Shanghai, China). DMAP, I2959, is referred to Irgacure 2959 and Dichloromethane (DCM) were supplied by Aladdin Reagent Co. (Shanghai, China). Sodium hyaluronate was provided by Shanghai Yuanye Biotechnology Co. (Shanghai, China). Sodium periodate (NaIO4), F127, is referred to Pluronic F127 and ε-polylysine (EPL) were gained from Aladdin Reagent Co. (Shanghai, China). Deionized water was self-made by laboratory instruments. All materials and solvents were used as received without any further purification unless otherwise noted.

## 2.2 Methods

### 2.2.1 Synthesis of poly (glyceryl-sebacate-acrylate) (PGSA) and preparation of PGSA nanospheres

The synthesis of poly (glycerol-sebacate) (PGS) was according to Wang's research (Wang Y D et al., 2002), then synthesized poly (glyceryl-sebacate-acrylate) (PGSA) by modifying PGS with acryloyl chloride (Nijst et al., 2007). The structure of PGS and PGSA was verified by ¹H nuclear magnetic resonance (¹H-NMR) and Fourier Transform Infrared (FTIR) spectrometry. The acylation degree of PGSA was computed from the ¹H-NMR according to previous study (Chen et al., 2018).

The preparation method of nanospheres was as follows: At first, 30 mg PGSA was mixed in a 3 mL of ethanol and distilled water solution ($V_E: V_W = 1:2$) to dissolve into a homogeneous solution. Photoinitiator solution was prepared by dissolving I2959 powder with deionized water, then 40 µL I2959 (2 mg/mL) was mixed with the above PGSA solution. Magnetic stirring was performed while irradiating the above solution with a UV lamp, and PGSA nanospheres (PNs) were obtained after 30 s.
## 2.2.2 Synthesis of oxidized hyaluronic acid (OHA) and preparation of injectable hydrogels

Hyaluronic acid (HA) was oxidized to oxidized hyaluronic acid (OHA) by sodium periodate (NaIO₄) (Hongyi et al., 2022). The structure of OHA was tested by ¹H-NMR and FTIR spectrometry. The oxidation degree of OHA was obtained through the hydroxylamine hydrochloride method (Hongyi et al., 2022). The proportion of aldehyde groups in the OHA molecular chain, is calculated as following:

$$ 
\begin{aligned}
\text{HA} - (\text{CHO})_2 + 2\text{NH}_2 - \text{OH} \cdot \text{HCl} & \rightarrow \text{HA} - (\text{CH} = \text{N} - \text{OH})_2 + 2\text{HCl} \\
\text{HCl} + \text{NaOH} &= \text{NaCl} + \text{H}_2\text{O} \\
n &= \Delta V \times \frac{C_{\text{NaOH}}}{2} \\
W &= n \times M_{\text{HA-CHO}} + m \times M_{\text{HA}}
\end{aligned}
$$

$$ 
\begin{aligned}
\text{oxidation degree (100\%)} &= n/(n+m) \\
&= n/[(W - n \times M_{\text{HA-CHO}})/M_{\text{HA}} + n]
\end{aligned}
$$

$\Delta V$ (mL) is the volume of sodium hydroxide solution consumed when adjusting the pH of the reaction solution; n (mol) is the number of repeating units of aldehyde groups on the OHA molecular chain; W (g) is the mass of OHA; $M_{HA-CHO}$ = 375 g/mol is the molecular weight of the OHA repeating unit; $M_{HA}$ = 379 g/mol is the molecular weight of the hyaluronic acid repeat unit.

The preparation of hydrogels was referred to Wang's research (Wang et al., 2019), briefly summarized as follows: OHA was freeze-dried and dissolved into an 80 mg/mL solution by using distilled water. $\epsilon$-EPL was dissolved into a solution by using the same method with a concentration of 50 mg/mL and 100 mg/mL. Then F127 was dissolved into a 400 mg/mL solution under 4°C. According to the volume ratio of F127: $\epsilon$-EPL: OHA as 3:0:0, 3:0:1, 3:1:1, the F127 solution and the $\epsilon$-EPL solution were sequentially mixed at 4°C, and the OHA solution was added after mixing evenly, then the solution was put in a thermostatic shaker for gelation, hydrogels were named as F127, FH, FHE5, FHE10, respectively.

PNs combined with FHE10 hydrogel were prepared by the above procedure. After mixing the F127 solution and the $\epsilon$-EPL solution at 4°C, PNs were dispersed in the mixing solution. Then the OHA solution was mixed with above solution at 37°C to obtain the hybrid hydrogel.

## 2.2.3 Characterizations and testing

The microstructure of PNs was observed by SEM and TEM, then the diameter of PNs was evaluated by ImageJ. The study of particle size and z-potential was using a DLS. The injectability of PNs was tested by needle injection to observe whether it can be extruded smoothly.

Using SEM to observe the microstructure of hydrogels, and the pore size was measured via ImageJ. Testing whether the FHE10 can be successfully extruded from a needle proves its injectability. Phosphate buffer solution (PBS) was dropped into the test tube with hydrogels several times until the volume no longer changes and then weighed the mass. The hydrogel was lyophilized and weighed, and its water absorption ratio was computed through formula (1):

$$ \text{Water absorption ratio (\%)} = \frac{W - W_0}{W_0} \times 100\% \quad (1) $$

When hydrogel was at swelling equilibrium, the weight was W, and when it was lyophilized, the weight was $W_0$.

The prepared hydrogel was weighed and placed in an environment of 37°C with a relative humidity of 70% and then the weight of it was recorded every 1 h. The water retention ratio was computed through formula (2):

$$ \text{Water retention ratio (\%)} = \frac{W_X - W}{W_0 - W} \times 100\% \quad (2) $$

In the start, the weight of the hydrogel was $W_0$, $W_X$ was the weight of the hydrogel at the X hour, and after lyophilized the weight was W.

The free radical scavenging rate of hydrogel was measured by DPPH assay to evaluate its antioxidant capacity. The fresh 0.1 mM 1,1-diphenyl-2-trinitrophenylhydrazine (DPPH)/ethanol solution was prepared in the darkness. 30 mg of hydrogel was mixed with 3 mL of DPPH solution and cultured for 3 h and 24 h in the dark. Then a UV-Vis spectrophotometer was applied to measure the absorbance of the solution at 517 nm. The scavenging rate was computed by formula (3):

$$ \text{Scavenging rate (\%)} = \frac{A - A_0}{A_0} \times 100\% \quad (3) $$

The absorbance of the control group was $A_0$, and A was the absorbance of the samples incubated with DPPH solution after 3 and 24 h.

## 2.2.4 Blood compatibility test in vitro

Blood compatibility test was carried out according to the method of GB/T 16886.4-2003 of China. The procedure is described below. 4 mL of carotid artery blood was collected from healthy SD rats, and anticoagulated whole blood was prepared using aqueous trisodium citrate solution, then freshly diluted anticoagulated whole blood was prepared using PBS buffered solution. The materials F127, FH, FHE5 and FHE10 were placed in centrifuge tubes, 3 mL of PBS solution was added to each of them, and the shaker's temperature was set at 37°C. Then put the centrifuge tubes in the shaker, and incubated for 30 min. After 60 µL of freshly diluted anticoagulated whole blood was added to each of them, the incubation was continued for 1 h. The tubes were centrifuged at 2000 rpm/min for 5 min to get the supernatant, and then test the supernatant at 545 nm to get the absorbance value L. The same volume of distilled water and diluted anticoagulated whole blood was added to the tube as positive control group and the absorbance was tested as M. The same volume of PBS buffer solution and diluted anticoagulated whole blood was added to the tube as negative control group and the absorbance was tested as N. Three parallel samples were tested in each group. The hemolysis ratio was calculated by formula (4):

$$ \text{Hemolysis ratio (\%)} = \frac{H_X - H_N}{H_P} \times 100\% \quad (4) $$

$H_X$ was the absorbance of the experimental group, and $H_N$, $H_P$ was the absorbance of the negative and positive group respectively.

## 2.2.5 Antibacterial performance test

Escherichia coli and Staphylococcus aureus were selected to evaluate the antibacterial performance of FHE series hydrogels. The procedure was as follows: 0.5 g of each of F127, FH, FHE5, and FHE10 hydrogels were accurately weighed in a tube and sterilized under UV for 30 min 10 µL of bacterial suspension was placed on top of the hydrogels, and after 12 h of contact, 10 mL of PBS solution was added to the tube, vortexed for 2 min, and then 10 µL of the above
vortexed solution was added to the culture medium, and the bacterial solution was evenly coated on the sheep blood agar (SBA) plates using the spread-plate method and placed in a constant temperature incubator for 24 h. After the culture was completed, the antibacterial effect of the hydrogel was evaluated by observing the growth of the photographed bacteria.

### 2.2.6 *In vitro* experiments

To assess the cytocompatibility of the materials, we selected human umbilical vein endothelial cells (HUVECs) and human foreskin fibroblasts (HFF-1). HUVECs and HFF-1 were purchased from Shanghai Cell Bank of Chinese Academy of Sciences (Shanghai, China). HUVECs were taken as an example to illustrate the culture of cells. HUVECs were grown in the culture flasks in endothelial cell medium (ECM) with 5% fetal bovine serum (FBS), 1% penicillin/streptomycin (P/S) and 1% endothelial cell growth supplement (ECGS). The flask was placed in a CO₂ cell incubator.

The proliferation of cells co-cultured with materials was evaluated by a Cell Counting Kit-8 (CCK-8) assay. The appropriate amount of lyophilized samples F127, FH, FHE5, and FHE10 were weighed, and a solution with a concentration of 10 mg/mL by soaking in serum-free medium for 24 h were prepared. HUVECs were seeded on a 24-well plate at 1×10⁴ per well and placed in a CO₂ cell incubator for 24 h, then the ECM medium was replaced by the ECM medium containing materials to culture cells for 1, 3, and 5 days. At a set point in time, 450 µL of ECM medium and 50 µL of the pre-warmed CCK-8 solution was placed to each well, then placed in a cell incubator for 1 h. After that, the solution was added to a 96-well plate, and then a microplate reader was used to survey the absorbance value at 450 nm wavelength.

The cytoskeletal proteins and nuclei were labeled by FITC-labeled phalloidin and 4',6-diamidino-2-phenylindole (DAPI) respectively. The same samples were used as the CCK-8 assay. The nucleus staining was visualized under the DAPI channel (excitation = 405 nm; emission = 437-552 nm), and cytoskeletal proteins were visualized under the GFP channel (excitation = 488 nm, emission = 505-545 nm). HUVECs were planted in a 24-well plate at a density of 2×10⁴ per well, and placed in a CO₂ cell incubator for 3 days. The cells were first washed three times with ice PBS for 15 min, and then covered with 4% cold paraformaldehyde for 15 min. For permeabilization, cells were treated with 0.1% Triton X-100 for 15 min and blocked with 1% BSA for 1 h. Finally, cells were incubated with primary antibodies at 4°C overnight. The cells were washed three times with PBS for 15 min among each step. After primary antibody incubations, cells were washed for 3 × 5 min in PBS-T (0.1% Tween 20%). Subsequently, fluorescent-labeled secondary antibody was used to incubated with cells for 2 h in dark and then DAPI was used on the cells for 15 min. All steps were washed with PBS-T. The results were observed *via* a confocal microscope.

Cell migration was tested by the transwell assay to measure the influences of materials, taking HUVECs as an example. HUVECs were planted in the upper chamber of a 24-well transwell, and 5×10⁴ cells were plated in each well. 600 µL of ECM medium containing materials were placed to the lower chamber and the ECM medium to the control group. The same samples were used as the CCK-8 assay. The culture plate was taken out after 8 h of culture and the migrated cells in upper chamber were washed with PBS for 3 times and fixed with 4% paraformaldehyde for 15 min, and stained with 0.1% crystal violet staining solution for 5 min and dried overnight. The cells were

<table><caption>TABLE 1 Primer sequences used for qRT-PCR.</caption><thead><tr><th>Gene</th><th>Gene bank</th><th>Primer sequences (5'-3')</th><th>Tm (°C)</th></tr></thead><tbody><tr><td rowspan="2">VEGF</td><td>NCBI</td><td>F: AGGGCAGAATCATCACGAAGT</td><td>61.2</td></tr><tr><td>geneID7422</td><td>R: AGGGTCTCGATTGGATGGCA</td><td>62.9</td></tr><tr><td rowspan="2">PDGF</td><td>NCBI</td><td>F: CTCGATCCGCTCCTTTGATGA</td><td>61.7</td></tr><tr><td>geneID5155</td><td>R: CGTTGGTGCGGTCTATGAG</td><td>60.2</td></tr><tr><td rowspan="2">bFGF</td><td>NCBI</td><td>F: AGAAGAGCGACCCTCACATCA</td><td>62.7</td></tr><tr><td>geneID2247</td><td>R: CGGTTAGCACACACTCCTTTG</td><td>61.2</td></tr><tr><td rowspan="2">GAPDH</td><td>NCBI</td><td>F: ACAACTTTGGTATCGTGGAGG</td><td>62.1</td></tr><tr><td>geneID2597</td><td>R: GCCATCACGCCACAGTTTC</td><td>61.1</td></tr></tbody></table>

observed and taken pictures under a microscope. Counting the number of migrated cells using ImageJ.

The materials' ability to promote HUVECs angiogenesis was tested by using Matrigel *in vitro*. 100 µL of well-mixed Matrigel was placed in the 48-well plate, then the plate was put in a cell incubator for 30 min. Each well of the plate was seeded with 3×10⁴ cells and ECM with materials was added to it, the control group was contained with ECM medium. Then the well plate was put in a cell incubator for 6 h. After this, the vascularized network structure on the surface of Matrigel was observed and photographed using a light microscope, and vessels were counted by ImageJ.

To explore the effects of materials on the gene expression of vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF), and basic fibroblast growth factor (bFGF), HUVECs and HFF-1 were placed on 6-well plates at a density of 4 × 10 per well. After 3 days, RNA was collected from different cell samples with TRIzol reagent. qRT-PCR was then achieved with ChamQ SYBR qPCR Master Mix. The gene expression levels were normalized to GAPDH, and the expression of related genes was analyzed by the 2⁻⁰⁻Δt method. The sequences of target genes and internal reference genes are shown in Table 1.

### 2.1.7 *In vivo* experiments

All animal experiments were authorized by the Animal Research Ethics Committee of the First Affiliated Hospital of Zhejiang University and were performed by the National Institutes of Health Guidelines for the Care and Use of Laboratory Animals. Forty healthy rats were randomly selected as experimental rats and divided into four groups, which were marked as the control group, PNs group, FHE10 group, and FHE10@PNs group. A rat skin defect model was established based on previous research methods (Wang et al., 2013). Each rat was weighed and 4%Pentobarbital (40 mg/kg) was used to anesthetize by intraperitoneal administration. After the injection, the vital signs of the rat were noted. Two symmetrical circular marks with a diameter of 2 cm were made on the depilation area on the back of the rat, and the full-thickness skin was cut off. The bleeding was compressed by the gauze until there was no obvious bleeding. In the next step, the wound was covered with materials and wrapped with gauze. After the operation, penicillin was injected to prevent infection, and the health of the rats was closely observed. Rats in each group were sacrificed 7 and 14 days after surgery, and the skin tissues at the edge of the wound were collected for immunohistochemical analysis.
<figure><img src="image_50.png" /><figcaption>FIGURE 1</figcaption></figure>

(A) The reaction mechanism route of PGSA; (B) ¹H-NMR spectrum of PGS; (C) ¹H-NMR spectrum of PGSA. (D) The reaction mechanism route of OHA; (E)

¹H-NMR spectrum of HA and OHA; (F) FTIR spectrum of HA and OHA. (G) Apparent morphology of PNs, as revealed by digital images; (H) Digital photographs of PNs dispersed in deionized water; (I) and (J) PNs were uniformly dispersed in deionized water under natural light, and Tyndall effect appeared after beam irradiation; (K) The diameter distribution histograms of PNs; (L) SEM image of PNs; (M) TEM image of PNs; (N) Particle size and potential of PNs; (O) and (P) Schematic illustration of the injectability of PNs. The red arrow means PNs were uniformly dispersed in the syringe.

The healing of skin wounds on days 0, 3, 7, and 14 after the operation was recorded by a camera, and the area of the wound was calculated by ImageJ. The wound healing ratio was computed by formula (5):

$$ \text{Wound healing ratio (\%)} = \frac{S_0 - S}{S_0} \times 100\% \qquad (5) $$

Wound area on day 0 was represented by S₀, and the wound area on days 3, 7, and 14 was represented by S.

The skin tissues were completely immersed in 10% formalin solution overnight, then in graded alcohol (75%, 85%, 95%, 100%) for dehydrating. Paraffin was used to embedded the cleared tissues, and the tissue block was cooled overnight then cut into 5 µm of sections.

The regeneration of wound epidermis was assessed by staining the tissue sections with hematoxylin and eosin (H&E). H&E staining was carried out according to previous study (Yoon et al., 2016). For H&E staining, paraffin-embedded sections were first deparaffinized and hydrate to water. Then, nucleus was stained with Hematoxylin Solution (Solarbio, China) for 15 min followed by rinsing in running tap water. Next, the tissue section was differentiated by Differentiation Solution (Solarbio, China) for 3 min and re-dyed with Eosin Y Aqueous Solution (Solarbio, China) for 1 min. After the sections were dehydrated and sealed, photos were taken and analyzed. The wound reepithelialization ratio was calculated using formula (6):

$$ \text{Reepithelialization ratio (\%)} = \frac{L_0 - L}{L_0} \times 100\% \qquad (6) $$

L₀ represents the original wound length, and L represents the incompletely healed wound epidermis length.

Masson's trichrome staining was accomplished on tissue sections to evaluate the collagen deposition. Masson's trichome staining were carried out according to previous study (Zhao et al., 2017). For Masson's trichome stain, a Masson trichrome staining kit (Sigma-Aldrich, America) was used to stain the tissue section according to the manufacturer's protocol. To observe and photograph the stained sections with an optical microscope. The area of the blue part was counted and quantitatively analyzed using ImageJ to evaluate the collagen deposition rate in the new tissue.

CD31 immunohistochemical staining and CD31/α-SMA immunofluorescence staining were performed on tissue sections to evaluate angiogenesis in wound tissue. CD31 immunohistochemical staining were carried out according to previous study (Kong et al., 2018). As for CD31 immunohistochemical staining: A two-step immunohistochemistry kit (Zhongshan Golden Bridge Biotechnology, China) was used for this study according to the manufacturer's protocol. To observe and photograph the stained sections with a light microscope, and two pathologists were asked to count the number of blood vessels in the stained sections under double-blind conditions.
<figure><img src="image_51.png" /></figure>

<figure><img src="image_52.png" /></figure>

<figure><img src="image_53.png" /></figure>

<figure><img src="image_54.png" /><figcaption>FIGURE 2</figcaption></figure>

(A) Schiff base reaction mechanism route; (B) Photographs of F127 and FHE10 hydrogel before and after gelation, and FHE10 proved its injectability by extruding SUES; (C) The SEM images of F127, FH, HE, FHE5 and FHE10 hydrogel; (D) and (E) The pore diameter distribution of FHE5 and FHE10, respectively; (F) The swelling ratio of F127, FH, FHE5, and FHE10 hydrogel; (G) The water retention ratio of F127, FH, FHE5, and FHE10 hydrogel.

### 2.1.8 Statistical analysis

Three parallel samples were prepared for the experiments, and the mean ± standard deviation was employed as the analysis data. One-way ANOVA was applied to process statistical differences between groups. When *p* < 0.05 (*) *p* < 0.01 (**) the differences had statistical significance.

## 3 Results and discussions

### 3.1 Characterization of PNs and FHE hydrogel

PGS and PGSA were successfully synthesized (Figures 1A–C and Supplementay Figure S1). The ¹H-NMR analysis indicates that the degree of acylation was 32%. OHA was obtained by modification (Figures 1D–F). The oxidation degree of OHA was 12.5% by the method of hydroxylamine hydrochloride (Wang et al., 2019). In the presence of photoinitiator, the aqueous ethanol solution of PGSA was subjected to ultraviolet irradiation and magnetic stirring to form nanospheres PNs with uniform particle size.

Figures 1G, H showed the state of the PNs under dry conditions and the state of dispersion in water, respectively. The PNs nanospheres

are easy to aggregate together in the dry state, which may be caused by their viscosity. When it was put in deionized water, it can be evenly dispersed. Using a beam to irradiate the PNs solution, the Tyndall effect can be observed from Figures 1I, J, indicating that the nanospheres were uniformly dispersed. It can be seen from the SEM image (Figure 1L) that PNs have a relatively smooth surface. Through ImageJ statistical calculation, it can be obtained that the particle size distribution of PNs (Figure 1K) is in the range of 0.2–1.2 µm, mainly in the range of 0.6–0.8 µm, which is relatively uniform. The dispersion of PNs can be seen from the transmission electron microscopy (TEM) image (Figure 1M). There is a certain difference between the particle size range measured by the dynamic light scattering (DLS) (Figure 1N) instrument and the results of scanning electron microscope (SEM). It is speculated that during the test process, the PNs aggregated due to the sedimentation of their weight, and the measured particle size reached several tens of microns. The surface charge of PNs is negative (Figure 1N). The extrusion experiment of a 1 mL syringe with a needle size of 0.45 × 16 mm proved that PNs could be extruded without sticking to the needle tube (Figures 1O, P). As can be seen the white PNs are evenly dispersed in the syringe and do not adhere to the wall. By
<figure><img src="image_55.png" /><figcaption>FIGURE 3</figcaption></figure>

(A) UV absorption of DPPH solution after co-incubation with F127, FH, FHE5, and FHE10 hydrogel; (B) Scavenging rate of F127, FH, FHE5, and FHE10 hydrogel; (C) Hemolysis ratio of F127, FH, FHE5, and FHE10 hydrogel; (D) Photographs of survival bacteria clones on SBA plates after coculturing with F127, FH, FHE5, and FHE10 hydrogel.

extruding on the glass slide, PNs have good injectability and will not block the needle tube.

F127 is a temperature-sensitive material. At low temperature, it is in a solution state, and when the temperature is raised, it transforms into a gel state. After mixing F127 with OHA and EPL, the whole solution has temperature sensitive properties. When the temperature was at 4°C, the mixed solution became a sol state, and when the temperature was raised up to 37°C, it became a gel, so it belonged to a thermosensitive gel (Figure 2B). After the gel was loaded into the syringe, it was slowly extruded into letters such as “SUES” (Figure 2B), and the extrusion was smooth without clogging, which proved that the hydrogel system has good injectable performance. It can be seen from Figure 2C that F127 and FH hydrogel had no pores, while FHE hydrogel had pores. The measurement of pore diameter was by ImageJ, and Figures 2D, E were the pore size of FHE5 and FHE10, respectively. Figure 2F is the swelling ratio of hydrogels. The swelling ratio of FHE5 and FHE10 hydrogel increased significantly, which may be due to the reticular pore structure formed after the chemical reaction, thereby improving the water absorption, and swelling ability of the hydrogel. Due to the prolongation of time, and the inevitable loss of water, the FHE5 and FHE10 hydrogel has a stronger water-locking ability (Figure 2G).

A UV spectrophotometer was used to measure the absorbance of the mixed solution, the UV absorption of the solution after co-

incubation with the material decreased (Figure 3A) and the color faded (Supplementary Figure S2) compared with the control group. The scavenging rate of F127, FH, FHE5, and FHE10 hydrogel reached more than 80% after 24 h (Figure 3B), indicating that the materials have good antioxidant properties. The ability to scavenge free radicals should come from many hydroxyl groups in the F127 molecule. In addition, the molecular chain of OHA also contains many hydroxyl groups, so the hydrogel has a strong ability to scavenge free radicals.

Supplementary Figure S2 showed that except for the control group, the centrifuge tubes containing the material were centrifuged after incubation, the red blood cells were deposited at the bottom without breaking, the supernatant was clarified, and the hemolysis ratio was less than 5% (Figure 3C), proving that the material did not cause hemolysis. The prepared hydrogels, only FHE5 and FHE10 have certain antibacterial properties (Figure 3D). Judging from the growth of bacteria in the control group and the F127 and FH groups, the bacterial colonies on the medium grew well, and it was speculated that the pure F127 hydrogel and FH hydrogel did not have antibacterial effects. Comparing the FHE5 and FHE10 groups, the growth of bacteria is inhibited, which is attributed to the fact that ε-EPL itself has a certain antibacterial activity, and the protonated amino group can damage the cell wall of bacteria and lead to bacterial death.