Cellular and Molecular Interaction of Human Dermal Fibroblasts with

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Biological and Medical Applications of Materials and Interfaces

Cellular and Molecular Interaction of Human Dermal Fibroblasts with Bacterial Nanocellulose Composite Hydrogel for Tissue Regeneration Evelyn Yun Xi Loh, Mh Busra Fauzi, Min Hwei Ng, Pei Yuen Ng, Shiow Fern Ng, Hidayah Ariffin, and Mohd Cairul Iqbal Mohd Amin ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.8b16645 • Publication Date (Web): 29 Oct 2018 Downloaded from http://pubs.acs.org on October 30, 2018

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ACS Applied Materials & Interfaces

Cellular and Molecular Interaction of Human Dermal Fibroblasts with Bacterial Nanocellulose Composite Hydrogel for Tissue Regeneration Evelyn Yun Xi Loh1, Mh Busra Fauzi2, Min Hwei Ng2, Pei Yuen Ng3, Shiow Fern Ng1, Hidayah Ariffin4, Mohd Cairul Iqbal Mohd Amin1, 5, * 1Centre

for Drug Delivery Research, Faculty of Pharmacy, Universiti Kebangsaan Malaysia, Jalan Raja Muda Abdul Aziz, 50300 Kuala Lumpur, Malaysia 2 Tissue Engineering Centre, Universiti Kebangsaan Malaysia Medical Centre, Jalan Yaacob Latif, Bandar Tun Razak, 56000 Kuala Lumpur, Malaysia 3 Drug and Herbal Research Centre, Faculty of Pharmacy, Universiti Kebangsaan Malaysia, Jalan Raja Muda Abdul Aziz, 50300 Kuala Lumpur, Malaysia 4Department of Bioprocess Technology, Faculty of Biotechnology and Biomolecular Sciences, Universiti Putra Malaysia, 43400 UPM Serdang, Selangor, Malaysia 5IDEA-UKM, Universiti Kebangsaan Malaysia, 43600 UKM Bangi, Selangor, Malaysia * Corresponding author: Mohd Cairul Iqbal Mohd Amin Faculty of Pharmacy, Universiti Kebangsaan Malaysia, Jalan Raja Muda Abdul Aziz, 50300 Kuala Lumpur, Malaysia. Phone: +6 03-92897690 Fax: +6 03-2698 3271 E-mail: [email protected] KEYWORDS: cell carrier, scaffold, wound dressing, skin wound healing, cell-material interaction

ABSTRACT The evaluation of the interaction of cells with biomaterial is fundamental to establish the suitability of the biomaterial for a specific application. In this study, the properties of bacterial nanocellulose/acrylic acid (BNC/AA) hydrogels fabricated with varying BNC to AA ratios and electron beam irradiation doses were determined. The manner these hydrogel properties influence the behavior of human dermal fibroblasts (HDFs) at the cellular and molecular levels was also investigated, relating it to its application both as a cell carrier and wound dressing 1

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material. Swelling, hardness, adhesive force (wet), porosity, and hydrophilicity (dry) of the hydrogels were dependent on the degree of cross-linking and the amount of AA incorporated in the hydrogels. However, water vapor transmission rate, pore size, hydrophilicity (semi-dry), and topography were similar between all formulations, leading to a similar cell attachment and proliferation profile. At the cellular level, the hydrogel demonstrated rapid cell adhesion, maintained HDFs viability and morphology, restricted cellular migration, and facilitated fast transfer of cells. At the molecular level, the hydrogel affected nine wound healing genes (IL6, IL10, MMP2, CTSK, FGF7, GM-CSF, TGFB1, COX2, F3). The findings indicate that the BNC/AA hydrogel is a potential biomaterial that can be employed as a wound dressing material to incorporate HDFs for the acceleration of wound healing.

1. INTRODUCTION The cell-biomaterial interaction is an essential evaluation to determine the functionality of a selected biomaterial for a particular biomedical and clinical application prior to in vivo studies1. In tissue engineering, the biomaterial characteristics—including pore size, porosity, hydrophilicity, topography and mechanical strength—determine the microenvironment for the cells, which will be reflected at the cellular level via cell adhesion, proliferation, spreading, migration and differentiation, as well as at the molecular level via regulation of gene expression, influencing the intracellular signaling pathways2,3. Inarguably, different types of applications require different characteristics, and different cells respond differently to these features. The normal surrounding for most cells, including fibroblasts, is the extracellular matrix (ECM), which is typically a type of hydrogel2. Hydrogels are three-dimensional (3D) water-swollen polymeric networks4. Fibroblasts contribute remarkably to the normal wound healing process by 2


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playing a role in the breakdown of fibrin clot, wound contraction, and laying down of new ECM components such as collagen, fibronectin, and hyaluronic acid5. Although most dermal wounds can heal naturally, the healing process can be hampered in the case of partial-thickness and fullthickness wounds, especially when a large area of injury is involved6. When the gold standard of treatment, i.e., split-skin grafting (SSG), is not a viable option because of the high risk of infection, severe pain, or extensive injury, cellular dermal scaffolds can act as a temporary wound cover and assist in the regeneration of the dermis. Several cellular dermal matrices such as TransCyte™ (human neonatal foreskin fibroblasts on nylon mesh with a silicon layer) and Dermagraft® (human neonatal fibroblasts on a polyglactin mesh) are commercially available7. A comparison study on silver sulfadiazine and TransCyte™ showed that wounds treated with TransCyte™ healed faster (average 11.14 days for TransCyte™ versus 18.14 days for silver sulfadiazine) and with less hypertrophic scarring8, while with Dermagraft®, complete wound closure was achieved by week 12 in 30% patients with chronic diabetic foot ulcers compared to 18.3% of control patients9. However, the disadvantages of TransCyte™ and Dermagraft® include high costs, potential risk of rejection, cold chain supply (from transport to storage) and/or short shelf life10. Thus, research into identifying the best dermal matrices are still underway7. Generally, natural polymers such as gelatin, collagen (COL), chitosan, cellulose, fibrin, and alginate are biocompatible, while synthetic polymers such as poly-lactic-co-glycolic acid (PLGA), polyethylene glycol (PEG), polyurethanes, and polycaprolactone (PCL) provide better mechanical and reproducible properties than those of naturally occurring polymers. There is an increasing interest in novel composite biomaterials obtained by the combination of two or more natural and synthetic polymers showing both biocompatibility and mechanical strength11. Lu et al.12 fabricated open and interconnected porous poly(l-lactic acid) (PLLA)-collagen and PLLA3



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gelatin hybrid scaffolds that improved fibroblast cell attachment and proliferation in vitro and promoted the regeneration of both the dermal and epidermal layers, in addition to reducing contraction in vivo. Bonvallet et al.13 prepared 70:30 COL/PCL scaffolds with 160 mm pores that promoted dermal fibroblast attachment, proliferation, infiltration, and ECM deposition. Moreover, the 70:30 COL/PCL scaffolds seeded with fibroblasts stimulated the regeneration of a more normal-appearing tissue compared to acellular scaffolds. Bacterial nanocellulose (BNC) is an attractive polymer because it demonstrates higher purity, crystallinity, degree of polymerization, and tensile strength than plant cellulose14. BNC can be obtained from the fermentation by-product of coconut water by Acetobacter xylinium. Considering the advantages of a natural/synthetic hybrid polymer, BNC in combination with acrylic acid (AA) is used to fabricate hydrogels. The grafting of AA into BNC via electron beam (EB) irradiation without the presence of cross-linking agents, catalysts, or additives eliminates any potential toxic effect. The degree of cross-linking can be controlled by varying the EB irradiation dose and the amount of AA15. Previously, a preliminary study was conducted using a 406035 (40% AA:60% bacterial cellulose (BC):35 kGy EB irradiation dose) hydrogel formulation, in which the fibroblasts did not proliferate and exhibited a round shape morphology in vitro16. Based on these findings, we produced various hydrogel formulations with less AA content in an attempt to improve hydrogel biocompatibility while maintaining its mechanical strength by increasing the EB irradiation dose. We also improved the cell transfer model in vitro. The aim of this study was to investigate the properties of the new BNC/AA hydrogel and its effect on the behavior of human dermal fibroblasts (HDFs) at a cellular and molecular level in vitro. As BNC is non-biodegradable in the

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human body due to the lack of cellulase, we evaluated its suitability to simultaneously function as both a cell carrier and a wound dressing material for skin wound healing application.

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2. MATERIALS AND METHODS 2.1 Preparation of BNC dispersion and determination of BNC fiber diameter BNC dispersion was carried out as previously described by Ahmad et al.15. BNC (1% w/v) was dispersed in distilled water and homogenized using a high-shear mixer (L4R, Silverson, USA) until the particle size reached 70-80 µm as tested by Mastersizer 2000 (Malvern, UK). The dispersed BNC (1 mL) was lyophilized and observed using a Field Emission Scanning Electron Microscope (FESEM Merlin, Zeiss, Germany). Specimens were mounted on aluminium stub and sputter-coated with platinum. The diameter of the BC fiber was measured using Smart Tiff software (version 2).

2.2 Preparation of BNC/AA hydrogel AA solution (Sigma-Aldrich, Czech Republic) was mixed with the dispersed BNC to obtain ratios of 20:80 and 30:70 (AA:BNC). The mixture (25 mL) was pipetted into each petri dish and exposed to a total of 45 or 60 kGy EB irradiation dose with 5 kGy each pass (3 MeV and 1 mA) at the Malaysian Nuclear Agency facility (EPS-3000, Japan). The hydrogels were labeled as 208045, 208060, 307045, and 307060 according to their AA to BNC ratio and irradiation dose. NaOH (1 M) was added to neutralize the hydrogel to achieve pH 7 ± 0.5 prior to autoclavation (121°C, 20 min).

2.3 Characterization of hydrogel 2.3.1 Swelling The swollen hydrogels were cut into 1.5 cm disks with 0.5 cm thickness before they were dried in the oven at 50°C to a constant weight (W1). They were immersed in 20 mL phosphate6


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buffered saline (PBS) (First Base, Singapore) at 37°C and were weighed every hour for the first 8 h and 24 h (W2); the excess buffer on the surface was blotted with tissue paper. The swelling ratios were calculated as follows:

Swelling ratio (%) = [(W2 − W1) / W1]  100

2.3.2 Hardness and adhesive force Hardness of swollen hydrogels was determined using the compression test on a texture analyzer (CT3-10kg, Brookfield, USA). A 36 mm diameter aluminium cylinder probe (TA-AACC36) was pushed at a speed of 0.5 mm/s for a distance of 2 mm into the hydrogel with holding time of 60 s and trigger load of 0.1 N. Three readings at different locations were obtained from each hydrogel. Adhesive force of swollen and dried hydrogels was determined according to the hardness test described above, with the exception of the use of fresh rat skin affixed to the probe. The inside of the skin was in contact with the hydrogel to represent the ECM of the wound.

2.3.3 Porosity Solvent replacement method was used to determine the porosity of the hydrogel. Lyophilized hydrogels were weighed (M1), immersed overnight in absolute ethanol (Merck, Germany), and blotted with tissue paper to remove excess ethanol from the surface before weighing (M2). The dimension of the hydrogel was determined by using a Vernier caliper. Porosity was calculated using the following equation:

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Porosity = (M2 − M1) / ρV Where, M1 and M2 = masses of hydrogel before and after immersion in absolute ethanol, respectively ρ = density of absolute ethanol V = volume of the hydrogel

2.3.4 Water vapor transmission rate (WVTR) Semi-dried hydrogels with a thickness ranging between 1.5 mm and 2.5 mm were placed inside the cap of glass vials containing 20 mL of distilled water, with an opening of 1.2 cm diameter on the cap. The vials were placed in a humidity chamber (Terchy HRM 80 FA, Taiwan) at 37°C and 75% humidity together with controls (glass vials with 20 mL distilled water without any covering at the cap opening). The amount of water evaporated was measured by weighing the vials every day for 5 days, and the WVTR was calculated by dividing the mean daily weight loss of water by the cap opening area.

2.3.5 Pore size Cross-sections of swollen lyophilized hydrogels were observed under a scanning electron microscope (SEM) (LEO 1450 VPSEM, Zeiss, Germany). Specimens were mounted on aluminium stub and coated with gold. Three different locations were captured from each hydrogel, and 40 random pores on each micrograph were measured for their size using the Image-J software (version 1.50i).

2.3.6 Contact angle 8


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Lyophilized and swollen hydrogels that were left at 20–24°C overnight were used to determine the contact angle. Distilled water (10 µL) was carefully dropped onto the surface of the hydrogel, and images were captured using a digital camera with continuous shooting mode (Fujifilm F850EXR). The contact angle was measured by using the Image-J software and enhanced with the ‘find edges’ plug-in.

2.3.7 Topography Topography images and surface roughness of lyophilized hydrogels were acquired using an atomic force microscope (AFM) (NTegra Prima, Russia). The area (10 µm x 10 µm) at three different locations per hydrogel was analyzed using the IA-P9 software (version 3.5.0.2069).

2.3.8 Bacterial Endotoxin Test (BET) BET was performed according to United States Pharmacopeia 39 (USP 39) standards based on the kinematic tubidimetric method. Briefly, two hydrogel disks were soaked in 25 mL sterile water for injection and incubated at 37 ± 1°C for 1 h ± 2 min. The extracting solution was mixed with reconstituted Limulus Amebocyte Lysate (LAL) reagent (Lonza, USA) and placed in an incubating spectrophotometer where turbidity was measured at 340 nm wavelength. The concentration of endotoxin in the hydrogel was calculated from a standard curve covering 5, 0.5, & 0.05 endotoxin unit (EU)/mL concentrations. The endotoxin release limit (ERL) was determined to be 1.6 EU/mL based on the following equation:

ERL = (K  N) / V Where, 9



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K = the endotoxin limit per device (20 EU) N = the number of devices tested (i.e., 2) V = the total volume of water used to extract the endotoxin from the number of devices (N) above (i.e., 25 mL)

2.4 Isolation and culture of HDFs This study was approved by the Universiti Kebangsaan Malaysia Research Ethics Committee (no. FF-2015-376). Redundant abdominoplasty skin tissue samples were obtained—with written informed consent—from three patients, and were processed as reported by Mazlyzam et al.17. The skin was rinsed in Dulbecco’s phosphate-buffered saline (DPBS) (Biowest, USA), cut into small pieces (1–2 cm2), and soaked overnight in 10 mL of serum free Epilife medium (Gibco, USA) containing 25.3 mg of Dispase (Gibco) at 2–8°C to separate the epidermis and dermis layers. The next day, the epidermis layer was removed from the dermis layer. The dermis was minced further and digested with 0.6% collagenase type I (Worthington, USA) for 4 h in an incubatorshaker at 37°C. The cell suspension was then centrifuged at 5,000 rpm for 5 min at 37°C, and the cell pellet was washed with DPBS. Finally, the cell pellet was resuspended in Ham’s F12:Dulbecco’s modified Eagle medium (F12:DMEM; 1:1) (Gibco) supplemented with 10% fetal bovine serum (FBS) (Biowest), 1% antibacterial-antimycotic (Gibco), 1% Glutamax (Gibco) and 2% 4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES) (Gibco). Cells were cultured at 37°C in 5% CO2 with medium changed every 2–3 days. HDFs were subcultured to a maximum of passage 4 upon confluency.

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2.5 Cellular interaction 2.5.1 Cell attachment HDFs (1.14  104/cm2) were seeded on hydrogels of different formulations, which were presoaked in F12:DMEM overnight. The cells were allowed to attach at 37°C with 5% CO2. The hydrogel was washed gently with DPBS every hour for the first 6 h and 24 h. The remaining (unattached) cells in DPBS were counted using a hemocytometer and 0.4% trypan blue solution (Sigma, USA). The percentage of cell attachment was calculated as per the equation below: Cell attachment (%) = [(Initial cell seeding − number of cells in DPBS) / Initial cell seeding]  100

2.5.2 Cell proliferation HDFs (3.40  104/cm2) were seeded on hydrogels of different formulations and on polystyrene cell culture plates. AlamarBlue® Cell Viability Assay (Invitrogen, USA) was conducted according to the manufacturer’s protocol. On day 1, spent medium was removed, the hydrogel was transferred to a new culture well, and fresh medium was added with 10% AlamarBlue® into each well. Negative control wells without cells were included in the experiment. After 4 h of incubation at 37°C, the solution was transferred to a 96-well plate and absorbance was measured at 570 and 600 nm using a spectrophotometer reader (Bio-Tek, Power Wave XS, USA). The samples were washed with DPBS, replaced with fresh medium, and allowed to grow further in the incubator at 37°C with 5% CO2. On day 7, the same procedure was repeated. The percent reduction of AlamarBlue® was calculated according to the manufacturer’s recommendation as below:

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Percent reduction = [(εOX)λ2Aλ1 − (εOX)λ1Aλ2] / [(εRED)λ1A’λ2 − (εRED)λ2A’λ1]  100 Where, (εOX)λ2 = 117,216 (εOX)λ1 = 80,586 (εRED)λ1 = 155,677 (εRED)λ2 = 14,652 Aλ1 and Aλ2 = Observed absorbance reading for test well at 570 nm and 600 nm respectively A’λ1 and A’λ2 = Observed absorbance reading for negative control well at 570 nm and 600 nm respectively

2.5.3 Cell viability HDFs (1.14  105/cm2) were seeded on the 307045 hydrogel and polystyrene cell culture plates. The cells were stained with LIVE/DEAD® Viability/Cytotoxicity kit for mammalian cells (Invitrogen) on day 1 and day 3 according to the manufacturer’s protocol. Cells were washed with DPBS and incubated with 2 µM calcein acetoxymethyl (AM) and 4 µM ethidium homodimer 1-red (EthD-1) working solution in DPBS, followed by observation using a Nikon A1R fluorescence microscope (Nikon, Japan) after 30 minutes of incubation at 37°C with 5% CO2.

2.5.4 Cell morphology HDFs (8.47  104/cm2) were seeded on the 307045 hydrogel. On day 1, 3, and 7, the construct was fixed with 4% glutaraldehyde (Sigma), dehydrated using serial dilutions of ethanol, and

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dried in a critical point dryer (Leica EM CPD300, Germany). The samples were sputter-coated with gold and observed using SEM (LEO 1450 VPSEM).

2.5.5 Cell migration The cytoplasm of HDFs was stained with green cell tracker CMFDA (Invitrogen), while its nucleus was stained with Hoechst Blue Dye (Invitrogen) at a concentration of 5 µM according to the manufacturer’s instructions. HDF (8.47  104/cm2) were seeded on the 307045 hydrogel, followed by 3D confocal imaging on day 0, 3, and 7 using a Nikon A1R confocal microscope.

2.5.6 Cell transfer The ovine collagen solution was prepared as described by Fauzi et al.18. The ovine collagen solution was neutralized with 1 M NaOH at 4°C and allowed to polymerize at 37°C to form the hydrogel that represented the ECM of the wound. HDFs (8.47  104/cm2) were stained with Hoechst Blue Dye as described above and seeded on the 307045 hydrogel. The hydrogel surface, which contained the HDFs, was flipped onto the ovine collagen hydrogel (OCH) after 4 h. On day 1, the BNC/AA hydrogel was peeled off from the OCH, and images of the cells on each of the individual hydrogels were captured. The BNC/AA hydrogel was placed on a new OCH and on the next day, images of the cells on the OCH were captured. This was repeated until day 3. The experiment was performed on both the submerge and airlift (using cell culture insert) models (Figure 1).

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Figure 1: Pictorial representation of the submerge and airlift models used for the cell transfer study.

The total number of cells (in percentage) at ten different locations on the BNC/AA hydrogel and OCH, as observed under a Nikon A1R florescence microscope, was determined using the ImageJ software.

To evaluate if living cells were transferred to the OCH, cell viability assay was performed on day 3 as specified above. The cells on the OCH were observed using a Nikon A1R florescence microscope, and the number (in percentage) of live cells versus dead cells at ten different locations was counted by using the Image-J software.

2.6 Molecular interaction 2.6.1 Total RNA extraction Total RNA from HDFs (1.41  105 cells/cm2) seeded on the 307045 hydrogel and polystyrene cell culture plates after 24 h was extracted using TRI Reagent (Sigma) according to the 14


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manufacturer’s protocol. The wells were washed with DPBS before 1 mL TRI Reagent was added in each well and were left for 30 min. The hydrogels were also washed with DPBS, cut into small pieces, and left in 1 mL TRI Reagent before being probe-sonicated using Sonics Vibra Cell (Sonics & Materials, Inc., USA) at 20% amplitude for 10 s and resting on ice for 1 min. The cycle was repeated 3 times. The solution was placed in a tube and 200 µL of chloroform (per 1 mL of Tri Reagent) (BDH, UK) was added. The solution was vigorously shaken for 15 s and left at 20–24°C for 10 min before centrifugation at 10,000 rpm for 15 min. The aqueous phase was carefully transferred into a new tube where 500 µL of isopropanol (per 1 mL of Tri Reagent) (BDH, UK) was used to precipitate the RNA. Polyacryl carrier (Molecular Research Center Inc., USA) (5 µL) was added to the mixture, flicked to mix, and left at 20–24°C for 10 min before centrifugation at 10,000 rpm for 8 min. The supernatant was discarded, and the RNA pellet was washed with 1 mL of 75% ethanol, followed by centrifugation at 7,500 rpm for 5 min. The supernatant was discarded, and the pellet was air-dried for 20 min. Finally, the RNA was dissolved in RNase-free water. All centrifugation steps were performed at 4°C. The concentration and purity of the RNA was measured using a spectrophotometer (Take 3, BioTek), while the integrity of the RNA was verified by 1.5% agarose gel electrophoresis where two distinct bands of 28S and 18S were detected.

2.6.2 cDNA synthesis cDNA synthesis was carried out using RT2 First Strand kit (Qiagen, USA) based on the manufacturer’s protocol. Briefly, genomic DNA removal reaction components were mixed with the RNA template (850 ng) and incubated for 5 min at 42°C. Reverse-transcription mix was added into the resulting solution, and the reaction was carried out in a Veriti Thermal Cycler 15



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(Applied Biosystems, USA) at 42°C for exactly 15 min and at 95°C for 5 min to facilitate the annealing, reverse-transcription, and inactivation of the reaction.

2.6.3 Gene expression by real-time quantitative polymerase chain reaction (RT-qPCR) cDNA and nuclease-free water were added to the RT2 SYBR Green ROX FAST Mastermix (Qiagen) as per the manufacturer’s instructions. Each well of the RT² Profiler™ PCR Array Human Wound Healing ring (PAHS-121Z) was filled with 20 µL of the resulting mixture using the QIAgility (Qiagen). The ring was sealed using the Rotor-Disc Heat-Sealing Film before RTqPCR was performed on the Rotor Gene Q 100 (Qiagen) with a hold time of 10 min at 95°C, followed by 40 cycles of 15 s at 95°C and 30 s at 60°C. The threshold cycle (CT) value of 84 genes related to wound healing, 5 housekeeping genes, 1 genomic DNA control (GDC), 3 reverse transcription control (RTC), and 3 positive PCR control (PPC) was determined. Fold changes (2−∆∆CT) between the control group (HDFs on cell culture plate) and experimental group (HDFs on hydrogel) after 24 h were analyzed using Qiagen’s PCR array data analysis web portal, where all the housekeeping genes (β-actin; β-2-microglobulin; glyceraldehyde-3phosphate dehydrogenase; hypoxanthine phosphoribosyltransferase 1; and ribosomal protein, large, P0) were used for normalization of the sample data. CT value > 33 was considered a negative call, and if the CT value of the genomic DNA control was > 30, no genomic DNA was evident19.

2.7 Statistical analysis For multiple group comparison with only one variable, statistical analysis was performed by using one-way ANOVA followed by Tukey’s post-hoc test, whereas, for two variables, two-way 16


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ANOVA was performed. For comparison between two groups, Student t-test was carried out by using Graph Pad Prism version 6.0. P-values < 0.05 were considered statistically significant. All characterization data values were obtained from three different hydrogel batches (n = 3), while all in vitro data values, in the cellular and molecular interaction studies, were obtained from three biological replicates (n = 3). All values are expressed as the mean ± standard deviations (SD).

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3. RESULTS AND DISCUSSION: 3.1 Diameter of BC fiber

Figure 2: Field emission scanning electron microscope (FESEM) image of bacterial cellulose (BC) fiber and its diameter.

Cellulose fibers can be considered nanosize if their diameter is less than 100 nm. This size can be obtained using high pressure homogenization, grinding, or refining20. In this study, nanosize cellulose fibers with a diameter less than 100 nm (Figure 2) were successfully produced using high shear homogenization, thus, they are known as BNC21.

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3.2 Characterization of hydrogel 3.2.1 Swelling, hardness, adhesive force, porosity, and WVTR

Figure 3: Physical evaluation of different hydrogel formulations. A) Swelling ratio; B) Hardness of wet hydrogels; C) Adhesive force of wet and dry hydrogels; D) Porosity; E) Water vapor transmission rate (WVTR) of different hydrogel formulations compared to control (uncovered 19



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surface). # indicates significant difference (p < 0.05) compared to 208045, § significant difference compared to 307060, * significant difference compared to all.

Characterization of the hydrogels was conducted in the form in which the hydrogels were intended to be used, i.e., neutralized and autoclaved. At 24 h, the 208045, 208060, 307045, and 307060 hydrogels achieved a swelling ratio of 2,495 ± 202%, 2,154 ± 195%, 2,337 ± 248%, and 2,016 ± 195%, respectively (Figure 3A). When the pH is higher than the pKa (swelling test performed at pH 7.4), the carboxylate groups are ionized, and the electrostatic repulsion between COO- groups causes the hydrogel to swell4. EB irradiation dose (45 kGy versus 60 kGy) significantly (p < 0.05) affected the swelling of the hydrogel. The swelling of hydrogels with higher irradiation dose (208060 and 307060) was significantly lower compared to that of its counterparts (208045 and 307045) because a larger number of reactive species and sites were produced with increased irradiation dose, increasing the degree of crosslinking4. On the other hand, the amount of AA (20% AA versus 30% AA) did not significantly (p > 0.05) influence the swelling ratio, even though there was more AA available to be cross-linked during the reaction. As the hydrogel acts as a wound dressing and a cell carrier, its water permeability is essential for the absorption of wound fluid and exudates and for the transfer of nutrients and metabolites via the hydrogel22. It was reported that the ability of a scaffold to absorb fluid over 80 times its initial weight was adequate for skin tissue engineering23. Considering that the hydrogel would be used in its wet form, hardness by compression was tested on swelled hydrogels and was significantly different across all formulations (Figure 3B), with 208045 having the least mechanical strength (8.01 ± 1.52 N), while 307060 being the hardest (16.38 ± 1.78 N). The higher the degree of cross-linking, which was achieved with increased 20


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irradiation dose and greater amount of AA, the higher was the hardness24,25. From our observation, 208045 hydrogel is quite fragile, difficult to handle, and may not be able to withstand mechanical abrasion. In contrast, 208060 and 307060 hydrogels are rather tough and may not be as flexible in conforming to any contour of the wound. Strong adhesiveness to the wound may result in secondary injury and pain during removal, therefore, bioadhesiveness is also an important property of a wound dressing. The inside of the rat skin was used to represent the ECM of the wound. Adhesive force was negligible when the hydrogel was dry. With increased irradiation dose, more AA was grafted on the hydrogel. Therefore, its adhesiveness increased with a higher amount of AA. Acrylics are common skin adhesives26. A comparison among the four formulations (Figure 3C) revealed that the 307045 hydrogel is mildly adhesive (0.69 ± 0.25 N) in its wet form, which makes it suitable as a wound dressing because it can adhere to the wound, but does not hurt the wound upon removal. Besides, once the hydrogel dries up, it is no longer sticky. The investigation into 56 commercial wound dressings on human skin found that the median values for adhesiveness, classified according to different groups were 2.25 N (hydrocolloid), 1.14 N (acrylate), 0.9 N (polyurethane), and 0.7 N (silicone) and the subjective pain intensity correlates with the adhesiveness27. The adhesive force value of the 307045 hydrogel is comparable to that of silicone (the group with the lowest adhesive force)27 and other wound dressing materials that are studied28,29. Porosity (Figure 3D), which is the percentage of void volume in a material, was the highest in the 208045 hydrogel (54.30 ± 5.26%) and the lowest in the 307060 hydrogel (21.40 ± 1.23%). All the four hydrogel formulations were not as porous as some other biomaterials investigated in literature that could reach 80–90% porosity30,31. Mechanical properties of the hydrogel are 21



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interrelated with porosity. The more porous a material is, the less mechanical strength it possesses to confer structural stability to the material32. Results from our study (Figure 3B and 3D) were in line with this observation. Porous 3D scaffolds are generally used in tissue engineering applications. The porosity, together with pore size, influences cell behavior and determines the final mechanical property of the scaffold32. There was a significant difference in water loss from an uncovered surface (control) compared to that covered with BNC/AA hydrogel (Figure 3E). The degree of cross-linking was not large enough to affect the WVTR as there was no significant difference (p > 0.05) between all four formulations (1,388 ± 99, 1,325 ± 177, 1,322 ± 128, and 1,286 ± 87 g/m2/day for 208045 to 307060, respectively). Wound dressings should be able to control water loss from a wound bed to provide a moist environment for healing. A very high WVTR can lead to the dehydration of a wound, whereas a very low WVTR can cause the accumulation of wound exudates33. It has been reported that human skin transpires water vapor at a rate ranging between 240 and 1,920 g/m2/day, whereas the WVTR of uncovered wounds could be in the order of 4,800 g/m2/day34. The WVTR of the hydrogel is in an acceptable range as it is within the WVTR range of the human skin.

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3.2.2 Pore size, contact angle, and topography

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Figure 4: Microscopic analysis of different hydrogel formulations. A) Cross-section scanning electron micrograph of hydrogel pores at 150 magnification and 13kV; B) Pore size distribution; C) Contact angle of wet (dried overnight) and freeze-dried hydrogels; D) Threedimensional (3D) topography image. # indicates significant difference (p < 0.05) compared to 208045, † significant difference compared to 208060, Ø significant difference compared to 307045, § significant difference compared to 307060.

As shown in Figure 4A, the pores of the hydrogel were irregular, with some closed pores. In addition, the 307060 hydrogel exhibited thicker walls compared to the other formulations. The number of pores with different pore size ranges for all the hydrogels was statistically nonsignificant (Figure 4B), with the majority of the pores between 50-199 µm in size. The optimal pore size is dependent on the cell type that will be finally used for the biomaterial application. In the case of fibroblasts for skin regeneration, it has been reported that a size range of 20–125 µm was desirable35. Previous work focusing on the topological organization of the BC/AA hydrogel discovered that large pores were not directly connected to each other, but were connected via a series of interconnected small pores, making the travelling path in the matrix complex and lengthy36. Interconnecting pores of suitable size are needed to facilitate cell growth, migration, and flow of nutrients. If pores are too small, cell migration is restricted, distribution of nutrients is limited, and removal of waste is reduced. Conversely, too large pores lack specific surface area available for cell attachment. Hence, maintaining the optimal pore size for cell migration and specific surface area for cell attachment is crucial37.

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As fibroblasts are anchorage-dependent, the scaffold surface properties should facilitate their attachment35. Hydrophilic surfaces favor cell adhesion and growth. Hydrophilicity can be assessed by measuring the contact angle via wettability of water on the material’s surface. The lower the contact angle, the more hydrophilic is the surface38. Hydrogel in its dry form exhibits different contact angles (Figure 4C). The contact angles were lower in hydrogels formulated with higher irradiation dose. This occurred because more AA was grafted onto the hydrogel, resulting in more carboxylic groups, thereby increasing the surface hydrophilicity39. However, in its relatively wet form, which is the form when cells were seeded on the hydrogel, there was no difference in contact angle between formulations, and the contact angle was in the 50°–60° range. The results obtained by Kim et al.40 indicated that 50°–60° was the optimal water contact angle for fibroblast cell adhesion and growth. Currently, there is a lack of general consensus on how surface roughness affects fibroblast behavior. Ribeiro et al.41 discovered that the material with the highest roughness demonstrated the maximum proliferation of fibroblasts; Bourkoula et al.42 observed that there is a roughness threshold where HDFs and mouse 3T3 fibroblast adhesion, proliferation, and morphology were significantly altered, while Cao et al.43 reported that the attachment and growth of human gingival fibroblasts increased as surface roughness decreased. However, these conflicting findings were not of utmost importance in our study because the surface roughness (Figure 4D) of the all hydrogels with different formulations appeared to be the same with grooves and ridges. As observed, the quantitative roughness value [arithmetic mean height (Sa)] differed depending on the measurement spot and ranged between 0.15 and 0.50 µm.

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3.2.3 BET As the characterizations for swelling, hardness and bioadhesiveness, which were significantly different between formulations, seemed to suggest that the 307045 hydrogel might be the most suitable formulation as a cell carrier and a wound dressing, the BET and most of the in vitro studies were only performed on the 307045 hydrogel. The BET result obtained upon testing the extracted solution from the 307045 hydrogel was < 0.05 EU/mL, much lower than the established USP 39 limit for medical device, i.e., 20 EU/device. Generally, BNC shows negligible amount of endotoxin and has been approved by the Food and Drug Administration (FDA) for use as MTA™ Surgical Sheet, Xylos® Vessel Guard, Xylos® Porous Surgical Mesh and Securian™ Tissue Reinforcement Matrix44. However, since BNC is a biomaterial produced by Gram-negative bacteria and the hydrogel may have potential for systemic exposure in a full-thickness wound, the presence of endotoxins was tested45.

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3.3 Cellular interaction 3.3.1 Cell attachment and proliferation

Figure 5: A) Cell attachment of human dermal fibroblasts (HDFs) on different hydrogel formulations at various time points; B) Cell proliferation of HDFs, quantified by using AlamarBlue® Assay, on different hydrogel formulations compared to cell culture plate on days 1 and 7. The asterisk (*) indicates significant difference (p < 0.05).

Properties of the biomaterial have long been recognized to affect the behavior of the cells at its cellular (attachment, proliferation, viability, morphology, migration, and transfer) and molecular (gene expression) level2. There was no statistical difference in terms of cell attachment (Figure 5A) and proliferation (Figure 5B) between the four hydrogel formulations at every time point. This is not surprising as the characterization of properties known to affect cell attachment and proliferation, such as hydrophilicity (Figure 4C), surface roughness (Figure 4D), and pore size (Figure 4B) were similar between all formulations. As the first step in a series of cellular events, cell attachment is important for cell function, which includes cell spreading, migration, and proliferation46. Cell attachment increased with time, and at 4 h, 82.91 ± 4.27%, 81.48 ± 2.47%, 84.33 ± 2.47% and 81.48 ± 6.53% of HDFs were attached 27



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on the 208045, 208060, 307045, and 307060 hydrogels, respectively (Figure 5A). Sutthikhum et al.47 showed that approximately 60% of fibroblasts attached on their silk fibroin scaffold after 24 h. This suggests that the BNC/AA hydrogel exhibits rapid cell attachment as more than 80% of the cells are already attached within 4 h. The excellent attachment may be due to the hydrophilicity of the hydrogel (Figure 4C) and reasonably large surface area available for the cells to attach as its porosity is rather low (Figure 3D). Therefore, cell transfer to the wound site can be carried out after a duration as short as 4 h; in view that it would be beneficial to deliver the treatment to the patient as soon as possible after an injury to the skin. Metabolic activity of cells causes the redox indicator in the AlamarBlue assay to change from an oxidized (non-fluorescent, blue) form to a reduced (fluorescent, red) form. Hence, the higher the percentage reduction of AlamarBlue is, the higher is the number of viable cells. The percentage reduction of HDFs on the 208045, 208060, 307045, and 307060 hydrogel formulations (35.74 ± 3.42%, 35.54 ± 7.15%, 36.31 ± 3.49% and 37.82 ± 6.33% respectively) was comparable with that of the control (46.74 ± 10.81%) on day 1 (Figure 5B). The lower non-significant percentage reduction on the hydrogel may be due to some cells flowing off the scaffold into the cell culture plate during seeding. Even though there was a slight increase in the percentage reduction on the hydrogel on day 7 compared to day 1, it was not statistically significant, especially when compared to the control at day 7. The results revealed that the hydrogel was capable of maintaining cell viability throughout 7 days, but HDFs did not proliferate on the hydrogel. In order to improve the biocompatibility of our previous preliminary BC/AA hydrogel formulation16, less AA was used to fabricate the hydrogels, but the environment was still not conducive to support cell proliferation. Possible reasons for this may be attributed to the low porosity of the hydrogels (Figure 3D) since scaffold with 80–90% porosity did facilitate 28


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fibroblast proliferation30,31 or non-interconnecting pore formations, which limited nutrient distribution and diffusion of waste48. 3.3.2 Cell viability and morphology

Figure 6: A) Live (green) and dead (red) assay of human dermal fibroblasts (HDFs) on the 307045 hydrogel formulation and cell culture plate on days 1 and 3 at 200 magnification; B) Scanning electron micrograph of HDFs on the 307045 hydrogel formulation on days 1, 3, and 7 at 1,000 magnification and 15kV.

As there was no difference in cell attachment and proliferation between formulations, and the characterization results (in line with the function of the hydrogel as a cell carrier and wound 29



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dressing) favored the 307045 hydrogel formulation, we further investigated cell-hydrogel interactions using the 307045 hydrogel. The results of the cell viability assay (Figure 6A) were consistent with the cell proliferation findings (Figure 5B). The image of the cells on the hydrogel showed live (stained green) HDFs and only a small number of dead (stained red) HDFs, which were comparable to those on the cell culture plate (control). The only difference was that on day 3, the quantity of HDFs on the plate increased, while the amount of HDFs on the hydrogel was similar to that on day 1. This confirms that the hydrogel supports cell viability but not cell proliferation. Considering that the intended use of the hydrogel is to serve as a cell carrier, with most of the cells being transferred to the wound site within several days, the ability to only maintain cell viability is deemed sufficient for its function. Compared to our preliminary study16, where HDFs were round-shaped, HDFs seeded on the 307045 hydrogel spread out in spindle-shape, maintaining their morphology throughout the 7 days (Figure 6B). This outcome was one of the improvements that we hoped to achieve when we fabricated the hydrogel with a lower amount of AA. This indicated that HDFs responded better on the 307045 formulation and exhibited good cellular attachment with the presence of lamellipodia49. When observed under SEM, HDFs were actually on the surface of the hydrogel and not within its matrix.

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3.3.3 Cell migration and transfer

Figure 7: A) Three-dimensional (3D) confocal image of cell migration on days 1, 3, and 7; B) Number of cells transferred from the 307045 bacterial nanocellulose/acrylic acid (BNC/AA) hydrogel to ovine collagen hydrogel (OCH) on days 1, 2, and 3; C) Number of live and dead cells on OCH on day 3 for both submerge and airlift model.

In accordance with the SEM observation, HDFs did not migrate into the hydrogel as the cells (cytoplasm stained in green and nucleus stained in blue) remained on the top surface of the hydrogel on days 0, 3, and 7 (Figure 7A). Although the suitable pore size for fibroblast was reported to be 20–125 µm35 and the majority of the pores lied in that range (Figure 4B), Barud et al.50 also discovered that even with an average pore size of 102 ± 5.43 µm on the surface of their BC/50% silk fibroin scaffold, no migration of fibroblasts into the scaffold was noticeable. The authors suggested that this occurred because the dense BC network does not possess large 31



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enough pore size to facilitate cell migration. This may be one of the factors preventing migration of cells into the BNC/AA hydrogel or perhaps, close pores may play a role as well (Figure 4A). Nevertheless, in order for this hydrogel to act as a cell carrier, limited or no cell migration is favorable as the cells are required to remain on the hydrogel surface so that it will quickly be transferred to the wound bed. Two models were investigated for the cell transfer study, namely, the submerge model, where the medium was in direct contact with both hydrogels, and the airlift model, using a cell culture insert, where only the OCH was directly in contact with the medium (Figure 1). The airlift model represents the actual situation more closely as the other side of the hydrogel, which does not contain cells, is not in direct contact with wound fluids. OCH, made up primarily of collagen type I (COL-I)18, was employed to mimic the ECM of the wound in vitro since COL-I is the main constituent of the ECM51. For the airlift model (Figure 7B), 60.52 ± 1.65% of HDFs was transferred from the BNC/AA hydrogel to the OCH on day 1, with the most amount transferred on the first day, followed by day 2 (16.33 ± 2.68%) and the least on day 3 (4.40 ± 1.21%). The total number of cells transferred in 3 days reached approximately 80%. A similar trend was observed for the submerge model as well (Figure 7B). This result showed an improvement compared to that in our preliminary study16 with a 10% increase of total HDFs being transferred, so that more cells would be available to accelerate the wound healing. The ability of the cells to be transferred from the BNC/AA hydrogel to the OCH was probably because of the dryness of the BNC/AA hydrogel compared to the OCH. This also applied to the submerge model as the medium was

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only added up to the same level of the top surface of the BNC/AA hydrogel. This could be the reason why the cell transfer results between the two models did not differ. Though the majority of HDFs were transferred to the OCH, it was important to investigate whether these cells were alive or dead. The cell viability assay revealed that on day 3 (Figure 7C), 87.23 ± 4.46% were live HDFs in the submerge model, while 84.79 ± 9.73% were live HDFs in the airlift model, quantified on the OCH. The excellent cell transfer, with predominantly live HDFs, suggests the potential of BNC/AA hydrogel to perform as a cell carrier.

3.4 Molecular interaction

Figure 8: Significant fold changes of gene expression in human dermal fibroblasts (HDFs) on the 307045 hydrogel relative to HDFs on cell culture plate (control) at 24 h, with its category. 33



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The threshold for the fold change in gene expression was set at 2 and the significant difference in gene expression was calculated with the p-value < 0.05. IL6, Interleukin 6 (interferon, beta 2); IL10, Interleukin 10; F3, Coagulation factor III (thromboplastin, tissue factor); MMP2, Matrix metallopeptidase 2 (gelatinase A, 72kDa gelatinase, 72kDa type IV collagenase); CTSK, Cathepsin K; FGF7, Fibroblast growth factor 7; GM-CSF, granulocyte-macrophage colony stimulating factor; TGFB1, Transforming growth factor, beta 1; COX2, Cyclooxygenase 2 (Prostaglandin-endoperoxide synthase 2 and prostaglandin G/H synthase). Since most cells were transferred from the BNC/AA hydrogel to the OCH on day 1, gene expression was measured in order to investigate which wound healing genes were upregulated or downregulated after being in contact with the hydrogel for one day. All 84 genes of interest related to wound healing (from Qiagen’s standard catalogue wound healing panel) with their fold changes and p-values were divided into several categories according to their role (Table S1). For the WNT signaling, cytoskeleton regulators, kinases, cell adhesion molecules, cell surface receptors, and ECM structural constituents category, there was either no significant difference (p-value > 0.05) between the control (HDFs on cell culture plate) and the hydrogel, and/or the change in gene expression was less than 2-fold. The BNC/AA hydrogel was not an inert biomaterial because even in the absence of injury, there were 8 genes that were upregulated and one gene that was downregulated compared to the control (Figure 8). For the inflammatory cytokines and chemokines, pro-inflammatory interleukin 6 (IL6) and antiinflammatory interleukin 10 (IL10) were upregulated by 3.57 and 8.75-fold, respectively. Even though prolonged inflammation impedes healing, it is a necessary occurrence to stimulate progression in the wound healing pathway52. A study comparing skin excision in wild-type and knockout (IL6 deficient) mice showed that the knockout mice exhibited delayed healing, leading 34


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to the conclusion that IL6 was important in the wound healing process, possibly by regulating leukocyte infiltration, angiogenesis, and collagen accumulation53. Furthermore, overexpression of IL10 is thought to play a part in decreasing the formation of scar in the same way as during fetal tissue regeneration54. Coagulation factor III (F3) or more commonly called tissue factor, is one of the initiator of clot formation. Although F3 is essential for coagulation, its expression is regulated differently at different sites. F3 was found to be downregulated around new blood vessels to enable the formation of endothelial sprouts and protect the newly generated leaky vessels from thrombosis55. Therefore, downregulation of F3 may also be beneficial in promoting revascularization. The ECM is always in a dynamic state where its components are deposited or degraded. The enzymes involved in the ECM remodeling include the matrix metallopeptidase (MMP) and papain-cysteine proteinases family56. Matrix metallopeptidase 2 (MMP2), also known as gelatinase A or type IV collagenase, is mainly responsible for the breakdown of gelatin and collagen IV. Besides modifying the wound matrix, MMP2 can also assist in the migration of cells by allowing infiltration of neutrophils and lymphocytes, or releasing chemoattractants when the basement membrane is degraded57. A 6.12-fold increase of cathepsin K (CTSK) could be advantageous as the findings by Runger et al.56 suggested that CTSK is an antifibrotic agent that is required to counteract the synthesis of new ECM proteins during scar formation so that the ECM returns to its normal state. Among the many growth factors regulating wound healing and repair and tissue regeneration, fibroblast growth factor 7 (FGF7), granulocyte-macrophage colony stimulating factor (GM-CSF) 35



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and transforming growth factor, beta 1 (TGFB1) were upregulated in the hydrogel group. FGF7, also termed as keratinocyte growth factor, is secreted by fibroblasts but acts in a paracrine manner through FGFR2IIIb receptor expressed only on keratinocytes. FGF7 is involved in reepithelialization by stimulating the growth and migration of keratinocytes, as well as in neovascularization58. Overexpression of GM-CSF also enhances re-epithelialization and neovascularization, in addition to increasing the production of pro-inflammatory cytokines59. TGFB1 was significantly the most upregulated (13.06-fold) among all the genes tested. TGFB1 signaling activates the intracellular SMAD pathway and has a wide spectrum of functions encompassing all stages of wound healing.

In the inflammatory phase, TGFB1 recruits

neutrophils and macrophages, while in the proliferation phase, TGFB1 is associated with angiogenesis, re-epithelialization, and formation of granulation tissue. In the remodeling phase, TGFB1 is involved in collagen production, inhibits the synthesis of MMPs, and promotes the generation of metalloproteinase inhibitor (TIMP1), all of it contributing to the accumulation of collagen. Therefore, if an excess of TGFB1 is present in the later part of the wound healing process, it may lead to over-healing, resulting in hypertrophic scarring and keloid58. Hence, the up-regulation of TGFB1 in HDFs on the hydrogel, to be used as part of the dressing in the early phase of wound healing and not later is ideal. Cyclooxygenase 2 (COX2) enzymes are induced upon injury and they convert arachidonic acid to prostaglandin E2 (PGE2), a primary mediator of inflammation and wound healing59. Previous studies60,61 showed that blocking COX2 with a COX2 inhibitor led to delayed re-epithelialization and angiogenesis, especially in the early stage of wound healing. Additionally, Fairweather et al.60 reported that blocking COX2 also inhibited the ECM production with two times less amount 36


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of alpha smooth muscle actin (αSMA) protein in the granulation tissue of COX2 inhibitor-treated mice. Considering that wound healing is a complex process, with each gene involved in multiple interrelated roles, it remains to be determined through in vivo studies whether the upregulation and downregulation of these mRNA is transient or lengthy and ultimately, whether it can truly help in the acceleration of wound healing with less scarring and minimal contraction.

4. CONCLUSION The characterization results of swelling, hardness, adhesive force, porosity, WVTR, surface roughness, pore size distribution, hydrophilicity, and bacterial endotoxin, together with the cellular and molecular interaction data, demonstrates that the 307045 BNC/AA hydrogel is a promising wound dressing and HDF cell carrier to expedite wound healing. At the cellular level, the hydrogel showed rapid cell attachment with more than 80% at 4 h, maintained the cell viability and morphology of HDFs with limited migration, and facilitated transfer of approximately 80% of HDFs comprising of mostly live cells by day 3. At the molecular level, the hydrogel upregulated eight important wound healing genes (IL6, IL10, MMP2, CTSK, FGF7, GM-CSF, TGFB1 and COX2) and downregulated one gene (F3) which can be beneficial during the wound healing process.

CONFLICTS OF INTEREST There are no conflicts of interest to declare. 37



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ACKNOWLEDGEMENT We thank Malaysian Nuclear Agency, Electron Microscope Unit of UKM, Centre for Research and Instrumentation of UKM, Cell Therapy Centre of UKMMC, Centre of Quality Control, NPRA and ALS Technichem (M) Sdn Bhd for the support given.

FUNDING SOURCES This work was supported by funding from the Ministry of Higher Education of Malaysia (FRGS/1/2017/SKK09/UKM/01/1).

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Table of contents graphic 320x150mm (300 x 300 DPI)



Figure 1: Pictorial representation of the submerge and airlift models used for the cell transfer study. 240x102mm (300 x 300 DPI)


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Figure 2: Field emission scanning electron microscope (FESEM) image of bacterial cellulose (BC) fiber and its diameter. 50x50mm (300 x 300 DPI)



Figure 3: Physical evaluation of different hydrogel formulations. A) Swelling ratio; B) Hardness of wet hydrogels; C) Adhesive force of wet and dry hydrogels; D) Porosity; E) Water vapor transmission rate (WVTR) of different hydrogel formulations compared to control (uncovered surface). # indicates significant difference (p < 0.05) compared to 208045, § significant difference compared to 307060, * significant difference compared to all. 299x329mm (300 x 300 DPI)


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Figure 4: Microscopic analysis of different hydrogel formulations. A) Cross-section scanning electron micrograph of hydrogel pores at 150× magnification and 13kV; B) Pore size distribution; C) Contact angle of wet (dried overnight) and freeze-dried hydrogels; D) Three-dimensional (3D) topography image. # indicates significant difference (p < 0.05) compared to 208045, † significant difference compared to 208060, Ø significant difference compared to 307045, § significant difference compared to 307060. 279x399mm (300 x 300 DPI)



Figure 5: A) Cell attachment of human dermal fibroblasts (HDFs) on different hydrogel formulations at various time points; B) Cell proliferation of HDFs, quantified by using AlamarBlue® Assay, on different hydrogel formulations compared to cell culture plate on days 1 and 7. The asterisk (*) indicates significant difference (p < 0.05). 279x99mm (300 x 300 DPI)


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Figure 6: A) Live (green) and dead (red) assay of human dermal fibroblasts (HDFs) on the 307045 hydrogel formulation and cell culture plate on days 1 and 3 at 200× magnification; B) Scanning electron micrograph of HDFs on the 307045 hydrogel formulation on days 1, 3, and 7 at 1,000× magnification and 15kV. 338x250mm (300 x 300 DPI)



Figure 7: A) Three-dimensional (3D) confocal image of cell migration on days 1, 3, and 7; B) Number of cells transferred from the 307045 bacterial nanocellulose/acrylic acid (BNC/AA) hydrogel to ovine collagen hydrogel (OCH) on days 1, 2, and 3; C) Number of live and dead cells on OCH on day 3 for both submerge and airlift model. 338x199mm (300 x 300 DPI)


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Figure 8: Significant fold changes of gene expression in human dermal fibroblasts (HDFs) on the 307045 hydrogel relative to HDFs on cell culture plate (control) at 24 h, with its category. The threshold for the fold change in gene expression was set at 2 and the significant difference in gene expression was calculated with the p-value < 0.05. IL6, Interleukin 6 (interferon, beta 2); IL10, Interleukin 10; F3, Coagulation factor III (thromboplastin, tissue factor); MMP2, Matrix metallopeptidase 2 (gelatinase A, 72kDa gelatinase, 72kDa type IV collagenase); CTSK, Cathepsin K; FGF7, Fibroblast growth factor 7; GM-CSF, granulocytemacrophage colony stimulating factor; TGFB1, Transforming growth factor, beta 1; COX2, Cyclooxygenase 2 (Prostaglandin-endoperoxide synthase 2 and prostaglandin G/H synthase). 184x150mm (300 x 300 DPI)


Cellular and Molecular Interaction of Human Dermal Fibroblasts with

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