In vitro and in vivo evaluation of a bioactive polysaccharide based self-healing hydrogel film for enhanced wound healing

Abstract

The objective of this research was to develop a carboxymethyl chitosan (CMCS) loaded borax cross-linked guar gum (GG) self-healing hydrogel film dressing for enhanced wound healing. Cross-links between borax and GG were confirmed through FTIR, SEM, and zeta potential measurement. The optimized formulation demonstrated favourable characteristics, including a thickness of 0.25 ± 0.008 mm, a water vapour transmission rate of 592.9 ± 55.1 g m^-2 day^-1, tensile strength of 37 ± 1 MPa in the wet state and 61.33 ± 4.16 MPa in the dry state, a swelling ratio of 557.84%, and 2, 2-Diphenyl-1-picrylhydrazyl (DPPH) radical scavenging activity of 53.132 ± 1.02%. Furthermore, the optimized formulation exhibited a healing efficiency of 43.949 ± 0.56%, as determined by tensile strength measurements of the self-healed hydrogel film dressings after 6 h. Antimicrobial testing revealed broad-spectrum activity against wound-associated pathogens, including Staphylococcus aureus and Escherichia coli. The therapeutic efficacy was evaluated through in vivo experiments using a rat full-thickness wound model, which demonstrated that the optimized formulation achieved greater wound closure than commercial Tegaderm and ofloxacin loaded dressings. The CMCS-loaded hydrogel film dressing showed significantly higher hydroxyproline levels than the commercial Tegaderm and ofloxacin dressings. Hydroxyproline levels increased from day 7 and day 14, indicating progressive collagen synthesis. Histopathology analysis revealed that the CMCS-loaded hydrogel film dressing promoted faster re-epithelialization, reduced inflammatory cell infiltration, and enhanced granulation tissue formation compared to Tegaderm and ofloxacin dressings.

Keywords

guar gum carboxymethyl chitosan borax self-heal enhanced wound healing

Background

Wound healing is a multifaceted biological process that replaces devitalized and missing cellular and tissue components.¹ Effective wound healing requires a moist environment; however, conventional dressings generally fail to provide sustained moisture retention. Therefore, advanced wound dressings are essential to maintain moisture, protect the wound site, and promote rapid tissue regeneration.² Among these, hydrogel dressings have emerged as promising materials because of their high water content, excellent biocompatibility, and ability to maintain a hydrated environment that facilitates the wound healing process.^3–5

However, traditional hydrogels are prone to movement-induced damage, leading to decreased performance and necessitate frequent replacement.⁶ Furthermore, structural damage may expose the wound to external contaminants, thereby increasing the risk of infection. Self-healing hydrogels have received significant attention because of their ability to autonomously repair damage, thereby enhancing wound protection and reducing the need for repeated replacement.⁷ These hydrogels can restore damaged regions, maintain an effective protective barrier, and reduce the risk of infection. Moreover, their self-healing capability reduces the frequency of dressing changes, improves patient comfort, and protects the wound from disruption.

GG, a galactomannan polysaccharide, has been widely utilized in tissue engineering owing to its biodegradability, biocompatibility, and non-toxic nature.^8,9 Its high water retention capacity creates an optimal moist environment that supports collagen deposition, granulation tissue development, and re-epithelialization.¹⁰ However, GG does not inherently possess self-healing properties, but the introduction of borax as a reversible cross-linking agent confers self-healing capability. Borax is an inexpensive compound with a reported lethal oral dose of approximately 15-20 g in adults.¹¹ It acts as a cross-linking agent, forming self-healing hydrogels via reversible borate-diol interactions.¹² Specifically, Borax forms reversible borate ester bonds with GG’s diol groups, enabling the hydrogel to restore itself after damage.

CMCS is a water-soluble, biocompatible, biodegradable, and non-toxic material suitable for wound healing.¹³ CMCS’s amino groups interact with negatively charged bacterial membranes, leading to bacterial eradication at the wound site.¹⁴ Moreover, CMCS mediates chemotaxis of neutrophils and macrophages, thereby aiding in the inhibition of wound infections.¹⁵ Transforming growth factor beta-1 (TGF-β1), alpha-smooth muscle actin (α-SMA), and matrix metalloproteinase-1 (MMP-1) protein expression play significant roles in the healing of wounds. α-SMA stimulates fibroblast-to-myofibroblast differentiation, which secretes collagen matrix crucial for the healing of wounds.¹⁶ MMP-1 breaks down collagen, which facilitates cell migration during the early phase and degrades excessive collagen accumulation in the later phase of wound healing.¹⁷ TGF-β1 helps promote the proliferation and differentiation of fibroblasts. CMCS promoted the expression of α-SMA, TGF-β1, and MMP-1, thereby enhancing wound healing.¹⁸

We developed a self-healing hydrogel film dressing composed of GG cross-linked with borax and loaded with CMCS. The developed system exists as a flexible hydrogel film dressing formed after drying the borax cross-linked hydrogel matrix. The hydrogel film dressing retains the characteristic swelling, moisture retaining, and self-healing properties of hydrogels while providing sufficient mechanical integrity for handling and direct wound application. The potential of this multifunctional hydrogel to accelerate wound healing while preventing infection and promoting tissue regeneration is systemically explored through in vitro characterization and in vivo biological assessment.

Methods

Materials

GG was procured from SD Fine Chem Ltd., India. Borax, glycerol, and ethanol were procured from Loba Chemie Pvt. Ltd., India. DPPH was procured from SRL, India. CMCS was obtained from Yarrow Chem, India. Ofloxacin was kindly gifted by Crest Life Sciences Pvt. Ltd, Himachal Pradesh, India. All other reagents of analytical grade were used.

Preparation of a hydrogel film dressing

Specific amounts of GG were incorporated into water and heated at 60 °C for 4 h (Table 1). The mixture was then allowed to swell overnight. Then glycerol was added (25% w/w of GG) as a plasticizing agent.¹⁹ A fixed amount of borax as a cross-linker was separately dissolved in water and incorporated into the above solution (Table 1). The resulting solutions were poured into Petri dishes (ø = 7.7 cm) and dried at 60 °C for 30 h, then left to air-dry till a constant weight was achieved. For the preparation of the CMCS-loaded hydrogel film dressing and ofloxacin loaded hydrogel film dressing, varying concentrations of CMCS and ofloxacin were added after the addition of glycerol, followed by the same procedure. After drying, the obtained material formed a flexible self-healing hydrogel film dressing suitable for direct application as a wound dressing scaffold.

Table 1.

Composition and thickness of hydrogel film dressing.

FormulationCode	Amount of GG (% w/v)	Amount of Borax (wt% of GG)	Amount of CMCS (wt% of GG)	Amount of Ofloxacin (wt% of GG)	Water (mL)	Thickness (mm)Mean±SD, n = 3
F1	0.5	-	-	-	40	0.11 ± 0.008
F2	0.5	1	-	-	40	0.13 ± 0.004
F3	1.5	1	-	-	40	0.17 ± 0.004
F4	2	1	-	-	40	0.21 ± 0.008
F5	2	1	9	-	40	0.25 ± 0.004
F6	2	1	12	-	40	0.25 ± 0.004
F7	2	1	15	-	40	0.25 ± 0.008
F8	2	1	-	1	40	0.24 ± 0.004

Fourier transform infrared (FTIR) spectroscopy

The FTIR spectra of GG and borax cross-linked GG-based hydrogel film dressing were analyzed using an FTIR spectrophotometer (Alpha-E, Bruker, USA) at 4000-400 cm^-1.^20,21 The powdered samples were subjected to mix with potassium bromide. The prepared mixture was subsequently formed into pellets through a hydraulic press for FTIR measurements.

Scanning electron microscopy (SEM)

A scanning electron microscope (SU3800, Hitachi, Japan) was used to evaluate the structural morphology of the hydrogel film dressing.^22,23 The dried hydrogel film dressings were mounted on conductive adhesive tapes using tweezers and sputter coated with gold under vacuum. SEM micrographs were obtained at an acceleration voltage of 20 kV to characterize the microstructure of the film samples.

Thickness of hydrogel film dressing

Hydrogel film dressing thickness was determined using a digital vernier calliper, and the average value was calculated.²⁴

Water vapour transmission rate

A hydrogel film dressing sample (1.5 cm diameter) was mounted on a water-filled tube using Teflon to evaluate the water vapour transmission rate (WVTR).²⁵ The tube was placed in a desiccator maintained at 37 °C with 75% relative humidity (RH), and the rate of mass change was measured. A mixture of 10.23 g of water and 20 g of NaCl was placed at the base of the desiccator to maintain the humidity.²⁶ After 24 h, the test tube was removed, and the weight loss was measured. The WVTR of the hydrogel film dressing was determined using the equation below.²⁷

W V T R (g . m^{- 2} . {d a y}^{- 1}) = (\frac{Δ m}{Δ t}) / A

(1)

Δm/Δt = Loss of water for 24 h (g/day), A = effective transfer area (m²).

Mechanical strength

The tensile properties of the hydrogel film dressings were assessed in both wet and dry conditions.²⁸ In wet states, hydrogel film dressings were soaked in phosphate buffer saline for 30 s, and then any extra water was eliminated using filter paper. Subsequently, tensile stress and % elongation at break were measured. The tensile stress of the hydrogel film dressing, defined as the stress at which the specimen breaks, was determined by dividing the applied force (N) by the cross-sectional area (mm²) (Equation (2)). The % elongation at break or tensile strain was calculated by dividing the stretched length (mm) by the original length (mm) in both dry and wet states (Equation (3)).²⁹

T e n s i l e s t r e s s (\frac{N}{{m m}^{2}}) = \frac{M a x i m u m f o r c e a t b r e a k (N)}{I n i t i a l c r o s s s e c t i o n a r e a ({m m}^{2})}

(2)

% E l o n g a t i o n a t b r e a k = \frac{O r i g i n a l l e n g t h (m m)}{S t r e t c h e d l e n g t h (m m)} \times 100

(3)

Determination of the degree of cross-linking

The degree of hydrogel cross-linking is reflected by the average molecular weight between neighbouring cross-links. The equation used for calculating the degree of cross-linking (cross-linked density) is given below.³⁰

C r o s s - l i n k e d d e n s i t y (q) = \frac{M c}{M r}

(4)

Where Mc = molecular weight between adjacent cross-links, Mr = molar mass of the repeat unit

M_{c} = - \frac{d_{p} v_{2, s} (v_{2, s}^{1 / 3} - v_{2, s} / 2)}{\ln (1 - v_{2, s}) + v_{2, s} + x v_{2, s}^{2}}

(5)

Where, v₂,s = volume fraction of the swollen hydrogel in the equilibrium state (mL/mol), dp = density of hydrogel sample (g/mL), x = solvent interaction parameter

v_{2, s} = {[1 + \frac{d_{p}}{d_{s}} (\frac{M_{a}}{M_{b}} - 1)]}^{- 1}

(6)

Where, ds represent the density (g/mL) of the solvent. Ma and Mb denote the weights (g) of the hydrogel in the swollen and dry conditions, respectively.

x = \frac{\ln (1 - v_{2, s}) + v_{2, s}}{v_{2, s}^{2}}

(7)

M_{r} = \frac{m_{G G} M_{G G} + m_{b o r a x} M_{b o r a x}}{m_{G G} + m_{b o r a x}}

(8)

Where, m_GG and mborax are the masses of GG and borax. M_GG and Mborax are the molar masses of GG and borax, respectively.

Self-healing property

The self-healing ability was measured by a gel-block fusion method.³¹ The hydrogel film dressings were cut into 10 × 10 mm pieces, sprayed with SWF, and left for a short period before visually observing the self-healing effect. Subsequently, the healing time and healing efficiency of the hydrogel film dressings were determined. A tensile test was used to measure the healing efficiency of the self-healed hydrogel film dressing at various time intervals using equation (9).³²

H e a l i n g e f f i c i e n c y (%) = \frac{{M e c h a n i c a l v a l u e}_{h e a l e d}}{{M e c h a n i c a l v a l u e}_{i n i t i a l}} \times 100

(9)

Mechanical value_(healed) is the tensile stress after self-healing, and Mechanical Value_(initial) is the original tensile stress of the hydrogel film dressing.

Swelling behaviour of the hydrogel film dressing

The gravimetric tea bag method was utilized to determine the hydrogel’s swelling capacity.³³ Tea bags containing the hydrogel film dressings were immersed in a beaker containing 50 mL of SWF maintained at 37 °C. The hydrogel film dressing specimens were weighed (W₀) before immersion in the SWF. After fixed time intervals, the tea bags were removed, and the residual fluid was blotted away with tissue paper. The mass of the swollen hydrogel film dressing (W_s) was then immediately determined. The % swelling ratio was calculated using equation (10).³⁴

S w e l l i n g r a t i o (%) = \frac{W_{s} - W_{o}}{W_{o}} \times 100

(10)

W₀ = dry weight of the hydrogel film dressing, and W_s = weight of the swollen hydrogel film dressing.

In vitro degradation

The hydrogel film dressings were cut into pieces with an area of 10 × 10 mm² and immersed in 10 mL of SWF at 37°C.³⁵ At predetermined time intervals, the samples were removed from the medium and dried in an oven at 37°C until a constant weight was achieved. The degradation medium was refreshed every 2 days. The percentage weight loss was calculated using equation (11).

W e i g h t l o s s (%) = \frac{W_{i} - W_{f}}{W_{f}} \times 100

(11)

W_i = initial weight of hydrogel film dressing at time 0 and W_f = final weight of hydrogel film dressing at time t.

Antioxidant activity

The antioxidant potential of the hydrogel film dressing was evaluated using the reported procedure.³⁶ 0.1 mM solution of DPPH in ethanol (3 mL) was added to the 3 mL of hydrogel film dressing solution. After mixing, the solutions were left to stand for 40 min in a dark environment at ambient temperature. The absorbance of both the DPPH solution, and hydrogel film dressing solution was recorded at 517 nm using a UV-visible spectrophotometer (UV-2450, Shimadzu, Japan). The transition of the solution’s colour from a deep purple to yellow was observed as an indicator of antioxidant activity over these intervals.³⁷ Scavenging activity was determined using the following equation.

% S c a v e n g i n g a c t i v i t y = \frac{A_{o} - A_{1}}{A_{o}} \times 100

(12)

A₀ = absorbance of the 0.1 mM DPPH-ethanolic solution (control) and A₁ = absorbance of the test sample.

Antimicrobial study

The in vitro antibacterial activity of CMCS-loaded and ofloxacin-loaded hydrogel film dressings was evaluated using the disc diffusion method against Gram-positive (Staphylococcus aureus) and Gram-negative (Escherichia coli) bacteria. The bacterial colonies were suspended in nutrient broth, and the turbidity was maintained to 0.5 McFarland standards. The suspension of bacterial strains (100 µL) was uniformly spread on the nutrient agar plates. Hydrogel film dressings, cut into 50 mm² specimens, were placed on the inoculated agar plates and incubated for 2 days. The diameter of the zone of inhibition was determined to assess the antimicrobial activity of the hydrogel film dressing.^38,39

In vivo wound healing study

Approval for the animal study was granted by the Institutional Animal Ethical Committee (IAEC) at the West Bengal University of Animal and Fishery Sciences, Kolkata, India (Registration No.-763/GO/Re-S/ReRc-L/03/CCSEA). The approval number and date of approval are S/ReRc-L/03/CCSEA/92/2025–26 and 10.12.2025, respectively. Male Wistar rats aged 3 months were used for the study. Rats underwent a one-week acclimatization period in an animal house at 50% relative humidity, 25 °C temperature, and a 12 h light/dark cycle using a standard diet and water ad libitum. The animals were assigned to five groups, each consisting of three rats, with group names reflecting the type of treatment administered.

Group 1: control.

Group 2: treated with hydrogel film dressing (F4).

Group 3: treated with CMCS loaded hydrogel film dressing (F7).

Group 4: treated with ofloxacin loaded hydrogel film dressing (F8).

Group 5: treated with marketed film dressing (Tegaderm).

All surgical procedures were carried out under aseptic conditions, subsequent to intraperitoneal administration of anaesthesia with xylazine hydrochloride (8 mg/kg) and ketamine hydrochloride (80 mg/kg). Hair was removed from the dorsal region.

The area was sterilized with 70% ethanol, and full-thickness skin wounds of 8 mm diameter were created using an 8 mm biopsy punch.⁴⁰ The wound was covered with a prepared hydrogel film dressing and a marketed dressing. The wounds covered with a surgical gauze bandage were designated the control. Following the operation, meloxicam (0.2 mg/kg) was given intramuscularly daily for 3 days to manage pain. The wound dressings were substituted with new hydrogel film dressing every 3 days up to 14 days post-injury. Wound closure was assessed by capturing photographs on days 3, 7, 10, and 14. The healing of wound was determined by measuring wound size through photographic analysis using ImageJ software. The following formula was used to calculate the wound closure.

W o u n d c l o s u r e (%) = \frac{A_{t}}{A} \times 100

(13)

Where A_t was the area of the wound at time (t), and A was the initial wound area.

Quantitative evaluation of hydroxyproline in wounds

Tissue samples were collected from the wound site on days 7 and 14 post-wounding and stored at −20 °C until analysis. The samples were then examined to measure the hydroxyproline in the tissue using a hydroxyproline assay kit (Cat. No: E-BC-K061-S, Elabscience) according to the manufacturer’s instructions.⁴⁰

Histomorphological assessment of wound healing

Wound bed tissues were collected on days 7 and 14 post-wounding for histological analysis. The tissues were then fixed in 4% formaldehyde, dehydrated through a graded alcohol series, and embedded in paraffin. The paraffin-embedded tissues were sectioned into 3 µm slices using a microtome and stained with Hematoxylin & Eosin and Masson’s Trichrome.⁴⁰

Statistical analysis

Data are presented as mean ± standard deviation. p < 0.05 was used to make a statistically significant difference.

Results and discussion

Development and characterisation of a hydrogel film dressing

Various hydrogel film dressings were made using GG (non cross-linked, F1) and borax cross-linked GG (F2-F8). The cross-linked hydrogel film dressings (F2-F4) were prepared by varying the concentrations of GG (0.5-2%) and maintaining a fixed concentration of borax (1 wt% of GG weight). In aqueous solution, borax breaks down to form boric acid and the tetrahydroxyborate anion (Figure 1). The free 1,2 cis hydroxyl groups of GG with the tetrahydroxy borate anion form a borate ester linkage, resulting in the network structure.¹¹ On increasing the GG concentration to 2.5 % w/v, a highly thick and viscous solution formed, which is unsuitable for the continued development of the hydrogel film dressing. Therefore, the GG, having a concentration of 2% w/v, is used to develop the hydrogel film dressing further. Then, different concentrations of CMCS (9, 12, and 15 % of GG weight) were incorporated in the 2% GG solution and cross-linked with borax to develop a hydrogel film dressing (F5-F7). Further ofloxacin was loaded in the 2% GG solution and cross-linked with borax to develop a hydrogel film dressing (F8).

The cross-linking effect of borax and GG was characterized through FTIR and SEM analysis. The FTIR spectrum of GG showed a broad peak at 3570 cm^-1 corresponding to the stretching vibration of –OH (Figure 2(d)). A peak at 2924 cm^-1 is attributed to the −CH stretching vibration. The bands around 1650 cm^-1 can be ascribed to the ring stretching and bending of water molecules. The band at 1150 cm^-1 is attributed to the glycosidic linkage of GG. The peak at 1020 cm^-1 corresponding to CH₂ twisting was observed. The peak at 870 cm^-1 is due to skeletal stretching vibrations of galactose and mannose units. Mudgil et al. observed similar FTIR peaks of GG.⁴¹

The FTIR spectra of borax cross-linked GG hydrogel film dressings of F2, F3, and F4 are represented in Figure 2(a)–(c). The absorption peak of -OH at 3570 cm^-1 shifts to a lower wave number (3450 cm^-1) in F2, F3, and F4 borax cross-linked GG hydrogel film dressings. The –OH group of GG may be involved in the formation of borate ester linkages. The peaks at 2924 cm^-1 due to C-H stretching of GG is shifted to 2893 cm^-1. The intense peak appears at 1630 cm^-1, which is attributed to borate ester bond formation within the hydrogel network.⁴² The peaks observed at 1416 and 1333 cm^-1 for stretching of B-O-C, confirming the GG and borax interaction.⁴³ The peak at 876 cm^-1 for B-O stretching vibrations of free borate ions and the peak at 650 cm^-1 for B-O-B bending vibrations.⁴³ Similar results have been documented by other researchers.⁴⁴ Thus, the FTIR analysis showed that GG underwent cross-linking with borax. The characteristic cross-linking bands of F2, F3, and F4 appeared at nearly identical wavenumbers, indicating that the same cross-linking mechanism was maintained despite differences in GG concentration. However, minor variations in peak intensity and band broadening were observed among the formulations, which may be attributed to differences in polymer chain density and intermolecular interactions. The absence of significant peak shifts among F2, F3, and F4 suggests that increasing GG concentration primarily influences the degree of physical network formation rather than altering the chemical nature of the borax–GG cross-linking process.

SEM analysis of non cross-linked (NCL) and cross-linked hydrogel film dressing are represented in Figure 3. The NCL hydrogel film dressing exhibited a loose and irregular surface due to weak internal bonding (Figure 3(a)). The polysaccharide chains of GG form bonds with water via hydrogen bonding, which promotes the aggregation of the polymer chains into bundles. Moreover, a larger and more heterogeneous pore distribution was observed, indicating poor network formation. Cross-linking with borax induces the formation of borate ester complexes between free borate ions and the cis-1, 2-hydroxyl groups of GG, resulting in a more rigid and tightly packed polymer network (Figure 3(b)). Moreover, smaller and more defined pores were observed. This effect arises from disrupted hydrogen bonding caused by increased borate–diol complexation.

The degree of cross-linking of the F2, F3, and F4 hydrogel film dressings was determined (Table 2). A significant increase in the hydrogel film dressing degree of cross-linking (p < 0.05) was observed with the rise in GG concentration from 0.5% to 2% (w/v). The hydrogel film dressing degree of cross-linking of F2, F3, and F4 was 1.52 ± 0.103, 2.44 ± 0.104, and 2.95 ± 0.133, respectively. A higher GG concentration provides more hydroxyl groups for interaction with borate ions, leading to the formation of additional cross-linking junctions. Consequently, the increased borate-mediated cross-links produce a denser network structure, resulting in a higher degree of cross-linking.

Table 2.

Degree of cross-link of the hydrogel film dressing.

Formulation code	Degree of cross-link (Mean ± SD, n = 3)
F2	1.52 ± 0.103
F3	2.44 ± 0.104
F4	2.95 ± 0.133

The thickness of the hydrogel film dressing is a critical parameter influencing its properties. The ideal thickness for a dressing material for wounds is considered less than the thickness of the dermis of human skin, which normally ranges from 0.5 to 2.0 mm.²⁹ It was reported from the literature that hydrogel film dressing thicknesses smaller than 0.5 mm are appropriate for the healing of wounds.²⁹ The hydrogel film dressing thickness varies with changes in polymeric concentration, ranging from 0.11 ± 0.008 mm to 0.25 ± 0.004 mm (Table 1). The prepared NCL hydrogel film dressing (F1) has a thickness of 0.11 ± 0.008 mm. While cross-linked with borax, the thickness of the hydrogel film dressing was increased significantly (p < 0.05) to 0.13 ± 0.004 mm (F2). The cross-linking with borax resulted in a dense, porous network structure characterized by an increase in small pores.⁴⁵ As the pore density increases, a slight increase in hydrogel film dressing thickness is observed.

A significant increase in hydrogel film dressing thickness (p < 0.05) was observed with the rise in GG concentration from 0.5% to 2% (w/v). The hydrogel film dressing thicknesses of F2, F3, and F4 were 0.13 ± 0.004 mm, 0.17 ± 0.004 mm, and 0.21 ± 0.008 mm. An increase in polymer concentration facilitates the formation of a well-structured, denser three-dimensional matrix during gelation, thereby resulting in a thicker hydrogel film dressing.⁴⁶ However, incorporating CMCS at 9–15 wt% of the gum (F5–F7) resulted in a hydrogel film dressing with approximately the same thickness (Table 1). A small amount of CMCS may be accommodated within the voids of the cross-linked network. Similarly, the hydrogel film dressing (F8) containing ofloxacin exhibited a thickness comparable to that of the CMCS-incorporated hydrogel film dressing (Table 1).

Following an injury, the wound surface may lose and evaporate water up to twenty times faster than healthy skin.⁴⁷ The WVTR for normal skin was reported at 204 g/m²/24 h. However, in injured skin, WVTR increases from 279 g/m²/24 h (first-degree burns) to 5138 g/m²/24 h (granulating wounds).⁴⁸ The wound dressing should maintain an appropriate WVTR to replicate the skin’s natural environment. A suboptimal WVTR creates an overly occlusive environment that hinders wound healing by promoting maceration, increasing the risk of infection risk, and impairing gas exchange. Conversely, an excessively high WVTR leads to increased moisture loss, which can dry the wound and delay healing. The ideal WVTR for commercial wound dressings has been reported in the range of 100-3300 g/m²/24 h.⁴⁹

The WVTR of prepared hydrogel film dressing is determined and represented in Figure 4. The NCL hydrogel film dressing (F1) exhibited significantly (p < 0.05) higher WVTR than the cross-linked hydrogel film dressing (F2). The WVTR of F1 and F2 were 1611.57 ± 29.65 g/m²/24 h and 1534.26 ± 49.48 g/m²/24 h, respectively. NCL hydrogel film dressing forms a loose, more open polymer network, contains more free volume and interstitial spaces. Water vapour can diffuse easily through these gaps. Borax cross-linking creates a dense, tighter network that restricts the diffusion of water vapour. Shalmon et al. reported that non-cross-linked cellulose nanocrystal films had poorer barrier properties and greater water vapor penetrability than cross-linked ones.⁵⁰ A further increase in GG concentration significantly reduced WVTR (p < 0.05), with F2, F3, and F4 showing values of 1534.26 ± 49.48 g/m²/24 h, 1403.35 ± 47.85 g/m²/24 h, and 1081.45 ± 44.96 g/m²/24 h, respectively (Figure 4). Higher GG concentration increases chain density, forming a tighter matrix that limits water vapor pathways and lowers WVTR. Another author made a similar observation.⁵¹ Increasing CMCS concentration significantly reduced WVTR (p < 0.05), with F5, F6, and F7 showing values of 1056.25 ± 50.48 g/m2/24 h, 951.61 ± 72.19 g/m²/24 h, and 592.9 ± 55.10 g/m²/24 h, respectively. A denser, more compact matrix may form, reducing the transmission rate.⁵² In F8, the addition of ofloxacin resulted in a WVTR similar to the drug-free formulation (F4), and showing that the loading of the drug did not alter the hydrogel film dressing’s properties. F8 exhibited a WVTR of 1077.19 ± 51.35 g/m²/24 h.

The mechanical performance of a hydrogel film dressing is generally related to its structural integrity.⁵³ Protecting wounds from rupture during application, handling, or storage is crucial for effective treatment. Hydrogel film dressings must balance ductility, elasticity, and strength to accommodate skin movement. However, hydrolysis and enzymatic degradation are expected to sharply reduce these properties during use.⁵⁴ Tensile stress refers to the maximum tension a hydrogel film dressing can endure before failure,⁵⁵ with skin typically tolerating 2.5–16 MPa.⁵⁶ Assessing a hydrogel film dressing’s mechanical strength in both dry and wet states is essential, as absorbing wound fluid can weaken its integrity and cause wear during use.⁴⁸

Tensile stress decreased from dry to wet state, indicating a significant loss in mechanical and structural integrity (Figure 5). In wet conditions, water weakens intermolecular forces, increasing chain mobility and softening the material. Formulation F1 showed tensile stress values of 0.84 ± 0.047 MPa in the wet state and 4.14 ± 0.094 MPa in the dry state. Borax cross-linking significantly (p < 0.5) increased the tensile stress (F2), reaching 2.53 ± 0.081 MPa (wet state) and 7.24 ± 0.34 MPa (dry state), both higher than F1. The cross-linker bridges polymer chains, forming a cohesive structure that resists tensile deformation.¹¹ Tensile stress increased significantly (p < 0.05) with higher polymer concentrations (F2–F4). F2 showed the lowest values, followed by F3, while F4 exhibited the highest tensile stress in both wet (5.58 ± 0.235 MPa) and dry (11.33 ± 0.188 MPa) states. Increasing GG concentrations means more -OH groups are available, and more cross-link points can form per unit volume. A denser and more interconnected polymer network showed greater resistance to deformation and rupture. Moreover, higher concentration of GG reduces chain mobility and creates entanglements, and the hydrogel film dressing becomes stronger.¹¹ Tahmonzi et al. reported that the tensile strength improved with the increase of gum concentration.⁵⁷ Furthermore, the incorporation of CMCS improves tensile stress. An increase in CMCS concentration significantly raises tensile stress from F5 to F7 (p < 0.05). F5, F6, and F7 show tensile stress values of 6.37 ± 0.163 MPa & 12.8 ± 0.249 MPa, 7.33 ± 0.047 MPa & 13.95 ± 0.094 MPa, and 8.58 ± 0.08 MPa & 14.6 ± 0.18 MPa in wet and dry states, respectively. CMCS contains -OH, -NH_2, and -COOH groups. These groups electrostatically link polymer chains, strengthening the network and increasing tensile stress. Ofloxacin-loaded F8 showed tensile stress of 5.28 ± 0.094 MPa (wet state) and 10.9 ± 0.249 MPa (dry state), similar to the drug-free F4. Formulation F8 did not contain CMCS; instead, it contained only 1 wt% ofloxacin within the borax cross-linked GG matrix. Unlike CMCS-loaded formulations, F8 lacked additional intermolecular interactions contributed by CMCS functional groups. Therefore, its tensile strength remained comparable to the drug-free formulation (F4) and lower than the CMCS-loaded formulations.

The tensile strain test is conducted to study the flexibility of the polymer hydrogel film dressing. The percentage tensile strain determines the stretchability of the hydrogel film dressing before breaking. The tensile strain of the hydrogel film dressing in the wet state showed higher values than in the dry state. Water penetrates the polymer network and acts as a plasticizer, which reduces intermolecular forces (like hydrogen bonds) between the gum chains. This increases the mobility and flexibility of the polymer chains, allowing the hydrogel film dressing to stretch more before breaking, which increases tensile strain (elongation at break). Wang et al. showed that collagen-cellulose composite thin films in the wet state exhibited an increased elongation compared to the dry state.⁵⁸

The tensile strain of cross-linked hydrogel film dressing (F2) showed high values compared to NCL hydrogel film dressing (F1) in both wet and dry states (Figure 6). In the wet and dry states, % elongation at break of formulation F1 was 11 ± 2.64 % and 4.33 ± 1.52 %, while that of formulation F2 was 18.66 ± 3.05 %, and 12.66 ± 4.16 % in wet and dry conditions, respectively. Borax cross-links function as dynamic and reversible bonds. These slip and reform during mechanical stretching. This allows the hydrogel film dressing to extend further before fracturing, increasing the tensile strain (elongation at break). In contrast, the non-cross-linked hydrogel film dressing lacks this flexibility and often breaks at a lower strain. Further, with the increase in polymeric concentration, the interconnection of polymeric chains results in more complexation of molecules, and tougher hydrogels are formed. It was demonstrated that an increase in polymeric concentration (F2-F4), an increase in the tensile strain of the hydrogel film dressing. The F2, F3, and F4 hydrogel film dressings showed tensile strain of 18.66 ± 3.05, 29.66 ± 3.21% and 37 ± 3.6% at the wet state and 12.66 ± 4.16, 19.66 ± 2.08% and 24 ± 2% in the dry state, respectively.⁵⁹ When gum concentrations increase, more -OH groups are available for borax to form cross-links. More cross-linking leads to a strong yet flexible network, capable of absorbing and redistributing stress efficiently. Moreover, with more polymer present, there is greater physical entanglement of the chains, in addition to chemical cross-linkers. These entangled chains can slide and redesign under tension, increasing the hydrogel film dressing’s deformability before rupture. Incorporation of CMCS in the hydrogel film dressing increases the tensile strain in both the wet and dry states. The tensile strain of F5, F6, and F7 was 44 ± 2, 50 ± 4%, and 61.33 ± 4.16% in the dry state and 28.33 ± 0.57%, 29.33 ± 2.3%, and 37 ± 1% in the wet state, respectively. CMCS introduces flexible chains and hydrophilic groups (-COOH and -OH), which can reduce the stiffness of the hydrogel matrix. This makes the hydrogel film dressing more stretchable, hence increasing tensile strain. CMCS, being a hydrophilic and flexible biopolymer, introduces more deformable regions into the hydrogel matrix. As its concentration increases, the hydrogel becomes more elastic and capable of sustaining greater deformation before fracture, resulting in higher tensile strain. In F8, the addition of ofloxacin resulted in a tensile strain similar to the drug-free formulation (F4).

Self-healing hydrogels have reversible dynamic connections within the network, and the self-healing property in hydrogels allows the restoration of the original structure after an injury without compromising its functions.^60,61 The hydrogel film dressings were cut into two separate pieces and subsequently brought into contact after spraying with SWF. The damaged interfaces gradually fused together without the use of any external stimulus, demonstrating autonomous self-repair capability (Figure 7). The prepared cross-linked hydrogel film dressings undergo outstanding self-healing due to dynamic borate ester linkage formation with the free -OH groups of the GG. In aqueous medium, borax dissociates into tetrahydroxyborate ions, which dynamically interact with -OH groups of GG to form reversible borate–diol linkages. Upon mechanical damage, these dynamic bonds temporarily dissociate and subsequently reform at the fractured interface, restoring the polymeric network structure. The complexation of borate ions and hydroxyl groups on the adjacent polysaccharide chain is very fast.⁶² Unlike permanent covalent cross-links, the reversible borate–diol interactions continuously undergo bond dissociation and reformation, allowing structural rearrangement and healing of the hydrogel matrix.

Quantitative evidence of self-healing of the hydrogel film dressing was evaluated by tensile stress measurements at different healing intervals (Table 3). The healing efficiency progressively increased with healing time, confirming gradual restoration of mechanical integrity. The concentration of borax and polysaccharides greatly influences the healing efficiency of the hydrogel film dressing. With the increase in GG concentration, the healing efficiency increased. The healing efficiency of F2, F3, and F4 hydrogel film dressing at 6h were 56.552 ± 5.651%, 62.602 ± 3.92%, and 63.582 ± 2.962% respectively. Higher GG concentration means more -OH groups are available. This permits additional dynamic borate diol bonds, enhancing the material’s ability to reform cross-links at the damaged interface.

Table 3.

Percentage of healing efficiency of cross-linked hydrogel film dressings.

Time (min)	% Healing efficiency (Mean ± SD, n = 3)
Time (min)	F2	F3	F4	F5	F6	F7	F8
0	0	0	0	0	0	0	0
30	25.371 ± 2.38	25.473 ± 3.0	39.185 ± 3.02	40.229 ± 1.17	42.923 ± 2.23	43.538 ± 0.73	38.23 ± 2.31
60	40.196 ± 2.86	42.445 ± 2.02	52.268 ± 2.41	52.852 ± 3.15	55.467 ± 2.87	56.15 ± 1.75	50.24 ± 1.7
360	56.552 ± 5.651	62.602 ± 3.92	63.582 ± 2.962	69.264 ± 2.746	73.713 ± 3.737	76.04 ± 2.323	62.21 ± 1.86

Further incorporating and increasing the concentration of CMCS in a hydrogel film dressing enhances its self-healing efficiency. The formulations F5, F6 and F7 showed healing efficiency of 69.264 ± 2.746%, 73.713 ± 3.737% and 76.04 ± 2.323% respectively. CMCS contains -OH, -COOH, and -NH₂ groups. These groups enable secondary interactions such as hydrogen bonding, electrostatic interactions, and ionic complexation. These interactions act as physical cross-links that are also reversible, contributing to energy dissipation and healing capacity. With increasing CMCS concentration, more functional groups become accessible for secondary interactions. The optimized formulation F7 exhibited the highest healing efficiency, indicating substantial restoration of mechanical properties after damage. In F8, the addition of ofloxacin resulted in a healing efficiency similar to the drug-free formulation (F4).

The swelling capacity of hydrogel film dressing is crucial, as exudate absorption affects drug release and bioadhesion. Excessive swelling can cause overhydration and maceration, while insufficient swelling limits fluid uptake and delays wound healing. All hydrogel film dressing showed rapid initial swelling in SWF, reaching equilibrium after 6 h, followed by a gradual decrease due to matrix degradation. The NCL film dressing (F1) and cross-linked film (F2) showed maximum swelling of 302.79 ± 1.38% and 244.30 ± 1.33 %, respectively, at 6 h (Supplemental Figure S1). Cross-linking significantly reduced fluid absorption due to fewer available –OH groups and increased network density, which limited water diffusion compared to the NCL hydrogel film dressing. The swelling of the hydrogel matrix increased significantly (p < 0.05) with increasing GG concentration. The swelling percentage of formulations F2, F3, and F4 at 6h were 244.30 ± 1.33%, 422.215 ± 2.13%, and 525.54 ± 2.75 %, respectively. GG is hydrophilic, and its –OH groups readily form hydrogen bonds with water, while the cross-linked network prevents dissolution and maintains structural integrity.

The swelling capacity of borax cross-linked hydrogel film dressing increases upon incorporation and with a higher concentration of CMCS. Formulations F5, F6, and F7 show a significant (p < 0.05) increase in swelling of 527.60 ± 2.62%, 537.11 ± 1.03%, and 557.84 ± 5.48% as these formulations are loaded with 9%, 12% and 15% CMCS. CMCS undergoes deprotonation in a basic medium (SWF) and ionizes to carboxylate ions (COO^-). These carboxylate ions generate electrostatic repulsion with adjacent carboxylate ions and phosphate ions of SWF; as a result, the polymer chains become more relaxed, allowing greater water penetration.⁶³ Moreover, the functional groups of CMCS (-OH, -COOH, -NH2) are highly hydrophilic and exhibit a strong affinity for water molecules.

In vitro degradation is needed to determine hydrogel film dressing biodegradability in human tissue.⁶⁴ The loss of mass of the NCL hydrogel film dressing (F1) dropped by 80.8% after day 1 and lost 96.6% after the 2^nd day (Supplemental Table S1). The hydrogel film dressing’s mass loss was reduced when cross-linked with borax (F2). The mass loss of hydrogel film dressing F2 was 63.2% after 1 day and 87.8% after 2 days. GG swells and degrades quickly due to its hydrophilicity, but cross-linking forms a dense network that limits water uptake, swelling, and degradation. Hydrogel film dressing degradation is inversely related to GG concentration. After 1 day, degradation levels for the F2, F3, and F4 hydrogel film dressings were recorded at 63.2 ± 3.1%, 38.8 ± 1.01%, and 17.7 ± 1.49%, respectively (Supplemental Table S1). After 3 days, the F2 hydrogel film dressing is entirely degraded, whereas the F3 and F4 hydrogel film dressing take about 5 days to degrade 100%. Although swelling increased with increasing GG concentration (Supplemental Figure S1), cross-linked network prevents dissolution and maintains structural stability, thereby slowing the degradation rate. An increase in CMCS concentration leads to enhanced hydrogel film dressing degradation. The formulations F5-F7 has shown a degradation rate of 14 ± 2.461%, 17.4 ± 1.364% and 20.8 ± 1.405% on day 1 (Supplemental Table S1). After 6 days, F5, F6, and F7 were completely degraded. CMCS ionizes in SWF to carboxylate ions, promoting fast chain expansion and water absorption. Consequently, an increase in CMCS concentration leads to greater swelling, chain separation, and accelerated degradation. In F8, the addition of ofloxacin resulted in degradation similar to the drug-free formulation (F4), showing that the loading of drug did not alter the hydrogel film dressing’s properties.

Antioxidants prevent reactive oxygen species (ROS) production in wounds, necessitating their determination.⁶⁵ Antioxidant activity was observed in all hydrogel film dressings, with activity increasing as GG concentration rose. At 3 h of incubation, F2, F3, and F4 showed antioxidant activities of 38.23 ± 0.351%, 42.92 ± 0.411%, and 43.74 ± 0.823%, respectively (Supplemental Figure S2). GG contain multiple –OH groups, and these -OH groups can act as weak hydrogen donors, neutralizing free radicals like DPPH. Increasing the GG concentration increases the total number of -OH groups in the hydrogel film dressing, leading to more potential radical-scavenging sites. Antioxidant activity increases with the incorporation and increasing concentration of CMCS. The antioxidant activity of F5, F6, and F7 were 45.542 ± 1.206%, 47.108 ± 1.747%, and 53.132 ± 1.024%, respectively. CMCS is a water-soluble derivative of chitosan that possesses –NH₂, –CH₂–COOH, and –OH groups.

These hydrogen-donating groups neutralize free radicals, and more CMCS increases their number, boosting scavenging activity.

Among the eight formulations (F1-F8), F7 exhibited the highest tensile strength (14.6 MPa in the dry state; 8.58 MPa in the moist state) and tensile strain (61.33% in the dry state; 37% in the wet state), outperforming the other hydrogel film dressings. F7 also demonstrated a high swelling capacity (557.84% at 6 h), sufficient to manage wound exudates while maintaining structural integrity. F7 showed WVTR of 592.9 g/m²/24 h, greater than the WVTR for normal skin (204 g/m²/24 h). The F7 hydrogel film dressing exhibited the highest healing efficiency of 76.04% at 6h and was entirely degraded in 6 days. In addition, F7 showed the highest antioxidant activity of 53.13%. Considering the above parameters, F7 was selected as the optimized formulation.

Microorganisms adhere to the dressing, thereby enhancing the chance of infection.⁶⁶ The antibacterial activity was evaluated against Staphylococcus aureus (Gram-positive) and Escherichia coli (Gram-negative) using hydrogel film dressings with 9–15% CMCS and 1% ofloxacin.

(Figure 8). The 9% CMCS-loaded hydrogel film dressing (F5) exhibited no zone of inhibition against either Staphylococcus aureus or Escherichia coli (Supplemental Table S2). The 12% CMCS loaded hydrogel film dressing (F6) showed a 2.02 cm inhibition zone against Escherichia coli, but was inactive against Staphylococcus aureus. However, the 15% CMCS-loaded hydrogel film dressing (F7) exhibited a 2.26 cm inhibition zone against Escherichia coli and a 1.56 cm zone of inhibition against Staphylococcus aureus. At 9%, the amount of CMCS in the hydrogel may be too low to interfere with bacterial membrane structure, resulting in no observable antimicrobial effect. At 12% CMCS, the concentration may be sufficient to disrupt Escherichia coli’s outer membrane but still insufficient to penetrate Staphylococcus aureus. Escherichia coli is a Gram-negative bacterium with a thinner peptidoglycan layer.⁶⁷ Staphylococcus aureus is Gram-positive with a thicker peptidoglycan wall, which can offer more resistance to penetration by CMCS.⁶⁸ At 15%, the CMCS concentration is sufficiently high to overcome the defences of both Escherichia coli and Staphylococcus aureus. The drug loaded hydrogel film dressing exhibited a 3.68 cm zone of inhibition against Escherichia coli and a 3.26 cm zone of inhibition against Staphylococcus aureus. The 1% ofloxacin-loaded hydrogel film dressing exhibited greater antimicrobial effectiveness against both Escherichia coli and Staphylococcus aureus than the 15% CMCS-loaded hydrogel film dressing.

Ofloxacin exhibits high potency, spreads readily through agar, and effectively eradicates microorganisms.⁶⁹ The antibacterial activity of CMCS is mediated through membrane disruption, with no development of microbial resistance. In contrast, ofloxacin may induce resistance, whereas CMCS is considered safer.⁷⁰ Hydrogel film dressings were more active against Escherichia coli than Staphylococcus aureus, as Escherichia. coli has a thin peptidoglycan layer, allowing the drug and CMCS to enter more easily. Damiana et al. carried out the antimicrobial study of ciprofloxacin loaded hydrogel prepared using trimethyl chitosan/sodium carboxymethyl xanthan gum against Escherichia coli and Staphylococcus aureus.⁷¹ Greater antibacterial effects against Escherichia coli than Staphylococcus aureus were observed with the developed hydrogel.

Wounds were produced in rats to assess the in vivo healing efficacy of the hydrogel film dressing. Hydrogel film dressing (F4), CMCS-loaded hydrogel film (F7), ofloxacin-loaded hydrogel film (F8), and the marketed dressing (Tegaderm) were applied to the wounds. Visual inspection during hydrogel film dressing replacement revealed a moist wound environment, indicating good performance. The animals remained in good health, with no adverse effects observed during the experiment. No redness, irritation, or signs of infection were observed. Photographs were captured on days 0, 3, 7, 10, and 14, and the percentage of wound closure was calculated (Figure 9).

At 3 days post-injury, the marketed wound dressing showed a significantly higher rate of wound closure than the other groups (p < 0.05). The extent of wound closure, expressed as a percentage, in the F4, F7, F8, and marketed dressing groups was 11.13 ± 0.005%, 30.68 ± 0.002%, 17 ± 0.002%, and 41.47 ± 0.084%, respectively (Supplemental Table S3). At 7, 10, and 14 days post-wounding, the CMCS-loaded hydrogel film dressing exhibited greater wound closure compared to all other groups. At 14 days post-wounding, the wound closure percentages in the F4, F7, F8, and marketed dressing groups were 80.81 ± 0.012%, 94.3 ± 0.008%, 89.56 ± 0.004%, and 94.03 ± 0.004%, respectively. Tegaderm showed rapid initial wound reduction due to its ability to maintain a moist environment and protect the wound from contaminants. Angiogenesis, cell migration, and tissue repair processes were enhanced after the inflammatory phase, with sustained CMCS release providing support during this stage, normally emerging after day 3.⁷² CMCS promotes wound healing by reducing inflammation and oxidative stress.⁷³ Additionally, during the later stage of wound healing, keratinocyte and fibroblast adhesion, proliferation, and migration occur, and CMCS contributes to these processes.⁷⁴

Collagen is a key structural protein that imparts strength to connective tissues, and hydroxyproline, a collagen-specific amino acid, serves as a marker for collagen content.^75,76 The hydroxyproline content was measured in the granulation tissue from the wound area on days 7 and 14, and the hydroxyproline content increased from day 7 to day 14 (Figure 10). Day 7 typically corresponds to the proliferative phase of healing, during which fibroblasts are activated and begin producing collagen. Day 14 marks the early remodelling phase, characterized by increased collagen deposition, maturation of granulation tissue, and improved wound strength. Therefore, the rise in hydroxyproline content reflects enhanced collagen synthesis and deposition over time. At both 7 and 14 days post-wounding, the CMCS-loaded hydrogel film dressing exhibited significantly higher hydroxyproline levels than all other groups (p < 0.05). At 7 days post-wounding, the hydroxyproline content in the F4, F7, F8, and marketed dressing groups was 0.75 ± 0.028 µg/mg, 1.73 ± 0.073 µg/mg, 1.5 ± 0.012 µg/mg, and 1.62 ± 0.049 µg/mg, respectively (Figure 10). At 14 days post-wounding, the hydroxyproline content in the F4, F7, F8, and marketed dressing groups increased to 1.23 ± 0.124 µg/mg, 2.19 ± 0.064 µg/mg, 1.8 ± 0.046 µg/mg, and 1.97 ± 0.02 µg/mg, respectively (Figure 10). CMCS has been shown to enhance the proliferation and migration of fibroblasts, the primary collagen-producing cells.⁷⁷ This directly contributes to increased collagen deposition, as reflected by higher hydroxyproline content. Additionally, CMCS contains functional groups (–NH₂ and –COOH) capable of interacting with cells and promoting the expression of collagen-related genes (COL1A1 and COL3A1).⁷⁸ This biomolecular stimulation results in the upregulation of collagen biosynthesis.

The histological evaluation of wounds treated with the developed hydrogel film dressings revealed various characteristics throughout the healing period (Figure 11). The microscopic observations were based on the presence and distribution of inflammatory cells (a), loosely packed granulation tissue (b), new blood vessels (c), fibroblast (d), epithelialization (e), dividing fibroblast (f), and densely packed granulation tissue (g). On day 7, wounds in the control group exhibited an abundance of inflammatory cells (a), indicating a prolonged inflammatory phase. The tissue architecture appeared disorganized with sparse fibroblast proliferation and limited epithelialization. Moreover, wounds exhibited poor epithelial coverage, haemorrhage, and insufficient granulation tissue. The F4 group exhibited reduced inflammatory cells (a), moderate fibroblast proliferation, early epithelial tongue formation, and loosely packed granulation tissue (b) indicating progression from the inflammatory phase toward tissue regeneration. Similarly, the F8 group showed diminished inflammatory cell infiltration (a) with moderate fibroblast activity (d) and re-epithelialization. The F7 group showed a reduction in inflammatory cells (a), enhanced re-epithelialization, early granulation tissue formation (b), and the presence of new blood vessels (c), indicative of angiogenesis. The Tegaderm group exhibited minimal inflammatory infiltration (a) along with modest epithelial restoration. Although fibroblast activity (d) and granulation tissue formation (b) were evident, the degree of tissue organization appeared less pronounced than in the F7-treated wounds.

On day 14, the control group still showed residual inflammatory cells (a), incomplete epithelial coverage, and poorly organized connective tissue, indicating delayed wound maturation. In contrast, the F4, F8, and Tegaderm groups showed considerable improvement in tissue regeneration. These groups showed reduced inflammatory infiltrates (a), increased fibroblast populations, and enhanced epithelialization (e). Granulation tissue became more organized, although complete tissue maturation had not yet been achieved. The F7 group exhibited a well-organized epidermal structure, complete epithelialization (e), and a mature granulation layer. The wound bed was characterized by densely packed granulation tissue (g), abundant fibroblasts (d), and well-developed connective tissue architecture with minimal residual inflammation.

Collagen deposition in the wound was evaluated using Masson’s trichrome staining (Figure 12). Histological examination focused on the distribution and organization of collagen fibres (a), keratin fibres (b), muscle fibres (c), fibroblasts (d), and newly formed blood vessels (e). On day 7, the control group showed loosely arranged collagen fibres (a), inadequate keratin fibre formation (b), and poorly organized muscle fibres (c), indicating incomplete tissue reconstruction. The F4 group showed a moderate increase in collagen fibre deposition (a). Keratin fibres (b) and muscle fibres (c) were visible, indicating active matrix production. Similarly, the F8 group showed increased collagen accumulation (a). Keratin fibres (b) and muscle fibres (c) were evident, indicating ongoing tissue regeneration. The Tegaderm group showed adequate collagen deposition (a) and fibroblast activity (d), accompanied by the presence of keratin fibres (b) and muscle fibres (c). On day 7, the F7 group exhibited denser collagen deposition (a) along with early development of the extracellular matrix. The presence of numerous fibroblasts (d), keratin fibres (b), muscle fibres (c), and newly formed blood vessels (e) indicates enhanced angiogenesis and accelerated tissue reconstruction and maturation. On day 14, the control group exhibited loosely arranged and disrupted collagen fibres (a). The F4, F8, and Tegaderm groups exhibited increased collagen deposition (a) compared with day 7, accompanied by enhanced fibroblast proliferation (d). Muscle fibres (c) were more prominent, indicating ongoing tissue regeneration. However, on day 14, the F7 group exhibited thick, well-organized collagen bundles (a), substantial population of fibroblasts (d), keratin fibres (b) and muscle fibres (c). On day 14, the F7 group showed elevated hydroxyproline levels, indicating a high healing efficiency.

Conclusion

A self-healing hydrogel film dressing was developed using borax cross-linked GG incorporated with CMCS to enhance its bioactivity and promote effective wound healing. The optimized hydrogel film dressing loaded with CMCS exhibited the highest tensile strength (14.6 MPa in the dry state; 8.58 MPa in the moist state) and tensile strain (61.33% in the dry state; 37% in the wet state), outperforming the other hydrogel film dressings. It also demonstrated a high swelling capacity (557.84% at 6 h), the highest healing efficiency of 76.04% at 6h and the highest antioxidant activity of 53.13%. The optimized hydrogel film dressing revealed broad-spectrum efficacy against Staphylococcus aureus and Escherichia coli. In a rat full-thickness wound model, CMCS-loaded hydrogel achieved superior wound closure compared to Tegaderm and drug-loaded dressings. Elevated hydroxyproline levels and enhanced collagen deposition on day 14 in the CMCS-loaded hydrogel group highlight biochemical and histological alignment in healing efficiency.

Supplemental material

Supplemental material - In vitro and in vivo evaluation of a bioactive polysaccharide based self-healing hydrogel film for enhanced wound healing

Supplemental material for In vitro and in vivo evaluation of a bioactive polysaccharide based self-healing hydrogel film for enhanced wound healing by Ayan Ranjan Hati, Chalamala Balganghadhar Reddy, Purnendu Ghosh, Kaushik Mukherjee, Amalesh Samanta, Samit Kumar Nandi, Tapan Kumar Giri in Journal of Biomaterials Applications

Footnotes

Acknowledgment

The authors are thankful to West Bengal University of Animal and Fishery Sciences, Kolkata for conducting in vivo study.

ORCID iDs

Chalamala Balagangadhar Reddy

Purnendu Ghosh

Kaushik Mukherjee

Amalesh Samanta

Samit Kumar Nandi

Tapan Kumar Giri

Ethical considerations

The in vivo study was approved by the Institutional Animal Ethical Committee (IAEC) at the West Bengal University of Animal and Fishery Sciences, Kolkata, India (Registration No.-763/GO/Re-S/ReRc-L/03/CCSEA).

Funding

The authors received no financial support for the research, authorship, and/or publication of this article.

Declaration of conflicting interests

The authors declared no potential conflicts of interest with respect to the research, authorship, and/or publication of this article.

Data Availability Statement

All data and materials are included in the manuscript.*

Supplemental material

Supplemental material for this article is available online.

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