Designing biomimetic hydrogels for neuro-therapeutic delivery devices for brain soft tissue injury: integrating antioxidant, cell viability and tissue adhesion properties to enhance neural regeneration via a synergistic approach

Abstract

Recent progress in brain research reflects an exciting interface of technology and biology, leading to the development of effective therapeutic compounds and site specific delivery systems to address complex neurological disorders. These innovative therapies also serve as novel diagnostic and therapeutic platforms for drug delivery (DD) to the brain and soft tissue in the context of brain injury, aiming to enhance drug penetration and targeting while improving efficacy and minimizing systemic toxicity. Hence, the innovative approach of this project lies in the development of a network structure in the form of hydrogels derived from bioactive sulphated polysaccharide & zwitterionic polymers by the copolymerization technique for the delivery of the neuroprotective & neurorestorative (citicoline) compound at the site of nerve injury. The biocompatibility, protein adsorption, antioxidant, mucoadhesion, drug delivery and cell-viability of rhabdomyosarcoma cell properties of hydrogels were analyzed. Hydrogels expressed 165 ± 0.19% cell viability of RD cells and promoted cell adhesion and proliferation, signifying their compatibility with mammalian cells. DPPH assay revealed 39.82 ± 1.65% free radical scavenging ability of the materials, highlighting their strong intrinsic antioxidant potential for neutralizing oxidative stress at the site of nerve injury. The mucoadhesion of the materials was signified from a force of 75 ± 4.00 mN, desirable for adherence to mucosal surfaces and helps in cell attachment and alignment during the nerve regeneration process. The citicoline anchored brain drug delivery carrier released the drug in simulated brain fluid in a sustained pattern and followed the non-Fickian diffusion mechanism. The release profile was best explained by the Hixson–Crowell kinetic model. The materials were also characterized by FESEM, EDAX, AFM, FTIR, ¹³C-NMR & XRD techniques. Overall, the presented synergistic therapy for treatment of brain injury involved the delivery of the bioactive nerve regenerating agent from functional materials. It will not only deliver therapeutic molecules to nerve injuries but its inherent antioxidant, haemostatic & non-cytotoxic nature with cell viability properties may also contribute to enhancing the nerve repair process of brain injury.

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1 Introduction

Recent brain research highlights advancements, particularly in designing new drug delivery systems for the treatment of neurological disorders. These innovative therapies act as new diagnostic tools for brain and soft tissue DD in the context of injury for enhancing drug penetration and targeting to improve efficacy and reduce toxicity. Brain-specific challenges include overcoming the blood–brain barrier, while soft tissue injuries may require localized delivery to the affected area. Globally, people are affected by severe nerve injuries or neurodegenerative disorders, presenting a significant challenge to human health. These disorders are devastating and can cause lifelong disabilities. The spinal cord is vulnerable to acute injuries that trigger severe physiological or immune responses, inflammation, cytotoxicity & oxidative stress.¹ These processes collectively impede functional recovery, resulting in persistent neurological deficits.² Despite improvement in various therapeutic strategies for nerve regeneration, functional recovery is still limited and the healing process is complicated due to the long recovery time, pathological mechanism and unique biological microenvironment of the central nervous system (CNS).^3,4 Conventional approaches inhibit inflammation at the site of nerve injury but at the same time, they also suppress immune response in other parts of the body & also increase susceptibility to bacterial infection.⁵ Emerging strategies have gained attention in addressing various issues related to brain drug delivery (BDD) devices for neural tissue regeneration owing to their potential to provide a constructive approach for delivering neuro-therapeutics and mitigate inflammation at the injury site.⁶

Advanced biomedical interventions for delivering drugs at the site of nerve injury involved the use of hydrogel based BDD platforms to promote the delivery of therapeutic agents to the target site and support the regeneration and reconstruction of damaged tissue.⁷ Hydrogel based BDD reduced the injury related neuro-inflammation and modulated the microenvironment, which resulted in reduced lesions and tissue repair and regeneration. These hydrogel based scaffolds are promising ways of releasing neuro-regenerative agents directly to the site of injury without getting affected by the blood–brain barrier (BBB) and the blood–medullary barrier.⁸ The injured area recruits local macrophages called microglia that activate & regulate neuro-inflammation. Persistent inflammatory responses led to apoptosis or neuronal atrophy.^9,10 Wang and coworkers¹¹ evaluated the release of quercetin from liposomal gel to achieve antioxidant, anti-inflammatory potential and neuro-protective effects for the treatment of CNS injuries. These gels facilitated the penetration of quercetin across the brain–spinal cord barrier and enhanced its bioavailability. In addition, they also promoted polarization of microglia, enhancing secretion of anti-inflammatory mediators and facilitated axonal growth along with reduced glial scars. Ai and coworkers¹² developed hydrogels for functional recovery in spinal cord injury by mitigating inflammation by enhancing axonal extension and myelin regeneration. Tao and coworkers¹³ developed agar-poly(pyrrole) based hydrogels for the repair of spinal cord injury. Materials exhibited swelling and conductivity traits, effectively filling the spinal cord cavity & reconstructing electrical conduction pathways. This facilitated electrical signal transmission, hence promoting locomotor recovery in rats.

In addition, bioactive materials used in nerve regeneration can add unique features and advantages. Carrageenan (CG) is a bioactive sulphated polysaccharide derived from seaweeds and exists in many forms owing to the degree of sulfation & position of sulfate groups in its structure. However, iota-CG is composed of more sulphated functionalities and has two subunits, namely 3-linked β-D-galactopyranose-4-sulphate (G4S units) and 4-linked 3,6-anhydro-α-D-galactopyranose-2-sulphate (DA2S units). In addition to sulfate groups, xylose, glucose, uranic acids are also found in the structure of CG.¹⁴ The presence of an additional sulfonate ester group in its structure makes it more efficient for various DD applications. These functional moieties provide anticoagulant, anticancer, anti-thrombogenic and antioxidant properties.^15–17 CG is widely used in the pharmaceutical industry to improve drug formulation properties, especially to prolong drug release. CG has a tendency to form complexes with drugs and hence can improve drug loading, improve the dissolution rate of poorly water-soluble drugs, and provide sustained DD.^18,19 Poly(2-(methacryloyloxy)ethyl dimethyl-(3-sulfopropyl)ammonium hydroxide) [poly(MEDSAH)] is a zwitterionic polymer used in hydrogel based applications due to its excellent hydrophilicity and biocompatibility properties.²⁰ Poly(acrylic acid) [poly(AAc)] is a hydrophilic, pH-responsive & bioadhesive polymer.²¹ Citicoline is a neuroprotective compound, particularly employed for nerve regeneration and nerve repair. Citicoline provides cytidine & choline that served as substrates for the synthesis of phosphatidylcholine, a key compound of neuronal cell membranes.²²

The innovative approach of this project is based on the fact that it is a development of network structure in the form of hydrogels from bioactive sulphated polysaccharide & zwitterionic polymers for delivery of the neuro-protective & neuro-restorative (citicoline) compound at the site of nerve injury. Drug-anchored hydrogels as brain DD carriers offered a functional platform that synergistically integrates biochemical, biophysical, and pharmacological attributes to facilitate nerve regeneration. Herein, sustained release from bioactive hydrogels may execute biomimetic action like the antioxidant action of enzymes, vitamins and hormones under the self-repair tissue regeneration process.

2 Materials and methods

2.1 Materials used

Iota-carrageenan [HiMedia Laboratories Pvt. Ltd], acrylic acid [Merck specialties Pvt Ltd (Mumbai) India], [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide (MEDSAH) [Sigma-Aldrich, USA], N N-methylenebisacrylamide (NN-MBA) and ammonium persulfate (APS) [Fisher Scientific India Pvt Ltd., Mumbai, India) were materials used to design network hydrogels. Citicoline was purchased from NEON laboratories Ltd., Mumbai India.

2.2 Synthesis of brain drug delivery carriers

Synthesis of hydrogels as brain drug delivery carriers was carried out by the graft copolymerization reaction of poly(MEDSAH) and poly(AAc) onto CG polysaccharide with the addition of the NN-MBA crosslinker and APS initiator. Polymer synthesis was initiated by adding a solution containing a fixed concentration of free radical initiator APS of 2.1 × 10⁻² mol L⁻¹ into a solution of definite content of CG = 10% w/v. Thereafter, a definite amount of vinyl monomers [MEDSAH] = 3.57 × 10⁻² mol L⁻¹, [AAc] 10.18 × 10⁻¹ mol L⁻¹ and a specified content of cross linker [NN-MBA] = 12.90 × 10⁻³ mol L⁻¹ were added to the reaction system. The reaction was then continued for 3 h at 65 °C temperature for completion of the polymerization process. Ultimately, the crosslinked network hydrogel was formed which was referred to as CG-cl-poly(poly(MEDSAH)-poly(AAc) hydrogel. Optimization of reaction parameters was obtained by altering the composition of monomers [MEDSAH] from 3.57 × 10⁻² mol L⁻¹ to 17.89 × 10⁻² mol L⁻¹, [AAc] from 4.36 × 10⁻¹ mol L⁻¹ to 13.08 × 10⁻¹ mol L⁻¹ and crosslinker from 6.4 × 10⁻³ mol L⁻¹ to 32.43 × 10⁻³ mol L⁻¹ during the copolymerization reaction, affecting the swelling capacity of the hydrogel. The maximum fluid absorption capacity of hydrogels was achieved under optimum conditions of [MEDSAH] = 3.57 × 10⁻² mol L⁻¹, [AAc] = 10.18 × 10⁻¹ mol L⁻¹ and [NN-MBA] = 12.90 × 10⁻³ mol L⁻¹. Optimized network hydrogels were subjected to DD analysis to evaluate their structural and functional performance and biomedical assessment. A pictorial representation of the synthesis of hydrogels and subsequent drug loading and release at the site of nerve injury from the hydrogel matrix is given in Scheme 1.

Scheme 1 A pictorial representation of the synthesis of hydrogels and subsequent drug loading and release at the site of nerve injury from the hydrogel matrix.

2.3 Characterization

Polymer characterization was performed using different analytical instruments. FESEM and EDAX were conducted on a JEOL-JSM6100 SEM; AFM analysis was performed using an INTEGRA, NT-MDT, Russia; FTIR analysis of the polymer sample in a KBr pellet was carried out on a BRUKER ALPHA-Platinum-ATR-IR; ¹³C-NMR solid state spectra of samples were recorded on a JEOL-ECZR600 NMR; and XRD analysis was performed on a PAN-ANALYTICAL X'PERT PRO INSTRUMENT.

2.4 Physico-chemical and biomedical properties

2.4.1 Swelling analysis. Swelling of copolymers was determined by the gravimetric method, wherein dry polymer samples were initially weighed and then immersed in buffer solution. Samples were then removed from the solution periodically and weighed to determine the swelling percentage relative to their initial weight.

2.4.2 Drug delivery properties. Drug encapsulation within hydrogels was performed using the swelling equilibrium method. For drug encapsulation, the sample was immersed in a drug solution of fixed concentration (citicoline = 500 μg mL⁻¹) at 37 °C for a specific time duration. Release of citicoline from encapsulated hydrogels was conducted by keeping the drug containing sample in a solution of different pH release media. Drug loading and release of citicoline were calculated from calibration curves of citicoline prepared in different media (distilled water λ_max = 270 nm), pH 2.2 buffer (λ_max = 210 nm), pH 7.4 buffer (λ_max = 270 nm) and ACSF (artificial cerebrospinal fluid) buffer (λ_max = 270 nm), using a UV spectrophotometer. The absorbance of the released drug was determined with a UV-vis spectrometer. The Ritger and Peppas equation (eqn (i)) was applied in the evaluation of the diffusion mechanism for the citicoline release gel matrix.^23,24

[(M_t/M_∞) = ktⁿ] (i)

where k (polymer characteristic) and n (diffusion exponent) are constants. M_t/M_∞ is a fraction of the drug released at time t. Release profile data were applied in various equations to describe the kinetic model for drug diffusion from infused samples.^25,26

2.5 Biomedical properties

Blood compatibility of the material was assessed by examining polymer–blood interactions in terms of thrombogenicity and haemolysis parameters.^27–29 The protein adsorption properties of hydrogels were determined using bovine serum albumin (BSA) as a reference protein by using the Lowry method.^30,31 Cell viability assay for the hydrogels was performed.³² The transformed cell line namely RD (rhabdomyosarcoma cells) was procured from Central Research Institute, Kasauli (India). Cells were maintained in DMEM (Dulbecco's modified Eagle medium) supplemented with 10% foetal bovine serum, 100 IU mL⁻¹ penicillin and 100 μg mL⁻¹ streptomycin. The cell lines were cultured in a 4% CO₂ incubator (95% humidity) at 37 °C. The cytotoxic effect of the polymer sample was checked after treating cells (RD) with a polymer sample of 1 mm length and 1 mm width. Then the effect of the polymer on RD cell lines (MTT Assay) was determined. Mean values of all measurements were used to calculate the cell viability percentage (eqn (ii)). The treated group of cells was compared with the control group in the absence of compounds.

Cell viability (%) = (A₅₇₀ of treated cells)/(A₅₇₀ of control) × 100 (ii)

Antioxidant activity (AOX) analysis of hydrogels was performed by 2,2′-diphenylpicrylhydrazyl (DPPH), Folin–Ciocalteu (FC) and phosphomolybdate (PM) reagent assays.^33,34 Mucoadhesion was determined by testing the detachment force required to separate the copolymer from the goat intestine membrane.³⁵ The gel strength (Nmm) of swollen hydrogels was determined by using a texture analyzer (Stable Micro system, UK) equipped with a 5 kg load cell to record the penetrating force required by immersing the probe into the hydrogel up to 3 mm depth. The polymer sample after 24 h water uptake was subjected to a trigger force of 0.01 N from the instrument and the cylindrical probe of 10 mm (P/10) diameter was allowed to penetrate into the hydrogel up to a penetration depth of 3 mm. At this depth the maximum force was recorded as resistance provided by the gel to penetration which gave the ‘gel strength’ of a sample (n = 3). The test was performed at fixed instrumental parameters i.e. pre-test speed = 1.0 mm s⁻¹, test speed = 0.5 mm s⁻¹, post-test speed = 1.0 mm s⁻¹, strain = 10% and acquisition rate = 200 points per second.

3 Results and discussion

3.1 Characterization

3.1.1 FE-SEM. FE-SEM analysis was carried out to characterize the surface morphology of the hydrogel (Fig. 1). FESEM images of the polysaccharide appeared smooth and homogeneous, whereas FESEM images of the hydrogel showed the heterogeneous & irregular morphology formed by the grafting or cross-polymerization reaction of monomers with CG.³⁶ These structures further provide a porous architecture to the polymer material and provide more surface area for easy passage of fluids or neuro-therapeutics to the target site.³⁷ Furthermore, these pores also serve as sites wherein hydrophilic functional groups interact with external stimuli and support more absorption and retention of drugs or body fluids. This characteristic makes the CG based hydrogel suitable for brain targeted DD applications, as it allowed efficient drug diffusion and promoted sustained drug release.³⁸

Fig. 1 FESEM images of (a) CG and (b) CG-cl-poly(MEDSAH)-poly(AAc) polymers, EDAX of (c) CG and (d) CG-cl-poly(MEDSAH)-poly(AAc) polymers and AFM of (e) CG-cl-poly(MEDSAH)-poly(AAc) polymers.

Khan and coworkers³⁹ revealed that FESEM images of poly(AAc) based hydrogels showed porous network structures owing to a covalent crosslinking reaction. These interconnected pores allowed easy flow of bioactive agents or drugs into and out of the gel matrix.⁴⁰ Bolanta and coworkers⁴¹ demonstrated that FESEM images of poly(AAc)–cysteine based hydrogels showed stacked or layered surfaces as a result of the copolymeric nature of hydrogels. The stacked morphology of the polymer favoured the formation of porous layers on swelling.⁴² Sairaman and coworkers⁴³ revealed that FESEM images of CG based hydrogels showed a porous morphology which is useful for tissue regeneration as it provides a matrix for differentiation of stem cells. Liu and coworkers⁴⁴ prepared tetramethylpyrazine loaded PVA hydrogels for sustained delivery at the nerve injury site. They revealed that these hydrogels are also featured with porous structure that supported axonal regeneration and synaptic reconstruction at the site of injury. Tao and coworkers¹³ developed a network structure for controlled drug release. These pores are beneficial to the DD carrier at the site of nerve injury.

Hence, hydrogels with an uneven network structure are useful in creating sustained DD systems owing to their ability to encapsulate therapeutic agents and subsequently release drugs in a controlled manner. Porous structures for DD of nerve regenerating agents allowed controlled release of neurotropic factors, directly at the site of nerve injury and hence facilitated growth of nerve cells to promote nerve regeneration.⁴⁵

3.1.2 EDAX. EDAX analysis is carried out to identify constituent elements present in the material (Fig. 1). EDAX spectra of CG contain peaks of C, O and S elements, whereas the copolymer spectrum displays an additional peak for nitrogen, along with an increase in the amount of sulphur content (from 0.94% to 9.62%). This clearly indicates the introduction of new components during the cross-polymerization reaction. EDAX spectra confirmed the crosslinking reaction and formation of effectively integrated poly(MEDSAH) onto CG in the presence of crosslinker NN-MBA, evidenced by the presence of the nitrogen peak and elevated sulphur content. Perumal and coworkers⁴⁶ demonstrated that the EDAX spectrum of the poly(MEDSAH) based composite exhibited peaks referring to C, S, and N, confirming the polymerization of poly(MEDSAH). Similarly, the presence of C, O, N & S elements in the EDAX spectrum verified the successful incorporation of zwitterions, poly(MEDSAH) polymers, in the presence of crosslinker NN-MBA.⁴⁷

3.1.3 AFM. AFM analysis of hydrogels allowed investigation of their surface topography & assessment of the surface roughness. Values of root mean square roughness and average roughness were found to be 34.426 nm and 26.853 nm, respectively (Fig. 1). The roughness in the polymer surface was attributed to the incorporation of poly(MEDSAH) and poly(AAC) onto CG by crosslinking reactions that yield a polymer with a rough morphology. Introducing new functional groups in polysaccharide chains formed covalent bonds and created a crosslinked network, which results in the formation of a rough polymer surface.⁴⁸

Polymers with a rough surface are beneficial in delivering the drug to the site of nerve injury by increasing the surface area for cell adhesion and promoting a more favourable environment for cell growth. Moreover, a rough polymer surface is useful for efficient drug encapsulation and subsequent drug release and hence improves therapeutic efficacy at the target site. Rajasekaran and coworkers⁴⁹ revealed that rough polymer surfaces provide a biomimetic environment by showing resemblance with the extracellular matrix that supported wound healing by supporting the cell regeneration process. Munz and coworkers⁵⁰ also found that hydrogels with rough topography provide better adhesion of the polymer matrix to the mucosal membrane and also improved drug release at the target site.

3.1.4 FTIR. FTIR spectra of CG revealed characteristic bands between 3251 and 3453 cm⁻¹ (ascribed to –OH stretching (str.), at 2961 cm⁻¹ (due to C–H str.), at 1269 cm⁻¹ (due to sulphate ester (O=S=O) symmetric str. of S=O), at 1150 cm⁻¹ (C–O str.) and at 1024 cm⁻¹ C–O–C str. of the glycosidic band present in CG. Characteristic bands of axial sulfate ester (C–O–SO₃) were observed at 844 cm⁻¹ and 806 cm⁻¹ str. vibrations of O-4 of G4S and O-2 of DA2S units, respectively⁵¹ (Fig. 2). The FTIR spectrum of CG-cl-poly(MEDSAH)-poly(AAc) demonstrated characteristic bands at 3260 cm⁻¹ [due to the –OH stretching mode of hydroxyl groups present in CG, poly(MEDSAH) & poly(AAc)], 2923 cm⁻¹ [due to –CH₂ stretching of methyl and methylene groups], peaks at 1717 cm⁻¹ & 1616 cm⁻¹ [due to C=O stretching of carboxylic groups present in poly(MEDSAH, poly(AAc)], 1545 cm⁻¹ [due to the quaternary ammonium (–N⁺R₃) group of poly(MEDSAH)], 1384 cm⁻¹ [due to O–H bending vibration], 1242 cm⁻¹ [due to C–N stretching of poly(MEDSAH)], 1182 cm⁻¹ [asymmetric stretching of S=O], and 1034 cm⁻¹ [symmetric stretching of S=O of poly(MEDSAH)], while the bands at 1034 and 987 cm⁻¹ are characteristic of the galactopyranose ring/glycosidic linkage and C–O of 3,6-anhydrogalactose, respectively. Ajmal and coworkers⁵² also observed the appearance of absorption bands corresponding to different functional groups of poly(MEDSAH) at 1720 cm⁻¹, 1482 cm⁻¹, 1175 cm⁻¹ and 1035 cm⁻¹ which confirmed the presence of the –C=O group, quaternary ammonium (–N⁺R₃) group, and asymmetric and symmetric stretching of the sulfonate group, respectively. This confirmed successful polymerization of poly(MEDSAH) onto hydrogels. Bashir and coworkers⁵³ revealed that FTIR spectra of karaya gum–poly(AAc) showed characteristic peaks of poly(AAc) at 2983 cm⁻¹ and at 1702 cm⁻¹ for the –CH₂ stretching vibration and –COOH group, along with peaks of gum, confirming successful grafting of poly(AAc) onto the polysaccharide. These FTIR spectral results supported the effective grafting & chemical interaction among polymer components, validating the formation of a crosslinked hydrogel network by chemically induced copolymerization reactions.

Fig. 2 FTIR of (a) CG and (b) CG-cl-poly(MEDSAH)-poly(AAc) polymers.

3.1.5 ¹³C NMR¹¹. ¹³C NMR of CG revealed peaks around 180 ppm because of the C=O carbon of uronic acid CG, and at 104 ppm (anomeric C1), 69 ppm (C2), 76 ppm (C3), 72 ppm (C4), 74 ppm (C5) and 61 ppm (C6) because of the G4S residue of CG (Fig. 3). The spectra also illustrated peaks at 94 ppm (C1), 74 ppm (C2), 76 ppm (C3 and C5), 78 ppm (C4) and 69 ppm (C6) due to DA2S unit residues of CG.⁵⁴ The ¹³C NMR spectrum of the CG-cl-poly(MEDSAH)-poly(AAc) hydrogel showed a peak at177 ppm due to carbonyl carbon C=O of poly(MEDSAH) &poly (AAc). Various peaks appearing between 67 and 103 ppm were associated with different carbon atoms within DA2S and G4S subunits of polysaccharide. The peak observed at 67 ppm corresponded to –CH₂–C̲H–SO₃⁻ of poly(MEDSAH), at 60 ppm corresponded to the –C̲H₂ group attached to the quaternary ammonium group of poly(MEDSAH), and at 50 ppm represented the carbon atom attached to oxygen (–O–C̲H₂–CH₂) of poly(MEDSAH); the peak at 40 ppm & 35 ppm was attributed to (–CH₂–C̲HX–) & (–C̲H₂–CHX) of poly(AAc) and poly(MEDSAH), respectively.^55,56 Furthermore, no peak for the unsaturated bond of –C=C was found in the range of 133 ppm–133 ppm, which represented the conversion of the vinyl monomer onto the polymer via the cross polymerization reaction.⁵⁷

Fig. 3 ¹³C NMR of (a) CG and (b) CG-cl-poly(MEDSAH)-poly(AAc) polymers.

Laishevkina and coworkers⁵⁸ also revealed that ¹³C NMR of poly(sulfobetaine methacrylate) based hydrogels showed a peak at 63 ppm which is related to carbon attached to the –SO₃⁻ group, a peak at 56 ppm which represents the carbon atom attached to oxygen, and a peak at 41 ppm which relates to the middle carbon of poly(sulfobetaine methacrylate). Das and coworkers⁵⁹ revealed that ¹³C NMR of dextrin–poly(AAc) based hydrogels showed a peak at 171 ppm due to carboxylic acid groups, and peaks at 41 ppm & 35 ppm were assigned to the sp³ hybridized carbon atom of poly(AAc). These peaks were formed through the polymerization reaction of poly(AAc). Inclusion of poly(MEDSAH) & poly(AAc) onto CG with the appearance of distinct peaks at 40 ppm and 35 ppm confirmed the crosslinking of the synthesized product by the chemically induced cross-polymerization reaction.

3.1.6 XRD. XRD spectra of CG showed sharp peaks at 2θ of 9.44°, 11.38°, 17.84°, 23.54°, 28.19°, 29.91°, 31.2° and 40.5° (Fig. 4). The presence of prominent & sharp peaks in the XRD spectrum is a result of inter/intramolecular hydrogen bonding interactions among various functional groups within polysaccharide units, contributing significantly to their crystalline framework. The strong, directional nature provided by these hydrogen bonds allowed saccharide units to align in a well-ordered & specific arrangement, along with a repeating pattern, hence giving a crystalline appearance to polysaccharide.²¹

Fig. 4 XRD of (a) CG and (b) CG-cl-poly(MEDSAH)-poly(AAc) polymers.

XRD spectra of the CG-poly(MEDSAH)-poly(AAc) hydrogel exhibited broad peaks, indicating an amorphous architecture of the polymer matrix (Fig. 4). This transition from crystalline to amorphous architecture was due to the interaction of amorphous components [poly(MEDSAH) & poly(AAc)] with polysaccharide that weakens the hydrogen bonds and disrupts the ordered arrangement & regular packing of polymer chains within the hydrogel network.⁵³ Jin and coworkers⁶⁰ also revealed that the addition of poly(MEDSAH) onto alginate showed a decrease in crystallinity owing to alteration in various intermolecular interactions.

Wang and coworkers⁶¹ revealed that polymerization of poly(AAc) onto chitosan led to a decrease in crystallinity and formed an amorphous material by decreasing inter/intramolecular hydrogen bonding interactions between polysaccharide units. Lazaridou and coworkers⁶² found that the addition of poly(MEDSAH) onto chitosan reduced their folding ability and hindered the formation of crystalline structure. Hence, the XRD spectrum with broadened peaks is indicative of successful incorporation of poly(MEDSAH) & poly(AAc) onto CG polysaccharide via grafting or crosslinking reaction.

3.2 Physiochemical and biomedical properties

3.2.1 Swelling analysis. Swelling characteristics of hydrogels were examined to study DD potential under physiological conditions. The influence of monomers and crosslinker content on network density was determined by assessing the swelling ability of hydrogels (Fig. 5 and Table 1). With a rise in [MEDSAH] content from 3.57 × 10⁻² mol L⁻¹ to 17.89 × 10⁻² mol L⁻¹, swelling decreases due to the dominance of crosslinking density over the hydrophilic character of the material. Also, the increase in [MEDSAH] content led to an increase in the number of polymeric chains in the hydrogel which makes the network more compact and restrains the mobility of macromolecular chains and hence limits the amount of solvent that can be absorbed.⁶³ With a rise in AAc content from 4.36 × 10⁻¹ mol L⁻¹ to 13.08 × 10⁻¹ mol L⁻¹, an increment in swelling index was attributed to the enhanced hydrophilicity of polymer chains.²¹ More hydrophilic sites are available for water molecules to interact with, hence allowing for greater water absorption and swelling due to more expansion and stability of porous polymeric networks. The NN-MBA crosslinker was added to the reaction mixture from 6.4 × 10⁻³ mol L⁻¹ to 32.43 × 10⁻³ mol L⁻¹, and it created covalent bonds between different polymer chains, forming a network structure. A higher crosslinker content forms a denser network grid within the polymer and hindered the penetration of solvent molecules onto the gel structure which results in limited swelling capacity.⁶⁴

Fig. 5 Effect of (a) [MEDSAH], (b) [AAc], (c) [NN-MBA] and (d) pH of swelling medium on CG-cl-poly(MEDSAH)-poly(AAc) hydrogels at 37 °C and (e) drug release profile of citicoline from drug loaded CG-cl-poly(MEDSAH)-poly(AAc) hydrogels in different media at 37 °C.

Table 1 Results of the swelling, gel strength, diffusion exponent n, gel characteristic constant k and various diffusion coefficients for the swelling kinetics of CG-cl-poly(MEDSAH)-poly(AAc) hydrogels

S. no.	Parameters	Swelling after 24 h (g g^−1 of gel)	Gel strength (Nmm)	Diffusion exponent n	Gel characteristic constant k × 10^2	Diffusion coefficients (cm^2 min^−1)
						Initial Di × 10^4	Average DA × 10^4	Late time DL × 10^4
Effect of MEDSAH × 10^2 (mol L^−1)
1.	3.57	11.72 ± 0.45	0.86 ± 0.07	0.67	1.20	1.06	1.02	0.91
2.	7.15	9.45 ± 0.33	1.84 ± 0.02	0.60	1.98	1.14	1.24	1.06
3.	10.73	9.04 ± 0.11	2.69 ± 0.09	0.64	1.50	1.23	1.23	1.08
4.	14.31	8.74 ± 0.49	4.26 ± 0.28	0.62	1.67	1.03	1.08	0.93
5.	17.89	8.33 ± 0.26	4.84 ± 0.23	0.60	1.87	0.95	1.05	0.88

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Effect of [AAc] × 10 (mol L^−1)
6.	4.36	10.59 ± 0.16	1.30 ± 0.08	0.76	0.74	1.48	1.07	1.13
7.	7.27	11.72 ± 0.20	0.86 ± 0.07	0.67	1.20	1.06	1.02	0.90
8.	10.18	12.76 ± 1.30	1.23 ± 0.03	0.61	1.56	0.95	1.02	0.88
9.	13.08	13.24 ± 0.06	1.06 ± 0.09	0.58	1.71	0.73	0.87	0.71

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Effect of NN-MBA × 10^3 (mol L^−1)
10.	6.40	18.35 ± 1.17	0.45 ± 0.01	0.72	0.84	1.10	1.12	0.89
11.	12.90	14.86 ± 1.00	1.00 ± 0.07	0.66	1.11	0.95	1.13	0.81
12.	19.45	13.52 ± 0.34	1.08 ± 0.07	0.67	1.10	1.04	1.21	0.88
13.	25.94	12.98 ± 0.77	1.16 ± 0.04	0.61	1.55	0.85	1.21	0.80
14.	32.43	12.76 ± 1.30	1.23 ± 0.03	0.61	1.56	0.95	1.28	0.88

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Effect of swelling medium
15.	pH 2.2 buffer	2.57 ± 0.73	14.67 ± 0.29	0.58	1.86	0.95	1.43	0.91
16.	Distilled water	14.86 ± 1.00	1.00 ± 0.07	0.66	1.11	0.95	1.12	0.84
17.	pH 7.4 buffer	8.81 ± 0.14	10.37 ± 0.21	0.71	0.77	0.82	0.94	0.73
18.	ACSF buffer	9.14 ± 0.38	3.69 ± 0.24	0.67	0.99	0.77	0.98	0.71

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CG and poly(AAc) also imparted pH sensitivity to the hydrogel and this characteristic is used to deliver drugs into specific areas of the body including the brain. Hydrogels for BDD are typically designed with ionisable functional groups that undergo protonation/deprotonation in response to pH changes. More swelling was observed at higher pH buffer medium mimicking brain fluid. Deprotonation of –SO₃H and –COOH groups occurred at higher pH values and these partially ionized –SO₃⁻ & –COO⁻ groups undergo repulsion between different polymeric chains, resulting in greater swelling of CG-cl-poly(MEDSAH)-poly(AAc) hydrogels.⁶⁵ The higher swelling capacity of hydrogels is beneficial as it increases drug diffusion and hence is suitable for drug release at the target site of nerve injury.

Hydrogels designed for BDD should be pH sensitive, thereby releasing the drug specifically at the site of nerve injury.⁶⁶ This targeted DD enhanced the therapeutic effect by minimizing damage to healthy brain cells and maximizing drug concentration at the site of injury.

3.2.2 Drug release properties. The drug release behaviour of citicoline from drug encapsulated hydrogels was determined in pH 2.2 buffer, distilled water, pH 7.4 buffer & ACSF buffer media. Higher drug release was obtained for the ACSF buffer medium, mimicking brain fluid, which can be explained on the basis of pH influence on the structure of the hydrogel (Table 2 and Fig. 5). At higher pH, the pore size increases due to expansion of the polymer network that was driven by ionization of functional groups (–SO₃H and –COOH) present in the hydrogel.^67,68

Table 2 Results of diffusion exponent n, gel characteristics constant k, various diffusion coefficients and correlation coefficients (R²) of different drug release models for the release profile of citicoline drug loaded CG-cl-poly(MEDSAH)-poly(AAc) hydrogels

		pH 2.2 buffer	Distilled water	pH 7.4 buffer	ACSF buffer
Diffusion exponent n		0.70	0.81	0.83	0.a81
Gel characteristics constant k × 10^2		1.33	0.69	0.60	0.70
Diffusion coefficients (cm^2 min^−1): initial Di × 10^4		1.65	2.30	2.45	2.34
Average DA × 10^4		0.91	1.05	1.06	1.05
Late time DL × 10^4		1.44	1.77	1.89	1.83
Zero order	R ^2	0.969	0.975	0.983	0.978
	K o × 10^3 (min^−1)	1.94	2.00	2.07	2.05
First order	R ^2	0.996	0.996	0.990	0.993
	K 1 × 10^3 (min^−1)	4.70	4.40	4.70	4.70
Higuchi	R ^2	0.997	0.998	0.999	0.998
	K H × 10^2 (min^−1/2)	5.28	5.44	5.61	5.57
Korsmeyer–Peppas	R ^2	0.993	0.992	0.994	0.993
	K KP × 10^1 (min^−n)	7.07	8.12	8.38	8.12
Hixson–Crowell	R ^2	0.998	0.999	0.999	0.999
	K HC × 10^3 (min^−1/3)	1.16	1.12	1.18	1.18

---

The pH sensitivity of the polymer matrix facilitated fluid intake, and hence more drugs can diffuse out of the hydrogel matrix easily. Release of a pharmaceutical compound through a targeted delivery system at the site of nerve injury can be achieved by the strategic design of pH-sensitive DD carriers to maintain the consistent therapeutic concentration for a prolonged time period. This approach enables precise and controlled release of active agents or growth factors, ensuring enhanced therapeutic efficacy, while minimizing off-target effects, leading to improved treating outcomes.

Drug–polymer interactions also played an important role in facilitating slow and sustained drug release which is beneficial for neuron regeneration.⁶⁹ Various functional groups present in citicoline drug interacted with functional moieties of the polymer matrix which provides controlled DD at the target site. Drugs or bioactive agents when incorporated in these network structures can access the external medium by diffusion through the pores.⁷⁰ Further, research studies also demonstrated that the crosslinking density of materials influences the diffusion of the drug molecule from the polymer matrix. The sustained release profile from the hydrogel loaded DD carriers could be further tailored by changing the crosslinking density of network hydrogels by changing the content of monomer and crosslinker during synthesis of hydrogels.⁷¹ The change in biological activity of drugs and different materials has also been discussed in other research studies.^72–74 The DD carrier enhanced the anti-tumor efficacy of docetaxel trihydrate. They revealed that these drug loaded carriers have the unique property of incorporating the hydrophobic drug in their internal network and their smaller particle size also facilitated the controlled release of the drug on the target site.⁷²

Lin and coworkers⁷⁵ revealed that citicoline loaded liposomal scaffolds showed a decrease in the inflammatory response of microglia cells & promoted axon elongation thereby creating a favorable microenvironment for nerve repair. Bekar and coworkers⁷⁶ studied the application of citicoline in functional recovery and regeneration of surgically transacted sciatic nerves in rats. Cavalu and coworkers⁷⁷ reported that the anti-inflammatory action of citicoline was mainly due to its ability to stabilize cell membranes & reduce excessive release of pro-inflammatory neurotransmitters that promote neuro-inflammation.

Results illustrated sustained delivery of citicoline, wherein initial rapid drug release was followed by slower release for an extended time period. These observations were further supported by diffusion coefficient values wherein the initial diffusion coefficient value (2.34 × 10⁻⁴ cm² min⁻¹) was higher compared to the later stage diffusion coefficient value (1.83 × 10⁻⁴ cm² min⁻¹), indicating more controlled drug release over time. Values of n indicated a non-Fickian transport mechanism for the release of the drug from the polymeric structure in all media. Mathematical kinetic models expressed diffusion kinetics of drugs from the polymer matrix and their regression coefficient values for formulations illustrated that release was best followed by the Hixson–Crowell kinetic model.

Citicoline (cytidine-5′-diphosphocholine), a neuro-protective & neuro-regenerative agent, played a crucial role in brain repair & functional recovery following brain injury. Its mode of action primarily involved biosynthesis of phosphatidylcholine, an essential phospholipid in the neuronal membrane. Phosphatidylcholine is further required for restoring, strengthening and reserving the structural integrity of the damaged neuronal membrane.⁷⁶ Citicoline is believed to undergo hydrolysis & dephosphorylation to yield cytidine and choline. Choline is then used for the synthesis of phosphatidylcholine through the cytidine diphosphate choline (CDP) pathway. Additionally, it facilitates the synthesis of acetylcholine, a key neurotransmitter essential for memory and cognitive function of the body. Citicoline also protects brain cells by reducing neuro-inflammation, oxidative stress by suppressing lipid peroxidation and formation of harmful free radicals, thereby safeguarding neurons from injury.^78–81 A pictorial representation of the mode of action of citicoline during the nerve regeneration process occurring at the site of brain injury^78–81 is shown in Scheme 2.

Scheme 2 A pictorial representation of the mode of action of citicoline during the nerve regeneration process occurring at the site of brain injury.^75–78

3.2.3 Blood compatibility.
3.2.3.1 Haemolytic and thrombogenicity assay. Haemocompatibilty testing of biomaterials is an essential aspect regarding the evaluation of the interaction of medical devices with blood or therapeutic agents that are associated with biochemical pathways without causing destructive coagulation or unwanted inflammation and any cellular or humoral reactions in the host (Table 3). The haemolysis index test & thrombogenicity percentage inferred that the present polymeric sample exhibited a haemolysis capacity of 0.83 ± 0.08% and a thrombogenicity value of 83.96 ± 1.67% during interaction with blood and hence it can be classified as a hemocompatible material. The key factor associated with this compatibility was the hydrophilic nature of materials owing to the composition (CG, poly(AAc) and poly(MEDSAH)) of hydrogels. Hydrophilic features prevented blood cell adhesion and rupturing of their membrane by forming a hydrated layer that weakens the interaction with the erythrocytes and hydrogel surface. The blood compatibility properties of the polymer matrix relieved haemolysis and also suppressed thrombus or clot formation, hence minimizing the immune response of the human body and they could be considered as materials that can be applied as a carrier for BDD.⁸²

Table 3 Results of thrombogenicity, haemolytic potential, protein adsorption, cytotoxicity, mucoadhesion and antioxidant activity properties of CG-cl-poly(MEDSAH)-poly(AAc) hydrogels

Properties	Thrombogenicity	Inference
Thrombose percentage (%)	83.96 ± 1.67%	Non-thrombogenic nature
Haemolytic index (%)	0.83 ± 0.08%	Non-haemolytic nature
Protein adsorption (%)	3.13 ± 0.24%	Biocompatible nature
Cell viability (%)	165 ± 0.19%	Cyto-compatible nature

---

Ex-vivo mucoadhesion
Peak detachment force (mN)	75 ± 4.00 mN	Mucoadhesive nature
Work of adhesion (Nmm)	0.25 ± 0.02 Nmm

---

Antioxidant activity
F–C reagent assay	25.22 ± 0.41 μg GAE		Antioxidant nature
Phosphomolybdenum assay	497.76 ± 2.58% μg AAE
DPPH reagent assay	Time	% inhibition
	24 h	39.82 ± 1.65

---

CG has provided enhanced haemostatic functions to materials which induced negligible haemolysis and did not modify the morphology of erythrocytes.^83,84 The anticoagulant effect of CG was attributed to sulphated groups making them optimal choices for soft tissue DD.¹⁵ Zwitterionic polymer poly(MEDSAH) has not induced erythrocyte lysis and maintains structural integrity of RBCs during prolonged blood exposure to polymers.⁸⁵ Moreover, these zwitterionic hydrogels with hydrophilic properties demonstrated a decrease in the aggregation of RBCs.⁸⁶ The material exhibited haemostatic and non-thrombogenic properties, indicating its potential suitability as a carrier for delivering nerve regenerative agents without inducing inflammation or toxicity. Further, Samadian and coworkers⁸⁷ revealed that incorporating citicoline in gelatin based scaffolds increased haemocompatibilty with low RBC lysis. Hence, citicoline also contributed to providing blood compatibility features to BDD carriers.

3.2.3.2 Protein adsorption. Addressing biocompatibility properties of hydrogels could also be correlated with protein adsorption characteristic of materials, as these interactions influence cellular responses and overall biological performances. BSA adsorption studies revealed 3.13 ± 0.24% albumin adsorption after 24 hour incubation at 37 °C (Table 3). In this study, it was observed that the hydrogels exhibited minimal BSA protein adsorption onto their surface due to hydrophilic groups of CG-cl-poly(MEDSAH)-poly(AAc) hydrogels. These groups formed a hydration layer around the hydrogel that prevents protein–polymer interactions and hence inhibited protein molecules from approaching and adhering to the polymer surface. Additionally, negative charges on both the hydrogel surface and the BSA protein molecules induced electrostatic repulsion, further reducing protein adsorption.⁸⁸

Wang and coworkers⁸⁹ revealed that incorporation of zwitterionic poly(MEDSAH) onto hydrogels has significantly reduced protein adsorption. Polar poly(MESAH) created a hydration layer around the hydrogels, effectively repelling proteins and preventing their adhesion to the polymer surface. Tavakoli and coworkers⁹⁰ revealed that a rise in CG content in hydrogels decreased protein adsorption due to the presence of negatively charged sulfate groups that repel protein molecules approaching the polymer surface. Carneiro and coworkers⁹¹ revealed that the presence of sulfonate groups in CG resulted in less interaction with albumin and prevented protein adsorption on the surface of CG-hydrogels.

Designing a DD system that successfully overcomes constraints imposed by the BBB required excellent blood compatibility potential with minimum side effects. These hemocompatible DD systems ensure efficient and sustained transport of bioactive agents to the brain by preventing haemolysis, thrombosis, immune responses and unwanted protein adsorption. A pictorial representation of the reduction in protein adsorption by the hydrogel matrix is shown in Scheme 3.

Scheme 3 A pictorial representation of reduction in protein adsorption by the hydrogel matrix.

3.2.4 Cell viability assay. The cyto-compatibility potential of hydrogels was evaluated to ensure their viability on exposure to mammalian cells. Results of cell viability tests unveiled 165 ± 0.19% cell viability on exposure of RD cells to the hydrogel material (Table 3). These findings signified that the hydrogel sustained the viability and enabled proliferation of cells with minimal cytotoxicity, facilitating enhanced cellular attachment and promoting cell proliferation and hence marking their suitability in biomedical contexts. The non cyto-toxic nature of the polymer material can be ascribed to the composition of the hydrogel [CG, poly(MEDSAH) and poly(AAc)], ensuring optimal interaction with biological systems and minimizing adverse cellular responses.^92–95 Research studies elaborately demonstrated the importance of cyto-compatibility and cell viability properties for repairing injury sites. Results suggested that the hydrogel has no significant cytotoxicity and has biocompatibility and neuroprotection features which have provided a microenvironment to protect against neurological damage and maintain nerve cell survival and repair besides taking care of wound or nerve injury.⁹⁶

CG demonstrated higher cyto-compatibility in other research findings. The polysaccharide containing sulfonate groups incorporates a high density of negative charge to polymeric materials which exhibit close resemblance to natural glucosamine present in the ECM. Sulfated domains of CG mimicked ECM facilitated cell adhesion, proliferation and migration.⁹⁷ Pettinelli and coworkers⁹⁸ reported that the viability of fibroblast cells seeded on the CG-derived hydrogel presented higher metabolic activity. The CG based hydrogels exhibited more proliferation, adhesion and spreading of fibroblast cells. They linked these observations to the structural similarity of the material to the ECM that improved adhesion and spreading of cells which in turn enhance adhesion and proliferation of fibroblast cells. Kim and coworkers⁹⁹ demonstrated that proliferation of mesenchymal cells was enhanced on incorporation of CG in the alginate–CaSO₄ hydrogel. They found that after 24 hours the majority of cells were viable and negligible cells were found dead, implying the cyto-compatible and non-toxic nature of CG. Tytgat and coworkers¹⁰⁰ also reported that cells were found viable on interaction with CG-gelatin based hydrogel films owing to the ECM mimicking trait of the material that helps in more cell spreading and proliferation. CG has stimulated pre-osteoblast proliferation and osteogenic differentiation, which has been considered as a potential factor for the promotion of bone regeneration.^101–103 Ganguly and coworkers⁶⁸ revealed that poly(AAc) based hydrogels with porous structure are found to be non-toxic to human cells and more cell growth was observed in the porous surface of the hydrogel.

Hence, the cell viability test signified the biocompatibility parameter of the material. These hydrogels are beneficial for brain targeted DD with cyto-compatible functionality and may be considered as a promising platform for advanced biomedical applications, meeting the demands for both biocompatibility and functional performance. Hence, the present polymer matrix is safe for attaching mammalian cells which is confirmed from the cell viability test of RD mammalian cells.

3.2.5 Antioxidant activity. The AOX defence mechanism of a material is a fact of consideration while evaluating its potential for neurological applications (Table 3). AOX hydrogels can help in mitigating oxidative damage and provide neuro-protection while delivering therapeutic agents to the brain. Results of AOX assays demonstrated 39.82 ± 1.65% DPPH free radical inhibition, F–C reagent assay indicated 25.22 ± 0.41 μg of GAE of AOX activity and PM reagent assay revealed an AOX activity of 497.76 ± 2.58% μg equivalent of ascorbic acid (AAE). CG based functional hydrogels act as radical quenchers due to their innate ability to neutralize toxic radicals.¹⁰⁴

The AOX potential of the material prevents oxidative damage to neuronal cells, suppresses neuro-inflammation and promotes tissue repair.¹⁰⁵ AOX compounds favour nerve repair by providing oxidative protection from ROS, preventing lipid peroxidation, and limiting cellular damage & the inflammatory process and hence favour the repair of damaged tissue.¹⁰⁶ Recently, the role of tissue repair by various agents and materials used at the brain injury or wound site has been recognized.^107,108 Biomaterials have the ability to intervene in immune responses and inflammation, triggered during neural dysfunction and facilitating tissue repair and nerve regeneration.¹¹⁰ Key strategies including attenuating ROS levels, maintaining the local environment and promoting axonal growth are promising ways to mitigate acute response.¹⁰⁹ Hence, by scavenging ROS, they protect damaged neurons and facilitate growth & nerve repair, promoting functional recovery after nerve damage.

Overall, the AOX activity of the present DD carriers may further reduce oxidative stress at the injury site to some extent. During brain tissue repair, the high rate of oxidative stress owing to high metabolic activity decreases the repair process of damaged tissue. Here, it may be said that bioactive hydrogels execute biomimetic action like the antioxidant action of enzymes (glutathione peroxidase), vitamins (E) and hormones (melatonin) under the self-repair tissue regeneration process. These hydrogel dressings may contribute to reducing oxidative damage and promoting protective tissue at the brain injury site. Further, the synergic action of sustained release of citicoline and the antioxidant action of BDD carriers may further facilitate the repair process.

Further, the synergistic AOX activity of materials not only enhanced the radical scavenging action but also supported neuronal regeneration in other research studies.^110,111 Antioxidants disrupted pro-inflammatory cascades and stabilized redox-sensitive pro-survival pathways. These combined molecular effects help in preserving the neurotrophic signaling pathway and create a microenvironment that directly couples antioxidant function with the enhanced neuronal regeneration. Zhou and coworkers¹¹⁰ extensively elaborated the repair process of spinal cord injury with an antioxidant and neural regenerative nanoplatform. These nanoparticles scavenge free radicals and prevent damage to neurons at the site of injury. A pictorial representation of the AOX activity of the hydrogel matrix is shown in Scheme 4.

Scheme 4 A pictorial representation of the AOX potential of the hydrogel.

Various reports also indicated that the type of neurological damage and delivery methods are deciding factors for AOX activity needed for neuro-protection. It was observed that no single value has been suggested ideal of antioxidants for neuro-protection. Different antioxidant agents demonstrated effectiveness in preclinical studies to reduce oxidative stress and cell death. But it remained a challenge to develop effective drug delivery carriers for the CNS and standardize a clinical protocol to translate this finding into promising therapy.¹¹² The different methods (DPPH, FC & and phosphomolybdate) demonstrate different values of AOX activities because different assays operate on different analytical scales and mechanisms. However, the role of DD carriers in establishing a consistent protocol in clinical trials for validation of neuro-protective potential could be another point of research needed in this direction.¹¹³

3.2.6 Mucoadhesion. The adhesive strength of the material with the bio-membrane was evaluated in terms of detachment force and work of adhesion. The hydrogel required a peak force of 75 ± 4.00 mN and the associated work of adhesion was 0.25 ± 0.02 Nmm, desirable for formulations to adhere to mucosal tissue (Table 3). Results revealed the mucoadhesive properties of this CG based hydrogel. Bioadhesive polymers with specific hydrophilic functionality (–COOH, –OH, –OSO₃⁻) in CG-cl-poly(MEDSAH)-poly(AAc) contribute significantly to adhesion. These groups interact with various functional groups present in the bio membrane surface by covalent and non-covalent interactions, ensuring strong bonding, allowing gradual drug release & improving long-term drug retention at the target site. The hydrophilic nature of materials has provided better cell attachment and alignment during nerve regeneration.^114,115 Poly(MEDSAH) has exhibited notable adhesive properties due to their zwitterionic nature, allowing strong interactions with bio surfaces in the literature.⁵² Poly(AAc) has also boosted mucoadhesion owing to carboxylic functional groups that adhere strongly to the bio membrane, required for efficient retention of the drug at the target site.¹¹⁶ Meng and coworkers¹¹⁷ developed poly(lipoic acid)–poly(dopamine) based adhesive hydrogels for the repair of spinal cord injury. These gels exhibited adhesive properties owing to synergistic interactions between poly(lipoic acid) & polyphenols.

Overall, the mucoadhesive properties of CG-cl-poly(MEDSAH)-poly(AAc) hydrogels make them suitable for designing DD vehicles and hence can assist DD carriers in targeting specific brain tissue. A hypothetical pictorial representation of mucosal adhesion of the hydrogel in DD and the nerve regeneration process is given in Scheme 5.

Scheme 5 A hypothetic pictorial representation of mucosal adhesion of the hydrogel in drug delivery and the nerve regeneration process.

3.2.7 Gel strength. The key factor determining gel strength and swelling behaviour of hydrogels is the degree of crosslinking within the polymer matrix. There is an inverse relationship between gel strength & the swelling characteristic of the material (Table 3). With a rise in [MEDSAH] content from 3.57 × 10⁻² mol L⁻¹ to 17.89 × 10⁻² mol L⁻¹, gel strength increased probably due to higher crosslinking, indicating that the samples have a more compact structure. With a further increase in AAc content from 4.36 × 10⁻¹ mol L⁻¹ to 13.08 × 10⁻¹ mol L⁻¹, a decrease in gel strength was observed owing to weak gel network formation that allowed more water uptake and increased swelling and subsequently decreased gel strength. An increase in NN-MBA content from 6.4 × 10⁻³ mol L⁻¹ to 32.43 × 10⁻³ mol L⁻¹ improved the mechanical strength of the material significantly. Higher crosslinking led to a stronger gel structure & limited available space for absorption of solvent molecules. Moreover, with a rise in pH, a decrease in gel strength was observed, attributed to higher swelling under alkaline conditions, resulting in weaker gel consistency. A designed network polymer matrix with covalent and non-covalent interactions can provide sufficient strength as a platform for nerve regeneration and ensure controlled drug release at the target site while maintaining its structural integrity & stability under physiological conditions. Swollen materials act like soft brain tissue and release of the encapsulated brain repairing drug from hydrogels can enhance neural regeneration, encouraging cell attachment, proliferation, and differentiation, supporting the formation of new synapses and hence the released drug is crucial for providing neuroprotective and therapeutic effects.¹¹⁸

The combination of sulfated polysaccharides & zwitterionic polymers has led to the development of hydrogels exhibiting some superior properties relative to other materials. The CG hydrogels demonstrated superior stability and maintained structural integrity over a broad range of physiological pH relative to chitosan, hyaluronic acid-based systems.^119,120 The sulfated polysaccharide showed higher biomimetic behavior, mimicking the ECM, and promoted cell adhesion, migration and proliferation compared to chitosan hydrogels.¹²¹ CG is a negatively charged sulfonated material which imparts superior haemocompatibilty by repelling negatively charged proteins and platelets thereby reducing haemolysis while chitosan containing positive amino groups may attract the negatively charged RBC membrane and reduce the blood compatibility of chitosan.¹²²

CG can trigger immune response in the body by acting as a pro-inflammatory stimulus by interacting with the receptors and activate innate immune pathways, leading to release of pro-inflammatory cytokines, and hence can contribute to inflammation in the body.¹²³ However, it was observed that after hydrogel formation, it can modulate immune responses and promote the immune system by controlling inflammation by capturing and neutralizing inflammatory cytokines, preventing migration of macrophages and promoting tissue regeneration. Three dimensional hydrogel networks allowed controlled release of anti-inflammatory drugs or therapeutic agents that also reduced system side effects. Moreover, the network structure and hydrophilic nature of the hydrogel closely resemble the ECM that allows growth and functioning of cells and helps in suppressing immune responses.¹²⁴ However, in some cases it was observed that CG also suppresses immune responses by the cytopathic effect on macrophages. The immunosuppressive effect of CG was due to its ability to damage the cells of mononuclear phagocyte systems.^125,126 Yermak and coworkers¹²⁷ reported that the inhibitory effect of CG on lipopolysaccharide induced inflammation was attributed to the immunomodulation effect of CG wherein it stimulates the biosynthesis of various mediators of the immune system and modulates the cellular activity of peritoneal leukocytes.

4 Future in vivo plans and scalability

The present drug encapsulating hydrogels could be applied for brain injury as wound dressings. However, further in vivo research is needed for their practical applicability and marketable products. The in vivo research involves (i) a brain injury model [selection of mouse species & creation of brain injury], (ii) dressing applications [selection of dressings along with negative/positive control and application of dressing with a consistent method], (iii) in vivo monitoring and assessment [wound healing, histological analysis and qualitative and quantitative data analysis] and (iv) ethical approval and regulatory compliance before clinical trials.¹²⁸

One of the most difficult clinical problems in the world is the regeneration of nerve defects/damage brought on by nerve injuries. One of the potential strategies for repairing damaged neural tissues is the creation of 3D replicas of nerves that closely match the natural ECM.¹²⁹ Bramhe and coworkers¹³⁰ elaborately discussed this model along with other tissue engineering aspects. The neural model should generally satisfy a number of requirements, including electro conductivity and elastic properties/micro-architecture to mimic various properties including mechanical & physicochemical properties of native nervous tissues, ECM and neuro compatibility to facilitate nerve cell attachment and proliferation. Further it is pertinent to mention that immunological rejection is also limiting the use of tissue models in clinical practice.

CG and acrylic monomers are cost effective materials with high scalability for synthesis of hydrogels at a large scale. The sterilization of the hydrogel material could be conducted using various techniques. Earlier sterilization of the hydrogel material has been performed with heat, radiation, and a gas sterilization method. However, recently some new methods of sterilization of these materials have emerged including ozone & supercritical carbon dioxide sterilization techniques without modification in mechanical aspects of materials.¹³¹ Thus, scalability requires an optimized synthesis protocol, cost effective raw material selection and validated standardization processes to ensure lab test transition onto industrially viable hydrogel devices for commercial viability.¹³²

5 Conclusions

The present analysis established the relationship between physicochemical properties and synthetic reaction parameters which has been established from cerebrospinal fluid sorption characteristics of network materials. FESEM images revealed irregular, heterogeneous morphology. Rough topography was unveiled from AFM images. A rough surface is beneficial as it provides more surface mucosal and tissue adhesion. FTIR & ¹³C NMR spectra confirmed incorporation of poly(MEDSAH) & poly(AAc) onto CG via a chemically induced cross polymerization reaction. XRD reflected their amorphous attributes during polymeric chain arrangement in the hydrogel matrix. Hydrogels expressed 165 ± 0.19% cell viability of RD cells and promoted cell adhesion & proliferation, signifying their compatibility with mammalian cells. DPPH assay revealed 39.82 ± 1.65% free radical scavenging, highlighting their strong intrinsic AOX potential for neutralizing oxidative stress at the site of nerve injuries. The bioadhesive ability of materials was signified from a force of 75 ± 4.00 mN, desirable for adherence to the mucosal surface and helps in providing cell attachment and alignment during the nerve regeneration process. The citicoline anchored brain DD carrier released the drug in simulated brain fluid in a sustained pattern and followed the non-Fickian diffusion mechanism. The release profile was best explained by the Hixson–Crowell kinetic model. Drug anchored hydrogels offered a functional platform that synergistically integrates biochemical, biophysical and pharmacological attributes to facilitate nerve regeneration. Herein, a sustained release from bioactive hydrogels may execute biomimetic action like the antioxidant action of enzymes, vitamins and hormones under the self-repair tissue regeneration process. Overall, the present synergistic therapy for treatment of brain injury involved the delivery of the bioactive nerve regenerating agent from functional materials. It will not only deliver therapeutic molecules to nerve injuries but their inherent antioxidant, haemostatic & non-cytotoxic nature with cell viability properties may also contribute to enhancing the nerve repairing process of brain injury.

Conflicts of interest

There are no conflicts (financial and research) of interest to declare.

Data availability

All the data supporting this article have been included in the manuscript in the form of tables/figures and no other research results, software or code have been included.

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