Injectable gelatin–PEG hydrogels obtained via cytochrome C-mediated polymerization

Abstract

Injectable hydrogels capable of in situ gelation under physiological conditions are highly attractive for minimally invasive surgery and locoregional drug delivery. We herein report three novel injectable hydrogels composed of gelatin methacrylate (GelMA) and poly(ethyleneglycol) dimethacrylate (PEGDA), crosslinked through a Fenton-like radical polymerization mediated by cytochrome C (CyC) in the presence of H₂O₂ and L-ascorbic acid. To the best of our knowledge, this is the first example of an injectable hydrogel formulation in which CyC is used as a redox mediator for radical polymerization. CyC enables the replacement of transition metals while maintaining polymerization kinetics comparable to those of hydrogels synthesized via traditional Fenton systems. The resulting hydrogels, undergoing a sol–gel transition within 1.0 and 2.0 minutes, are biocompatible, and their properties are highly tunable. Indeed, rheological analysis showed that mechanical properties and the linear viscoelastic region (LVR) can be easily modulated by varying the concentrations of the starting methacryl-functionalized gelatin and the crosslinker. Structural characterization and biodegradation studies revealed that enzymatic degradation is strongly dependent on the degree of crosslinking. All hydrogels were readily injectable and showed no detectable cytotoxicity in conditioned-medium assays. Sustained release of rhodamine 101, as a drug-mimicking system, reached ∼70% over 7 days.

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Introduction

Hydrogels are three-dimensional (3D) networks composed of cross-linked hydrophilic polymer chains capable of absorbing and retaining large amounts of water or biological fluids.¹

The chemical and physical properties of hydrogels can be finely tuned through the appropriate choice of building blocks/crosslinkers, making these materials versatile and useful for applications in tissue engineering and healing,² biosensing/imaging³ and drug delivery systems.⁴ Among these materials, injectable hydrogels have come to the fore as particularly promising platforms for drug delivery applications.⁵ They are able to encapsulate therapeutic agents within their three-dimensional network, enabling controlled and sustained release.⁶ A major advantage of injectable hydrogels is their minimally invasive administration, as they can be delivered via injection without requiring a preformed scaffold.⁷ Moreover, being liquids at the injection point, they can fill small spaces and adapt to the shape and volume available. Finally, their intrinsic biocompatibility and porous architecture facilitate oxygen and nutrient diffusion, thereby supporting cell viability.⁸ Injectable hydrogels can be classified according to three main criteria: (i) polymer source (natural or synthetic), (ii) structural organization (homo- or copolymeric), and (iii) crosslinking mechanism.

The main crosslinking strategies include ionic interactions⁹ and covalent bond formation (radical polymerization,¹⁰ click chemistry,¹¹ Michael addition,¹² Schiff base formation,^5a and enzymatic crosslinking¹³). Other approaches rely on hydrophobic interactions, host–guest complexation,¹⁴ hydrogen bonding, and stereo-complexation.¹⁵

Among these, radical polymerization represents one of the most effective strategies due to its short reaction time, which gives gelation within minutes.⁷ A common method to trigger radical polymerization involves the use of light irradiation;¹⁶ however, this approach raises concerns regarding cytotoxicity, site reachability, and patient safety. To address these limitations, redox initiators can be employed to trigger radical formation.¹⁷

One of the most widely used approaches to initiate polymerization via a redox mechanism is the Fenton reaction (FR), capable of generating hydroxyl radicals (˙OH) from hydrogen peroxide in the presence of iron ions.¹⁸

The FR has been widely applied to initiate radical polymerization of acrylate- and methacrylate-based polymers and hydrogels,¹⁹ and it has found wide applicability in biomedical²⁰ and environmental applications.²¹ Although several hydrogels obtained by conventional FR systems have been described,²² the use of a biological iron source could be more attractive for in vivo applications. We herein report three novel injectable gelatin-based hydrogels obtained via radical polymerization triggered by a redox system catalyzed by cytochrome C (CyC), a low amount of the oxidant H₂O₂ and the water-soluble reducing agent L-ascorbic acid (Fig. 1). Importantly, the use of CyC provides a heme protein-based alternative to free Fe species, which can chelate with proteins/enzymes and alter their biological functions.²³ Control and scavenging experiments supported the role of CyC as a redox mediator, suggesting that radical generation is associated with the redox activity of the heme center. Hydrogel formation occurs in 1–2 minutes, depending on the formulation. The gelation times are comparable to those of reported injectable hydrogels²⁴ or composite hydrogels²⁵ obtained through conventional Fenton systems, while the storage moduli (G′; 10³–10⁴ Pa) are comparable to those of reported gelatin-based injectable hydrogels.²⁶

Fig. 1 (A) Schematic illustration of the preparation of injectable hydrogels via Fenton polymerization. Loading and release of a drug-mimicking (Rhodamine 101) system. (B) Pictures of the three hydrogels obtained: (I) Gel-PEG A; (II) Gel-PEG B; and (III) Gel-PEG C.

Beyond rheological properties, the biodegradability, cytocompatibility, and release characteristics of our Gel-PEG hydrogels highlight the potential of such materials for locoregional drug delivery over several days.

Results and discussion

Hydrogel synthesis

Injectable Gel-PEG hydrogels (Fig. 1A) were obtained via radical polymerization initiated by a Fenton-like reaction mediated by cytochrome C (CyC).

To ensure biocompatibility, we selected gelatin (gelatin from porcine skin, type A) as the polymeric building block; however, pristine gelatin is not reactive in radical polymerization, and thus it was functionalized with methacryl groups.

We decided to keep the methacrylation degree low to preserve a fraction of positively charged amino groups which are essential for promoting tissue and cell adhesion; thus, gelatin methacrylation was conducted under slightly acidic²⁷ conditions (pH = 5.5, MES buffer; Fig. S1). ¹H NMR (Fig. S1) and IR (Fig. S2) analyses confirmed the presence of methacrylate/methacrylamide groups in the GelMA polymer.

The total amount of methacryl groups (MA) per mg of gelatin was estimated to be 2.6 × 10⁻⁴ mmol_MA mg_GelMA⁻¹ (see the SI). In line with NMR, the Kaiser test conducted pre- and post-functionalization of the gelatin revealed slight variation in the number of amino groups (Fig. S3).

The methacryl-functionalized GelMA polymer so obtained was therefore employed in the synthesis of injectable Gel-PEG hydrogels, using the commercially available PEGDA (M_n = 700) as the crosslinker. Radical polymerization of GelMA and PEGDA was performed by using H₂O₂ as the oxidant and CyC as the iron catalytic source. As reported in the literature,²³ the heme iron center of CyC can react with H₂O₂, leading to the formation of reactive intermediates that may generate hydroxyl radicals. L-Ascorbic acid (AA) was employed as a sacrificial reducing agent, since ascorbate (AscH⁻) and the ascorbate radical anion (Asc˙⁻) can reduce Fe_CyC³⁺ back to Fe_CyC²⁺, restoring the catalytic center while producing a dehydroascorbate (DHA) molecule.²⁸ Based on literature reports of Fenton systems involving chelated iron species,²³ we proposed a tentative reaction mechanism as summarized in eqn (1)–(4):

Fe_CyC²⁺ + H₂O₂ → Fe_Cyc(H₂O₂)²⁺ (1)

Fe_Cyc(H₂O₂)²⁺ → Fe_CyC³⁺ + ˙OH + OH⁻ (2)

Fe_CyC³⁺ + AscH⁻ → Fe_CyC²⁺ + Asc˙⁻ + H⁺ (3)

Fe_CyC³⁺ + Asc˙⁻ → Fe_CyC²⁺ + DHA (4)

To the best of our knowledge, this represents the first example of an artificial hydrogel synthesized by Fenton-like radical polymerization using CyC in place of iron salts. With this procedure (see Fig. 1), three injectable hydrogel formulations – Gel-PEG A (GelMa 0.47% m/v + PEGDA 8.0% m/v; Fig. 1B), Gel-PEG B (GelMa 0.53% m/v + PEGDA 8.9% m/v; Fig. 1B, II), and Gel-PEG C (GelMa 0.58% m/v + PEGDA 9.8% m/v; Fig. 1B, III) – were prepared by simultaneously increasing the amount of both gelatin and the crosslinker.

The hydrogels exhibited very high and comparable swelling capacities, with swelling ratios (SR%) of 90%, 89%, and 88% for Gel-PEG A, Gel-PEG B, and Gel-PEG C, respectively.

Extrusion tests demonstrated that all the synthesized materials could be smoothly extruded in PBS through a syringe equipped with an insulin needle, while maintaining structural integrity and without needle occlusion (see Movie S1). This property is critical for clinically viable injectable systems and indicates that the polymer ratios in the formulations do not compromise flowability.

The Gel-PEG B formulation (Fig. S4) was used as a model system to qualitatively assess the role of CyC in the polymerization mechanism.

Control experiments were performed by selectively removing from the reaction mixture (a) the oxidant (H₂O₂), (b) the reducing agent (AA) and (c) the catalytic source (CyC). Additionally, to assess the involvement of radicals, a scavenging test (d) was performed using 2,2,6,6-tetramethylpiperidine 1-oxyl (TEMPO).

In detail, when the reaction was performed in the absence of H₂O₂ (a; Fig. S5) or AA (b; Fig. S6), no gelation was observed, which suggested that the radical polymerization is initiated by the H₂O₂/AA system through the generation of reactive radical species, which may include hydroxyl radicals.²⁹ Accordingly, the control experiment (c) without CyC suggested that the H₂O₂/AA system can generate initiating radical species; however, it results in a slower (>5 minutes) polymerization process and the formation of a poorly cross-linked gel (Fig. S7). Finally, the scavenging experiment (d) showed that in the presence of TEMPO, no gelation is observed, consistent with the trapping of propagating radicals and suppression of the radical polymerization (Fig. S8).

The outcomes of the above-described experiments are consistent with a radical-mediated polymerization mechanism and with the Fenton-like pathway proposed in eqn (1)–(4). Overall, our results support the role of CyC as a redox mediator contributing to faster polymerization kinetics.

Hydrogel rheological characterization

Gelation kinetics were then evaluated at 37 °C to assess the suitability of the hydrogels for in situ gelation (Fig. 2A). All hydrogels exhibited rapid gelation, with the sol–gel transition defined by the G′–G″ crossover occurring within 2 min (1.7 min for Gel-PEG A, 1.5 min for Gel-PEG B, and 1.0 min for Gel-PEG C). The polymerization reaction reached completion in less than 5 min for all hydrogels, as evidenced by the plateau.

Fig. 2 (A) Gelation kinetics of Gel-PEG A–C; (B) elastic modulus of Gel-PEG A–C (data are represented as mean); (C) step-strain recovery test of Gel-PEG A–C at a strain of γ% = 1 (t = 2 min) and at a strain of 500% (t = 1 min).

This rapid gelation is advantageous for localized drug delivery, as it minimizes the diffusion of liquid precursors and ensures precise spatial confinement of the therapeutic agents. Dynamic mechanical properties were evaluated through both the strain test and the oscillation frequency model (Fig. 2B and C and S9). Increasing the polymer concentration from Gel-PEG A to Gel-PEG C resulted, as expected, in progressively stiffer hydrogels, as reflected by an increase in the storage modulus (G′) (Fig. 2B), consistent with a denser polymer network and thus enhanced mechanical rigidity.³⁰ Notably, the rise in stiffness from Gel-PEG A to Gel-PEG B was accompanied by an expansion of the linear viscoelastic region (LVR), with the strain limit (γ) increasing from 13% for Gel-PEG A to 33% for Gel-PEG B (Fig. S9). This behavior suggests a balance between covalent crosslinking and dynamic non-covalent interactions, generating a network capable of withstanding deformation before structural breakdown. In contrast, further increasing the polymer concentration from Gel-PEG B to Gel-PEG C resulted in a sharp reduction in the LVR (γ = 4%), consistent with a more brittle structure (Fig. S9).

These findings indicate that increasing the polymer content enhances stiffness (higher G′) but reduces the strain range over which the hydrogels behave linearly, consistent with a lower critical strain. Oscillatory frequency sweeps performed at a fixed strain amplitude within the LVR (Fig. S9) showed predominantly elastic behavior (G′ > G″) at low frequencies, with a G′–G″ crossover occurring at 9–10 Hz for all hydrogels. Moreover, a step-strain recovery test was performed to evaluate the self-healing properties of the formulated hydrogels (Fig. 2C).

The results showed that Gel-PEG A only partially recovers its stiffness, and this trend remains consistent across all creep–recovery cycles. This behavior is likely attributable to the breaking of covalent bonds within the hydrogel network, which limits the complete reformation of the three-dimensional structure. Conversely, Gel-PEG B and Gel-PEG C exhibited full recovery of both G′ and G″ values after each creep phase, suggesting that the applied deformation did not significantly affect the crosslinking network. Interestingly, the order of mixing the components is critical for obtaining gels with the described properties. In fact, the syringe must be loaded with the solution containing GelMA, the PEGDA crosslinker, and CyC first, and then with the solution containing the oxidant and ascorbic acid (AA).

Hydrogel structural characterization

The microscopic structure of the hydrogels was assessed by scanning electron microscopy (SEM) analysis of freeze-dried samples (Fig. 3A; see SI). It should be noted that the observed pore morphology corresponds to the lyophilized hydrogel state and may differ from the native hydrated structure due to structural rearrangements induced during freeze-drying. Nevertheless, since all hydrogel formulations were subjected to the same lyophilization protocol, SEM analysis still enables comparative evaluation of differences in pore density, pore size uniformity, and overall network organization between the samples.

Fig. 3 (A) SEM images of (I) Gel-PEG A, (II) Gel-PEG B, and (III) Gel-PEG C – scale bar is 50 µm; (B) degradation kinetics of Gel-PEG A–C in collagenase A (1 mg mL⁻¹ in PBS, solid lines) and PBS (dashed lines); (C) ATR-IR spectra of GelMA and Gel-PEG A–C – spectra normalized by the intensity of the C–O frequency (2880 cm⁻¹).

The SEM images show that all hydrogels possess a porous three-dimensional microstructure with interconnected pores, while also displaying differences in pore density, pore size uniformity, and overall network organization depending on the formulation. Compared to Gel-PEG A and Gel-PEG B, Gel-PEG C exhibits a lower pore density and a more compact microarchitecture (Fig. 3A-III). This could explain the more brittle structure of Gel-PEG C, as the denser network may limit chain mobility and reduce the network's ability to dissipate mechanical energy under load.³¹ The degradation profiles of the Gel–PEG hydrogels after incubation with collagenase A showed a gradual decrease in mass loss with increasing polymer concentration, ultimately resulting in no detectable degradation of Gel-PEG C (Fig. 3B). This trend can be attributed to the higher density of chemical crosslinks, which reduces the susceptibility of GelMA to enzymatic hydrolysis. To confirm the successful incorporation of PEGDA into the hydrogel matrix, ATR-IR (attenuated total reflection infrared spectroscopy) spectra of lyophilized samples were recorded (Fig. 3C). The C–O stretching vibration (1080 cm⁻¹) showed increased intensity compared to pristine GelMA, consistent with the ether bonds of the PEGDA crosslinker. Moreover, the spectra exhibited a stronger C–H sp³ stretching band at 2880 cm⁻¹, again attributable to the PEGDA chains. A decrease in the C=C stretching band, typical of GelMA,³² at 1640 cm⁻¹ was also observed, consistent with the conversion of methacrylic double bonds in the radical polymerization.

Thermogravimetric analysis (TGA) confirmed that Gel-PEG A–C undergo complete thermal degradation (Fig. S11). A first mass loss at 162–169 °C was attributed to the release of water strongly bound to the gelatin through polar interactions, as previously reported for gelatin materials.³³ The second degradation event, occurring at 270–290 °C, corresponded to GelMA decomposition. A third step at 376–380 °C was assigned to the degradation of the PEGDA units, followed by a final event at 440–450 °C associated with the residual GelMA fraction, leading to the full thermal decomposition of the network.

Hydrogel biocompatibility and release profile

Cell culture assays were carried out to evaluate the biosafety of compounds released from the gels after gelation. Since the proposed hydrogels have been developed within the context of locoregional therapeutic strategies for hepatic applications, Human Hepatoma G2 (HepG2) cells were selected as a model cell line for preliminary biological assessment. In detail, HepG2 cells were treated with Gel-PEG A, B and C and cell viability tests were performed after 24 and 48 h (Fig. 4A and B) using the MTT assay (λ_abs = 565 nm; see the SI). After 24 and 48 h (Fig. 4A and B), no differences were observed in cell viability when comparing untreated cells (labelled as Untr) and cells treated with the hydrogels (labelled as Gel-PEG A, B and C). The results obtained suggested that none of the hydrogels releases cytotoxic by-products.

Fig. 4 Cell viability test at (A) 24 h and (B) 48 h of incubation; results are expressed as mean ± SD (N = 3) – one-way ANOVA statistics analysis with Šídák's post hoc multiple comparisons test was used (ns: p > 0.05); (C) release of free Rhodamine 101 embedded in the hydrogel in PBS at 37 °C (data are represented as mean).

We next sought to perform sustained-release studies on Gel-PEG B, which was found to be the formulation combining the best properties in terms of rapid gelation, preserved injectability, and an extended linear viscoelastic region.

A release assay was performed using Gel-PEG B and the water-soluble dye Rhodamine 101³⁴ (Fig. 4C) as a drug-mimicking compound. The dye was loaded into Gel--PEG B at a concentration of 0.13 mg mL⁻¹ and then incubated (37 °C under a humidified atmosphere) in PBS. Cumulative release was determined by measuring the emission intensity (λ_em = 595 nm) of the supernatant at each time point (see SI). After the initial burst that reaches 50% within 24 h, the release profile gradually approached a plateau, reaching about 70% cumulative release after 7 days. After 15 days, more than 90% of the dye was released from the hydrogel. The observed burst release (≈50% in 24 h) may be attributed to the heterogeneous distribution of Rhodamine 101 within the hydrogel network. In detail, the dye molecules located close to the surface diffuse rapidly into the surrounding medium; conversely, the dye molecules located in the core of the hydrogel network display limited diffusion, resulting in a slower release.

Conclusions

Three fast-gelling (1–2 minutes), biodegradable, and injectable Gel-PEG hydrogels (Gel-PEG A–C) were synthesized via Fenton-like radical polymerization reactions mediated by cytochrome C (CyC). We have demonstrated that it is possible to modulate the gelation time, stiffness, and degradation kinetics by simply varying the concentrations of GelMA and PEGDA. Our combined structural, rheological and release studies showed that Gel-PEG B is the most promising formulation, in that it combines rapid gelation (1.5 min) and preserved injectability, conserved linear viscoelastic region (LVR) and mechanical strength along with biodegradability and good cargo-release performance (70% release over 7 days).

Overall, the hydrogels presented in this work represent a promising platform for the construction of fast-injectable formulations allowing for localized drug release.

Author contributions

AF: synthesis of the hydrogels, characterization and biological tests; ES: synthesis of the hydrogels, characterization and biological tests; AA: synthesis of the hydrogels; HSV: SEM experiments and data analysis; VB: synthesis of the hydrogels; LM: mechanistic experiments, supervision, conceptualization, data analysis, and writing & review; LDC: conception of the idea, supervision, writing & review, and funding acquisition. All the authors discussed and commented on the manuscript.

Conflicts of interest

The authors declare no conflicts of interest.

Data availability

The data supporting this article have been included as part of the supplementary information (SI). Supplementary information is available. See DOI: https://doi.org/10.1039/d6ob00479b.

Acknowledgements

We gratefully acknowledge Project Next Gen EU and MUR, PNRR M4C2, “Mano-HCC”, PE_00000019, HEAL ITALIA, CUP E93C22001860006. H.S.V. sincerely thanks the Alexander von Humboldt Foundation for financial support.

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Footnote

† These authors contributed equally to this work.

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