# Water-based synthesis of dextran-methacrylate and its use to design hydrogels for biomedical applications

Bastien Darmau, Antoine Hoang, Andrew J. Gross, Isabelle Texier

► To cite this version:

Bastien Darmau, Antoine Hoang, Andrew J. Gross, Isabelle Texier. Water-based synthesis of dextran-methacrylate and its use to design hydrogels for biomedical applications. European Polymer Journal, 2024, 221, pp.113515. 10.1016/j.eurpolymj.2024.113515. hal-04912916

HAL Id: hal-04912916

https://hal.science/hal-04912916v1

Submitted on 5 Dec 2025

HAL is a multi-disciplinary open access
archive for the deposit and dissemination of sci-
entific research documents, whether they are pub-
lished or not. The documents may come from
teaching and research institutions in France or
abroad, or from public or private research centers.

L'archive ouverte pluridisciplinaire HAL, est
destinée au dépôt et à la diffusion de documents
scientifiques de niveau recherche, publiés ou non,
émanant des établissements d'enseignement et de
recherche français ou étrangers, des laboratoires
publics ou privés.

<figure><img src="image_1.png" /></figure>

Distributed under a Creative Commons Attribution 4.0 International License
# Water-based synthesis of dextran-methacrylate and its use to design hydrogels for biomedical applications

Bastien Darmau<fnref n="1" />,<fnref n="2" />, Antoine Hoang<fnref n="1" />, Andrew J. Gross<fnref n="2" />, Isabelle Texier<fnref n="1" />*

## Abstract

Polysaccharide-based hydrogels are desirable for biomedical applications owing to their biocompatibility and physicochemical tunability. The chemical modification of polysaccharides with photo-sensitive groups, such as methacrylates is a common method to obtain new hydrogel materials. This study introduces a non-toxic water-based method to effectively functionalize dextran with methacrylate groups. The methacrylation reaction with methacrylic anhydride in water in the presence of NaOH was rapid and efficient (85% - 92% yield), permitting degrees of substitution (DS) up to 62% within 60 minutes. An unconventional solid-state photo-crosslinking method was employed to form chemically crosslinked dextran-based hydrogels with LAP (lithium phenyl-2,4,6-trimethylbenzoylphosphinate) as the photoinitiator. By varying the polymer formulations (DS, polymer mass, LAP concentration), a wide range of hydrogels were obtained with various swelling ratios (40 - 250%) and release kinetics of model drug and protein biomolecules. Compression modulus values ranged from 32 ± 1 to 342 ± 10 MPa (dry state) and 87 to 8500 kPa (swollen state). Cytotoxicity experiments indicated good biocompatibility for the crosslinked dextran hydrogels. The green synthesis protocols and obtained dextran-based hydrogels with high mechanical strength open up perspectives for applications from tissue engineering to the design of medical devices such as microneedles.

## Keywords

Dextran methacrylation, UV visible photo-crosslinking, hydrogel biomaterials, compression modulus, drug delivery

<footnote n="1">¹ Université Grenoble Alpes, CEA, LETI-DTIS, 38000 Grenoble, France</footnote>

<footnote n="2">² Département de Chimie Moléculaire (DCM), Univ. Grenoble Alpes – CNRS (UMR 5250), 570 rue de la Chimie, 38041 Grenoble, France</footnote>

<footnote>* [Isabelle.texier-nogues@cea.fr](mailto:Isabelle.texier-nogues@cea.fr); tel +33 438 784 670</footnote>
## Highlights

* A catalyst-free water-based synthesis of dextran-methacrylate was developed

* The reaction was rapid (1 h), efficient (≈ 90% yield) and required pH control

* Polymer degrees of substitution up to 62% were achieved

* Solid-state photo-crosslinking to obtain dextran-based hydrogels

* Biocompatible hydrogels displayed high mechanical strength

## Graph abstract

<figure><img src="image_2.png" /></figure>
# 1. Introduction

Biomedical applications require functional materials that are soft and biocompatible yet mechanically-resistant to ensure robust and ideally inflammation-free integration with biological systems. Biopolymers such as polysaccharides, proteins, or natural polyesters, have attracted much attention these last decades as an alternative to synthetic polymers for applications such as prosthetic coatings, fabrication of implants and tissue substitutes, or drug delivery systems [1-3]. In particular, polysaccharide-based hydrogels have aroused considerable interest due to their unique properties. Their hydrophilicity and ability to absorb and retain water (swellability) confer them with structural properties close to those of the extracellular matrix of connective tissues, with high water content, polysaccharide content, and porosity. This porosity allows the diffusion of oxygen and nutrients when the material is used, for example, as a prosthetic coating or a tissue substitute, and can be controlled to modulate the diffusion kinetics of active ingredients loaded into the matrix for drug delivery applications. The hydrogel infusible three-dimensional crosslinked polymer network confers them with stability and directs their mechanical properties that can be tuned to match those of soft and hard tissues [3-4]. The stability, swelling behavior, and mechanical properties of hydrogels largely rely on the physical and chemical interactions between polymer chains. These interactions depend on parameters such as the chemical nature of the polymer and the crosslinks, the molecular mass of the polymer chains, the effective network density (i.e. the mesh size), and the fabrication process [4].

Recently, much interest has been focused on the synthesis of hydrogels based on green chemistry principles, especially considering their biomedical applications [5]. Green chemistry principles applied to hydrogel synthesis include the use of energy efficient reactions and less hazardous solvents (e.g. water) and chemicals (such as genipin [6] or tannic acid [7]) for polymer crosslinking, and the use of polymers issued from renewable resources. In this perspective, nature-derived polysaccharide-based hydrogels are considered to bring important ecological benefits while presenting intrinsically high levels of biocompatibility, biodegradability and bioresorbability. Polysaccharides and their derivatives can also be degraded naturally by hydrolysis using carbohydrate-active enzymes that react with high specificity and mild conditions. This enzymatic degradation pathway can be exploited to accelerate and/or release therapeutics.
Photo-chemical crosslinking of polymers modified with photosensitive groups is a simple and versatile approach to obtain functional hydrogels [8]. After chemical modification, the photo-crosslinking process involves activating the photosensitive groups grafted onto the polymer chains by light (typically in the near UV range of 365 nm to 405 nm) in the presence of a photosensitive initiator. The crosslinking process involves the production of radicals followed by free radical polymerization that connects the polymer chains to each other to form 3D-structured crosslinked networks [9]. Although the photo-chemical crosslinking route requires the chemical pre-modification of the polymer, the reactions can be fast and performed under mild conditions with a good level of temporal and spatial control. Moreover, cytocompatible and water-soluble photoinitiators such as lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP) and Irgacure™ 2959 are nowadays commercially available [10-12]. Methacrylates are attractive photosensitive functional groups to graft onto polysaccharide backbones because of their photo-crosslinking efficiency and the wide range of functional derivatives available. Various methacrylating agents have been reported in the literature, including methacrylic anhydride (MAh) and glycidyl methacrylate. Glycidyl methacrylate is a good methacrylating agent for polysaccharides, thanks to the presence of the highly reactive epoxy function [13-16]. However, glycidyl methacrylate presents an important toxicity. The epoxy-based reaction also lacks specificity, and its fast kinetics could introduce heterogeneity into functionalized polymers, making it less attractive compared to MAh for the preparation of polysaccharide hydrogels for applications in the biomedical field [15,17]. MAh has been largely employed for the methacrylation of polysaccharides due to the well-known controllable reactivity of the anhydride function. MAh was used for the methacrylation of hyaluronic acid [15-16], sodium alginate [13], sodium carboxymethylcellulose [18-19], chondroitin sulfate [20], cellulose [21-23], methyl cellulose [22], starch [24], and chitosan [25-26]. Interestingly, the reaction can be performed in water but usually requires 12-24 hours of reaction. The main drawback of MAh methacrylation comes from the large excess of methacrylic anhydride that is typically required to efficiently modify the polysaccharides. We herein describe a new protocol for the synthesis of dextran-methacrylate (Dex-MA) that requires a more reasonable excess of MAh.

Dextran is an extracellular polysaccharide produced from bacteria that is essentially composed of α-1,6 linked D-glucopyranose residues with α-1,2, α-1,3, or α-1,4 linked side chains [27]. Dextran, which has a neutral structure and is highly water-soluble at
room temperature, has been used for many applications, from food and drug formulations, to cosmetics and tissue regeneration scaffolds [27-30]. In particular for drug delivery, dextran based hydrogels have demonstrated their high safety, their properties to limit non-specific protein adsorption [31], and their ability to protect antitumor and other drugs from chemical and enzymatic degradations increasing their stability and bioavailability [27,32]. Dextran based hydrogels benefit from a high level of robustness in aqueous media, yet can be easily degraded by enzymes (dextranases). Enzymatic degradation can therefore be used to control drug payload release with kinetics depending on drug concentration and interactions with the hydrogel matrix, the nature of the dextran hydrogel, and various release conditions (pH, temperature,...) [33]. Additionally, dextran materials crosslinked through ester bonds, like dextran-methacrylate hydrogels, can be degraded by abundant human body esterases, making them of particular interest as drug delivery systems [27]. Since the first report on dextran acrylation by Edman et al., there has been high interest in dextran functionalized with photoactivatable acrylate groups [34]. In the method of Edman et al., dextran-acrylate was prepared by reacting dextran with glycidyl acrylate in carbonate buffer at pH 11 for 5 days. A very limited degree of substitution (DS) of 5 % was obtained that corresponded to < 10% of the initial amount of glycidyl acrylate. A highly efficient synthesis of dextran-methacrylate was later proposed by van Dijk-Wolthuis et al., in which dextran was methacrylated with glycidyl methacrylate in the presence of 4-dimethylaminopyridine (DMAP) as organic catalyst in dimethyl sulfoxide (DMSO) under nitrogen [35-36]. DMAP is highly toxic and particularly dangerous because it can be absorbed through the skin. More than 90% of the reacting glycidyl methacrylate was introduced onto the dextran polymer with a good level of control over the methacrylation to give polymers with DS from 4.5% to 45%. Several years later, Kim et al. described the methacrylation of dextran in lithium chloride/dimethylformamide (LiCl/DMF) mixture under nitrogen, using methacrylic anhydride and triethylamine as a catalyst [37]. Although a higher DS up to 75% could be obtained with increasing temperature and reactant concentration, a lower control over the reaction, methacrylation efficiency and polymer purity were observed compared to the procedure of Dijk-Wolthuis et al. Even if other routes have been reported for the methacrylation of dextran [38-39], the two procedures that employ MAh or glycidyl methacrylate in organic solvent with a catalyst or base are the mostly commonly adopted. Like DMAP, DMF and LiCl are toxic, hence toxicity remains an important and
overlooked issue for dextran methacrylate synthesis, especially in the context of the
growing importance of green and sustainable chemistry, and the importance of
minimizing the use of hazardous materials for biomedical applications [40-41].

We report here a focused study on the first aqueous synthesis of dextran methacrylate (Dex-MA). The optimized reaction employed MAh with a low number of equivalents and a small amount of NaOH, and required careful consideration of the pH. The reaction was otherwise surprisingly simple and ecofriendly. The reaction was performed in air and without toxic products, unlike the methods of Dijk-Wolthuis et al. and Kim et al., and was also carried out at room temperature. The methacrylation reaction of dextran was carefully examined using FTIR, NMR and GPC methods. Various parameters were investigated to obtain a high level of control over the reaction, and access a wide range of methacrylated dextran polymers with different DS. The polymers were subsequently crosslinked using a green synthesis photo-chemical crosslinking method at 405 nm to yield resistant but swellable hydrogels with tunable properties. The mechanical properties and chemical/enzymatic degradation of the materials were investigated, and a proof-of-concept demonstration of model drug release in aqueous buffer was subsequently demonstrated.

**2. Experimental section**

**2.1. Materials**

Dextran T70 (Dex T70, $\overline{M_w}$ = 70,000 g/mol) and dextran T20 (Dex T20, $\overline{M_w}$ = 20,000 g/mol) were purchased from Pharmacosmos (Holbæk, Denmark), methacrylic anhydride (MAh, 94%), lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP, ≤ 95%), potassium hydroxide (KOH, ≥85%), potassium carbonate ($K_2CO_3$, ≥ 99%), triethylamine (Et₃N, ≥ 99%), fluorescein isothiocyanate-labelled bovine serum albumin (BSA-FITC), ketorolac tris salt (≥ 99%), HPLC-grade acetonitrile (≥ 99,5%), trifluoroacetic acid (TFA, ≥ 99%), phosphate buffer saline (PBS: 137 mM NaCl, 2.7 mM KCl, 10 mM $Na_2HPO_4$, 1.8 mM $KH_2PO_4$, pH = 7.4) and dextranase from Chaetomium erraticum (liquid mixture and used as such) were purchased from Sigma-Aldrich (Saint-Quentin Fallavier, France). Sodium hydroxide (NaOH, ≥ 98%) and sodium carbonate ($Na_2CO_3$, 99%) were purchased from Laurylab (Brindas, France). NIH-3T3 murine fibroblast cells were purchased from the American Type Culture Collection (ATCC). WST-1 assay kit was purchased from Roche Diagnostics.
## 2.2. Synthesis of dextran-methacrylate (Dex-MA)

Dex (5 g) was dissolved in 100 mL of distilled water in a beaker. After complete dissolution, different equivalents (0.0625 to 0.5 eq.) of MAh, compared to the Dex hydroxyl groups of the polymer, were added dropwise into the polymer solution. The solution was stirred at room temperature (RT) for 1h. The pH was adjusted using NaOH (3 mol/L) to keep the pH at around 9-11 during the reaction. The obtained Dex-MA solution was dialyzed against distilled water for 1 week (12-14 kDa molecular weight cut-off membrane, Carl Roth, Lauterbourg, France), and lyophilized. The white solid was stored at -20°C prior to characterization and use.

## 2.3. Characterization of Dex-MA

Fourier transform infrared (FTIR) spectroscopy was performed with a Shimadzu IRAffinity-1 spectrophotometer with a DLATGS detector (deuterated triglycine sulphate doped with L-alanine). The measurements were carried out under a dry atmosphere at RT with lyophilized samples compressed against a KBr disk. ¹H and ¹³C Nuclear Magnetic Resonance (NMR) spectra were recorded in either D₂O or DMSO-d₆ using a Bruker Avance-300 spectrometer operating at 298°K, at 300 MHz and 75.5 MHz, respectively. UV-visible absorption spectra were recorded using a Varian Cary 300 UV-Vis spectrophotometer at 25 °C and a quartz cell with a 1 cm path length. Gel permeation chromatography coupled with a multi-angle light scattering detector (GPC-MALS) was performed using a Shimadzu HPLC chain composed of a SIL-20A auto-sampler, a LC-20AD high pressure pump, a DGU-20A 3R degasser, and a CTO-20A column oven. The eluent was composed of 0.1 M sodium nitrate and 0.03% sodium azide at a flow rate of 0.5 ml/min. The samples prepared in the eluent (2 mg/mL) were filtered (0.2 µm filters) and 100 µL were injected on two Shodex OHpak SB 806 M HQ columns placed in series at a temperature of 30°C. The elution time was 70 minutes. Detection was carried out with a Wyatt refractometer, a Shimadzu UV detector set at 280 nm, and a Wyatt 3-angle light scattering detector. The value of the refractive index increment (dn/dc) used for the analysis was 0.142 mL/g. The analyses were performed on three different samples.

## 2.4. Preparation of Dex-MA hydrogels

PBS solutions comprising a mixture of Dex-MA at 10%, 20% or 30% m/m (mass fraction) polymer concentration, and LAP photoinitiator at 0.1%, 0.5% or 1% m/m, were
prepared. The solutions were then cast into a Teflon ≈ 1 cm³ mold and left to dry under air for 24 hours. Each mold contained the same polymer formulation weight (500 mg) irrespective of the formulation. After total drying of the polymer solution and demolding, the solid was exposed to 405 nm light (UVAHand, Eleco-Panacol, Gennevilliers, France, 75 mW/cm²) for 60 s. Swollen hydrogels were obtained by equilibrating the dried photo-crosslinked pellets for 24 hours in PBS. To obtain BSA-FITC or ketorolac loaded hydrogels, Dex-MA (10%, 20% or 30% m/m), LAP (0.1% or 1% m/m), BSA-FITC (0.1% m/m), or ketorolac (0.1% m/m), were solubilized in the Dex-MA mixtures, then dried, photo-crosslinked in the dry state as described above, then re-hydrated with PBS.

## 2.5. Swelling, insoluble mass fraction, chemical and enzymatic degradation studies

Dried pellet samples were weighed ($m_0$) after UV photo-crosslinking. The samples of around 20 mg were then incubated in a vial containing 1 mL of PBS at RT.

**Swelling ratio (SR):** Samples were periodically removed from the vial, quickly and delicately wiped with a paper towel to remove liquid excess, immediately weighed ($m_t$), then placed back into the vial containing PBS. This operation was repeated until no further mass change was detected (24 hours maximum). The percentage of SR was calculated according to the formula:

$$ \%SR = \frac{m_t - m_0}{m_0} \times 100 $$

**Insoluble mass fraction (IM):** The samples were incubated in 1 mL of PBS at 37°C for one week. After one week, samples were removed, dried completely in an oven at 80°C, then weighed to obtain the insoluble mass $m_{insoluble}$. The percentage of IM was calculated according to the formula:

$$ \%IM = \frac{m_{insoluble}}{m_0} \times 100 $$

**Chemical degradation (CD):** The samples were incubated in 1 mL of PBS at 37°C for 90 days. The samples were removed periodically, dried completely in an oven at 80°C, and then weighed to obtain the dry weight, $m_d$. The percentage of chemically degraded (CD) polymer network at a given time was calculated according to the formula:

$$ \%CD = \frac{(m_0 - m_d)}{m_0} \times 100 $$
Enzymatic degradation (ED): Aliquots of around 30 mg of equilibrated swollen hydrogels were first weighed ($m_0$). The samples were then incubated in 1 mL of a mixture of dextranase/PBS (v/v 1/10) solution at 37°C. The samples were removed from their vials every 2 hours, quickly wiped with a paper towel to remove excess solution, then immediately weighed ($m_d$). After each measurement, the sample was placed back into the dextranase/PBS solution at 37°C. This operation was repeated until further mass changes were minimal. The percentage of enzymatically degraded material (%ED) was calculated according to the formula:

$$ \%ED = \frac{(m_0 - m_d)}{m_0} \times 100 $$

For each measurement, three experiments were performed in parallel for each sample.

## 2.6. Compression tests

Compression tests were performed using a TAXT.Plus texturometer (Stable Micro Systems) on dried and equilibrium-swollen hydrogels. The compression speed was set at 0.5 mm.s⁻¹ with a maximum compression force of 50 N. The values of the compression modulus (E) were calculated from the slope of the linear portion of the stress-strain curves between 5% and 15% strain. The ultimate compression strength (UCS) corresponds to the maximum applied load before material failure. The ultimate elongation strength (US) corresponds to the elongation achieved at the point of material failure. Three experiments were performed in parallel for each sample.

## 2.7. Scanning electron microscopy (SEM)

Two Dex-MA hydrogels ((10% m/m, DS=13%, 0.1% m/m LAP), and (30% m/m, DS=62%, 1% m/m LAP)) were analyzed. Some material samples were characterized in the dry state after storage under ambient conditions while others were prepared as cryogels after immersion of the hydrogels in PBS for 24 hours (swollen), freezing in liquid nitrogen, then lyophilization. The dry hydrogels and cryogels were fractured, placed vertically in a holder, and metalized with a thin layer of a few nm of Au/Pd prior to analysis. SEM micrographs were recorded using a FEI/Quanta FEG 250 scanning electron microscope (Thermo Fisher Scientific Co., Waltham, MA, USA) operating with an accelerating voltage of 2.5 kV and a secondary electron detector.
## 2.8. LAP, photo-irradiated LAP, and drug release studies

LAP, photo-irradiated LAP, ketorolac, and BSA-FITC release from Dex-MA hydrogels was studied. Plain (control) or drug/protein-loaded crosslinked hydrogels were immersed in a closed and static vial containing 1 mL of PBS at RT. The release media were collected for measurement at regular times. After each collection, the initial vial was topped up with fresh PBS. The percentage of photoinitiator, drug, or protein release, was quantified using either high-performance liquid chromatography (HPLC) (LAP, photo-irradiated LAP, ketorolac) or UV-visible spectroscopy (BSA-FITC) via calibration curves. HPLC was performed using a Waters Alliance 2695 separation module equipped with a Waters 2487 dual absorbance detector and a reverse phase C18 column (3.5 µm / 4.6 mm x 75 mm). Data were processed using the Empower™ 2 software. The elution used a gradient mobile phase consisting of (A) water + 0.1% TFA and (B) acetonitrile + 0.1% TFA. The gradient started at 95% of A and decreased to 0% of A (100% of B) at 20 min. After 5 min, the percentage of buffer A linearly increased in 1 min from 0% to 95%, followed by 4 min equilibration. A flow rate of 1 mL/min was used at RT with a 20 µL injection volume. The UV detector was set at 270 nm to detect LAP and photo-irradiated LAP with a retention time of 9.74 ± 0.06 min and 12.71 ± 0.01 min, respectively, and at 322 nm to detect ketorolac with a retention time of 9.43 ± 0.01 min. To quantify BSA-FITC, UV-visible measurements at 494 nm were performed using an Infinite-M1000 Tecan microplate reader. Three experiments were performed in parallel for each sample.

## 2.9. Cytotoxicity

NIH-3T3 fibroblasts were cultured under a humidified atmosphere (90%) of 95% air and 5% CO₂ at 37°C in culture medium with a high glucose content (Eagle medium supplemented with 10% fetal bovine serum and 1% of antibiotics (penicillin and streptomycin; Dubelco)).

The cytotoxicity tests were carried out according to ISO-10993-5 with crosslinked Dex-MA T70 hydrogels (20% m/m, DS=37%, 1% m/m LAP). All samples were sterilized in an ethanol/water 7/3 mixture for 3 hours. The ethanol/water solution was then removed and replaced twice by sterile PBS for 1 hour to eliminate all traces of ethanol. The samples were then set in a 24-well plate and immersed in 1 mL of culture medium (3 cm² of material for 1 mL of cell culture medium) for 24 hours, 48 hours, and 7 days. Control plates without materials and cell culture medium only were used as negative
controls. The release media were then completely collected in Eppendorf tubes after 24 hours, 48 hours, and 7 days, then frozen at -20°C.

The cells were seeded with a cell density of 5,000 cells per well in a 96-well plate for 24 hours at 37.5°C with 5% CO₂. After 24 hours, the growth medium was replaced by 100 µL of material release medium after thawing at 37.5°C, negative control medium, or medium with 10 mM of H₂O₂ as the positive cytotoxicity control. After 24 hours, the cell culture media were removed and replaced with 100 µL of fresh medium and 10 µL of the WST-1 reagent (Roche). After 2 hours of incubation at 37.5°C with 5% CO₂, the cell viability was calculated by reading the absorbance at 450 nm, corrected for the absorbance at 650 nm, with an Infinite-M1000 Tecan microplate reader according to the following formula:

$$ \frac{\text{Absorbance}_{\text{sample}} - \text{Absorbance}_{\text{positive cytotoxicity control}}}{\text{Absorbance}_{\text{negative cytotoxicity control}} - \text{Absorbance}_{\text{positive cytotoxicity control}}} \times 100 $$

Experiments were conducted using three different cell cultures in quintuplicate.

## 2.10. Statistical analysis

All data were expressed as mean ± standard deviation. Statistical differences between three or more sets of data were determined by a one-way ANOVA test, between two sets of data by a Student t-test.

# 3. Results and Discussion

## 3.1. Synthesis of dextran-methacrylate in aqueous medium

The synthesis in aqueous medium of Dex-MA was demonstrated using dextran polymers with different molecular mass, Dex T70 (70,000 g/mol) and Dex T20 (20,000 g/mol). The reaction scheme shown in **Figure 1a** was based on a nucleophilic attack from the hydroxyl groups of the dextran at the carbonyl groups of methacrylic anhydride (MAh). The methacrylating agent was fragmented with the esterification of the polysaccharide on one side, and the formation of methacrylic acid as by-product on the other side. Methacrylic acid was neutralized during the reaction using sodium hydroxide (precise pH-controlled conditions) to direct the reaction towards the esterification.
<figure><img src="image_3.png" /></figure>

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

<figure><img src="image_5.png" /><figcaption>**Figure 1.** Dextran methacrylation in distilled water at room temperature in air. **a)** Reaction scheme. **b)** Time-evolution of Dex T70 DS methacrylation in water using different eq. of MAh. **c)** Time-evolution of unreacted (circles, red) and hydrolyzed (triangles, black) MAh during the methacrylation of Dex T70 using 0.125 eq of MAh.</figcaption></figure>

### 3.1.1. Kinetic study

The kinetic study of the reaction conducted with Dex T70 and different number equivalents of MAh versus the number of hydroxyl groups of a Dex unit is presented in **Figure 1b**. Crude aliquots were taken at regular intervals then analyzed by ¹H NMR in D₂O after freeze-drying. The Dex-MA degree of substitution (DS) was determined by the integration of the proton signals at 5.8 and 6.25 ppm of the vinyl groups of the modified polysaccharide, and the protons of the Dex C₂, C₃, C₄, C₅ and C₆ carbons that give a signal between 3.4 and 4.2 ppm (Supporting Information, **Figure S1**). The signals of the C₁ anomeric protons were also noticeable at 5.1 ppm and 5.4 ppm, corresponding respectively to the α-1,6 linkage and α-1,3 linkage. A classical value of 4% was observed for the α-1,3 linkage. The methyl protons from the methacrylate groups were also present (1.9 ppm). An unexpectedly fast and efficient methacrylation reaction was observed, providing the pH was carefully monitored and controlled during the reaction (**Figure 1b**). Methacrylation DS values were obtained from 13% to 62% in 60 min as a function of the number of MAh equivalents from 0.0625 to 0.3 eq. After
purification by dialysis, yields of 85% to 92% with a linear evolution of the DS versus the number of eq. of MAh were obtained (Figure S2). The linear evolution across a wide DS range, combined with the rapid kinetics and high yield, highlighted a precise level of control over the reaction. We note that the DS values determined by ¹H NMR spectroscopy (Figure S2) were validated using a UV-visible spectroscopy protocol. The data showed good agreement with a difference lower than 3% for each sample across the 13% to 62% DS range. The UV-vis protocol was based on the absorption of the methacrylate moiety at 269 nm (Figure S3).

A white solid formed in increasing quantity from 15 to 60 minutes was observed during the reaction with 0.5 eq. of MAh. ¹H NMR analysis of this white solid re-suspended in DMSO-d6 showed it was dex-MA with an average DS of 84% (Figure S7). The water-soluble product obtained at 60 minutes of reaction with 0.5 eq. of MAh displayed a DS of 62%, similar to that achieved with 0.3 eq. of MAh. Moreover, yields in water-soluble Dex-MA DS=62% were 85% and 9% with 0.3 eq. and 0.5 eq. of MAh, respectively. The precipitation of Dex-MA above the critical DS of ≈ 62% was attributed to the high number of hydrophobic methacrylate groups that decreased the solubility of the modified Dex in water.

The production of the methacrylic acid by-product during the reaction with 0.125 eq. of MAh was also followed, based on the ¹H NMR signals of the vinyl groups of the methacrylic acid appearing at 5.4 ppm and 5.7 ppm. The amount of hydrolyzed and unreacted methacrylic anhydride could then be determined and was plotted as a function of the reaction time (Figure 1c). The data revealed a clear competition between hydrolysis and esterification during the reaction. After 60 minutes of reaction i.e. reaction completion, 69% of MAh had reacted with hydroxyl groups from the dextran, 29% were hydrolyzed by NaOH, and 2% remained unreacted. At 90 minutes, all of the MAh was consumed. Similar results were obtained with Dex T20, as described in Figure S4.

### 3.1.2. Base and pH effect

The methacrylation of polysaccharides using MAh in water involves the production of methacrylic acid and consequent changes in pH towards acidic values. The pH impact on the synthesis of Dex-MA was studied using 0.125 eq., with NaOH added as the base for pH control. The final DS was determined by ¹H NMR for products obtained by controlled reaction at different pH ranges: pH 6-8, pH 9-11, and pH > 12 (Figure S5).
No methacrylation took place without pH control (the addition of NaOH), consistent with the deactivation of Dex hydroxyl groups at acidic pH. At a pH value between 6 and 8, based on unpurified products, a competition between polymer methacrylation and MAh hydrolysis was observed. A 50:50 ratio between the two reactions was determined and attributed to partial deactivation of hydroxyl groups in this pH range. The optimal reaction pH was determined between a range of pH 9 and pH 11. A 69:31 ratio between Dex methacrylation and MAh hydrolysis was observed at this basic pH range. Further highlighting the importance of pH control, when the pH was higher than 12, only a few percent of methacrylate groups were introduced, while a large quantity of methacrylic acid was observed. At strongly basic conditions, the dominant reaction was MAh hydrolysis (Figure S5 and Table S1).

The impact of the nature of the base employed during the reaction was also investigated. Although quite similar results were obtained when NaOH, KOH, and $Et_3N$ were employed, the highest DS was obtained with NaOH. The methacrylation of Dex-MA in water was less efficient with weaker bases ($Na_2CO_3$ and $K_2CO_3$). These results are different to those obtained for the methacrylation of cellulose reported by Chen *et al.* where the reaction was most effective using $K_2CO_3$ [21]. In that case, however, a low cellulose methacrylation DS of 2-6 % was obtained in comparison to the DS of 62% reported here. Similar MAh number equivalents in comparison to the hydroxyl groups of the polymer were used here compared to the work by Chen *et al.*

### 3.1.2. Dex-MA polymer characterization

Purified Dex-MA polymers were characterized by complementary ¹H NMR, ¹³C NMR, FTIR, and GPC studies. The ¹H NMR spectra in $D_2O$ of Dex and the dialysis-purified Dex-MA T70 are presented in **Figure 2**. The spectra recorded in $D_2O$ were free of signals from MAh and other impurity traces, indicating well-purified polysaccharide polymers. Similar results were obtained with Dex T20, as shown in **Figure S6**. The ¹H NMR spectra of Dex and Dex-MA were also recorded in DMSO-d6 and are shown in **Figure S7**. Unlike the NMR spectra recorded in $D_2O$, the protons from the hydroxyl groups were visible in the DMSO-d6 spectra, offering further insight into the reaction. From the DMSO-d6 spectra, a good correlation was determined between the increase of the DS and the decrease of the hydroxyl signals (**Figure S7** and **Table S2**). Moreover, a similar decrease for each hydroxyl group was observed, indicating their similar reactivity during the methacrylation reaction. The ¹³C NMR spectra obtained in
387

D₂O for Dex and Dex-MA (DS=37%) are presented and interpreted in **Figure S8**. The structure of Dex-MA was confirmed by HSQC NMR as well (**Figure S9**).

388

<figure><img src="image_6.png" /></figure>

389

390

<caption>**Figure 2.** ¹H NMR spectra in D₂O at 300 MHz of dextran T70 (a), and Dex-MA T70 with DS=13% (b), DS=26% (c), DS=52% (d), DS=62% (e).</caption>

391

392

393

394

395

396

397

398

399

400

401

402

The FTIR spectra of Dex and Dex-MA (T70 and T20) evidenced the characteristic bands of Dex at 3100-3600 cm⁻¹ (-OH) and 2930 cm⁻¹ (-CH₂) (**Figure S10** and **S11**, respectively). The successful incorporation of the methacrylate groups on the Dex was confirmed by the apparition of ester bonds, as evidenced at 1710 cm⁻¹ (C=O), 1636 cm⁻¹ (C=C), 1299 cm⁻¹ and 1147 cm⁻¹ (C-O). The number of methacrylate groups appeared to increase with increasing DS based on the intensifying signal of these characteristic peaks. A decrease of the hydroxyl band between 3100-3600 cm⁻¹ (-OH) was noticeable as the DS increased, indicating an increasing esterification of these functions.

403

404

405

406

The GPC spectrum of Dex T70 and Dex-MA T70 DS=52% (**Figure S12**) showed an increase of the mean molecular mass of the linear polymer, M̄n, from 63,900 g/mol⁻¹ for the unmodified Dex polymer, to 78,300 g/mol⁻¹, for Dex-MA with DS=52%, consistent with the effective incorporation of methacrylate moieties on the Dex backbone. This
result was quite close to the estimated theoretical values (Table S3). The narrow polydispersity of the unmodified polysaccharide starting material was conserved after the methacrylation reaction (polydispersity index (PDI) of 1.14 and 1.09 for Dex and Dex-MA, respectively), seemingly indicating a homogenous methacrylation of the polymer chains. A higher elution time, despite an increase of the $M_n$ of the Dex-MA, was also noticeable. This was attributed to a decrease of the gyration radius of the polymer chains due to the incorporation of hydrophobic methacrylate moieties. A similar phenomenon was also observed by Van Dijk et al., although no significant difference in elution time between a Dex T40 and Dex-MA T40 DS=11% was observed in that case [35]. Finally, no additional peaks were observed in the chromatogram of Dex-MA, showing that the initial polysaccharide polymer chains were not cleaved during the methacrylation reaction.

## 3.2. Preparation of photo-crosslinked Dex-MA materials

The photo-chemical reaction scheme is shown in Figure 3a. PBS solutions containing LAP (0.1% to 1% m/m) and Dex-MA at a fixed mass but different concentrations (10% to 30% m/m) were first poured into Teflon molds. After water evaporation and air-drying, the hydrogels were crosslinked using an unconventional photo-crosslinking reaction at the solid (dried) state rather than the classic wet-state approach (Figure 3). After a short UV-crosslinking reaction of 1 minute, the dried materials were swollen in PBS to yield hydrated and stable Dex-MA hydrogels. Compared to photo-crosslinking in the wet state, this process allowed one to control the material shape by avoiding shrinkage that can occur during wet crosslinking. The dry Dex-MA pellet materials were initially transparent then turned slightly yellow following UV irradiation. For materials prepared from Dex-MA with a high DS above 52%, and especially when a high LAP concentration was used (1% m/m), the yellow coloration was particularly obvious. This yellowing of the material was attributed to the photoreaction of LAP (Figure S13a) that resulted in photo-fragmented LAP products (Figure S13b). After immersion in PBS, the materials that were initially yellow turned transparent again, indicating that the colored residues were not chemically attached to the polymer and could be easily washed out. The LAP and photofragmented molecules were fully released after immersion in PBS for around 4 hours, based on HPLC analysis (Figure S14).
<figure><img src="image_7.png" /></figure>

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

<figure><img src="image_9.png" /><figcaption>**Figure 3.** Preparation of Dex-MA hydrogels. **a)** Photo-chemical crosslinking reaction at the dried state; **b)** Scheme of the crosslinked Dex-MA hydrogel preparation process and **c)** photographs of material samples prepared from Dex-MA DS=13% 10% m/m solutions with 0.1% m/m LAP before (i) and after (ii) photo-crosslinking at the dried state, and after 1 hour swelling in PBS (iii).</figcaption></figure>

The efficiency of the dry-state photo-crosslinking process was assessed by measuring the insoluble mass fraction that corresponds to the fraction of the polymer that was efficiently crosslinked and could not diffuse out into the liquid phase. For these experiments, the crosslinked dry materials were weighed, immersed in PBS for 1 week, then dried and weighed again. Dex was completely soluble in PBS, confirmed by a 100% mass loss. For Dex-MA hydrogels prepared with 1% m/m LAP, only a small percentage of polymer mass was washed out after 1 week (Figure 4a). If between 12.2 ± 0.3% and 16.2 ± 0.4% of material mass was lost when a 10% m/m polymer solution was used, while only 8.8 ± 0.2% to 12.1 ± 0.2% of the material mass was solubilized when a 30% m/m polymer solution was used, no statistical difference was observed in insoluble mass fractions for samples prepared at different concentrations from polymers with different DS with 1% m/m LAP. The data in Figure 4a clearly show that
the photo-crosslinking reaction was effective, even at the solid-state, at least for the conditions reported here and with 1% m/m LAP. We subsequently explored the use of a lower LAP concentration to effectively crosslink Dex-MA via dry-state crosslinking (Figure 4b). The efficiency of the photo-crosslinking reaction decreased when 0.5% and 0.1% m/m of LAP were used. The insoluble mass fractions were 71.7 ± 4.1 % and 55.8 ± 5.2 % for Dex-MA DS=13% with 10% m/m crosslinked with 0.5 and 0.1% m/m LAP, respectively. The possibility to effectively crosslink Dex-MA via dry-state crosslinking was again revealed. A small decrease in the insoluble mass fraction and therefore crosslinking efficiency was observed for the smaller Dex-MA T20 polymers compared to the Dex-MA T70 polymers. For example, insoluble mass fractions were 88.6 ± 2.3 % for T70 versus 82.8 ± 3.1 % for T20 for polymer formulations prepared with the same DS of 26%, 30% m/m polymer concentration, and 1% m/m LAP (Figure S15). The slightly reduced crosslinking efficiency with the smaller dextran pointed to an effect of the shorter polymer chain length that reduced the probability of a successful crosslinking reaction.

<figure><img src="image_10.png" /><figcaption>**Figure 4.** Insoluble mass fractions (IM) at equilibrium for Dex-MA (T70) materials: a) influence of the polymer DS and polymer concentration at a LAP concentration of 1% m/m; b) influence of the LAP concentration (% m/m) and polymer concentration for Dex-MA (T70) materials with DS=13%. NS: non-statistically different; ***: p<0.005.</figcaption></figure>

### 3.3. Swelling and structural properties of photo-crosslinked Dex-MA materials

The swelling properties in PBS of the photo-crosslinked Dex-MA materials with different DS as well as polymer and LAP concentrations were subsequently explored. The swelling equilibrium was achieved after around 1–2h for all formulations as exemplified with 10-30% m/m polymer with 1% m/m LAP formulations (Figure S16). The equilibrium swelling ratio (SR) decreased with an increase in the polymer concentration irrespective of the DS or LAP concentration. For example, the SR of
Dex-MA T70 DS=13% with 1% m/m LAP was 111.8 ± 3.0 % at 10% m/m, decreasing to 84.4 ± 2.3 % at 30% m/m (Figure 5a). Similarly, an increase in LAP concentration induced a decrease of the SR, consistent with an increase in the crosslinking efficiency. For materials prepared from 10% m/m solutions of Dex-MA DS=13%, the SR decreased from 238.2 ± 5.6 %, with LAP at 0.1% m/m down to 111.8 ± 3.0 % with LAP at 1% m/m (Figure 5b). For a fixed polymer and LAP concentration, an increase of the SR was observed with a decrease in polymer DS. For example, for Dex-MA at 30% m/m with 1% m/m LAP, the SR increased from 40.2 ± 1.5 % at a DS of 62% to 84.4 ± 2.3 % at a DS of 13%, consistent with a higher crosslink density for the polymer with DS=62%. A comparison of the SR obtained for Dex-MA T70 and Dex-MA T20 polymers with the same DS (DS=26%), polymer amount (10% m/m), and LAP concentration (1% m/m) showed a slightly higher SR value for T70 polymer (66.1 ± 3.1 %) than for T20 (61.6 ± 3.2 %) (Figure S17).

<figure><img src="image_11.png" /><figcaption>**Figure 5.** Swelling ratios (SR) at the equilibrium recorded after 24 hours in PBS for Dex-MA (T70) materials: a) influence of the polymer DS and concentration at a LAP concentration of 1% m/m; b) influence of the polymer and LAP concentration for Dex-MA DS=13%. \*: p<0.05; \*: p<0.01; \*: p<0.005.</figcaption></figure>

The mesh size and the crosslinking density of the swollen hydrogels were calculated based on the SR measurements according to the formula of Peppas *et al*. that was considered for dry-state crosslinked hydrogels (Table S4) [42]. Further details are provided in the Supporting Information. The calculated mesh size was considered non-significant when the value was below the size of one glycoside unit (around 15.5 Å) i.e. for weakly swelling and dense materials with DS > 26%. For materials issued from polymers with DS=13%, the calculated mesh size decreased with an increase in LAP and polymer concentrations. For example, the estimated mesh size was 54.1 ± 2.3 Å
for materials issued from 10% m/m solutions of Dex-MA and 0.1% m/m LAP (Table S4), and only 16.7 ± 1.0 Å for materials issued from 30% m/m Dex-MA solutions and 1% m/m LAP (Dex-MA DS=13%). In parallel, the calculated density of crosslinks increased with the DS as well as increasing polymer and LAP concentrations, going from 1.36 ± 0.09 mmol cm⁻³ to 27.22 ± 2.15 mmol cm⁻³ for the least and most crosslinked materials, respectively (Table S5).

The microstructure of Dex-MA T70 materials with ‘high’ and ‘low’ PBS swellability were characterized by SEM. The analyzed samples were Dex-MA with high SR, high mesh size, and low crosslinking density (Dex-MA with DS=13% at 10% m/m with 0.1% m/m of LAP) and Dex-MA with low SR, low mesh size, and high crosslinking density (Dex-MA with DS=62% at 30% m/m with 1% m/m of LAP). SEM cross-section images of the materials in the dry state, before swelling, and the dry state (lyophilized), after swelling, are shown in Figure 6.

<figure><img src="image_12.png" /><figcaption>**Figure 6.** Cross-section SEM micrographs of dried Dex-MA materials after photo-crosslinking (a), b), scale bar of 200 µm), and after PBS swelling then freeze-drying (swollen state) (c), d), scale bar of 20 µm). a), c): Materials issued from 10% m/m</figcaption></figure>
solutions of Dex-MA T70 with DS=13% and 0.1% m/m of LAP. **b), d)**: Materials issued from 30% m/m solutions of Dex-MA T70 with DS = 62% and 1% m/m of LAP.

For both cases, after material drying and photo-crosslinking, a smooth cross section surface was visible, reflecting limited porosity and high density of the dried materials. Similar results were reported by Kim *et al.* for a non-swollen crosslinked Dex-MA material [37]. At the swollen state, no significant microstructure changes were observed for the 'low swellability' material, consistent with its high density and low hydrogel swelling that was too low to reveal porosity by SEM. For the 'high' swellability hydrogel material (SR=238.2%) with lower crosslinking density and larger mesh size, the micrographs at the swollen state revealed a honeycomb porous microstructure with an estimated pore size distribution between 2.7 µm and 15.7 µm and a mean value of 8.3 ± 3.0 µm (Figure S18).

### 3.4. Mechanical properties of photo-crosslinked Dex-MA materials

The compression moduli of Dex and Dex-MA photo-crosslinked dried materials prepared using different polymer and LAP concentrations were measured and are reported in **Figure S19** and **Table S6**. The stress-strain curves for three different Dex-MA T70 crosslinked hydrogel materials are shown in **Figure S19**. It is noted that the mechanical characterization data of dried materials were limited to compression moduli values since the maximum force of 50 N that our equipment could apply was insufficient to reach sample fracture or deformation. Only the polymer concentration and molecular mass of the Dex-MA significantly impacted the dried material compression properties. The stiffest hydrogel materials with compression modulus values of ≈ 320 to 340 MPa were obtained with Dex-MA T70 at 30% m/m with 1% m/m LAP. Less resistant materials with compression modulus values of 30 to 40 MPa were obtained for Dex-MA T70 with DS=13% at 10% m/m with 0.1% to 1% m/m LAP. Interestingly, only weaker materials with compression modulus ≤ 30 MPa were obtained for the T20 Dex-MA, even for formulations with 30% m/m polymer and 1% m/m LAP. The unmodified dextran T70 was also 10-fold stiffer than the unmodified Dex T20 (Table S6).

The compression modulus (E), ultimate compression strength (UCS) and ultimate strain (US) were also measured on swollen hydrogels after PBS swelling to equilibrium (Figure S20, Table 1).
<table><caption>**Table 1.** Compression modulus, ultimate compression strength, and ultimate strain for equilibrium swollen Dex-MA hydrogels at different DS, polymer, and LAP concentrations. All data were statistically different with p<0.001 (***), except for E and US values of Dex-MA T70 DS=62% versus DS=52% for which p<0.1 (*) for 10% m/m and 30% m/m and non-significantly different for 20% m/m.</caption><thead><tr><th rowspan="2">Hydrogels</th><th>10% m/m</th><th>20% m/m</th><th>30% m/m</th></tr><tr><th>E (kPa)<br>(UCS (kPa); US (%))</th><th>E (kPa)<br>(UCS (kPa); US (%))</th><th>E (kPa)<br>(UCS (kPa); US (%))</th></tr></thead><tbody><tr><td>Dex-MA T70 DS=62%<br>1% m/m LAP</td><td>1320 ± 89<br>(275 ± 12; 25.7 ± 2.2)</td><td>3311 ± 218<br>(436 ± 12; 14.5 ± 0.3)</td><td>8500 ± 119<br>(&gt; 500; &gt; 6.9)</td></tr><tr><td>Dex-MA T70 DS=52%<br>1% m/m LAP</td><td>1327 ± 143<br>(314 ± 13; 26.5 ± 2.2)</td><td>3107 ± 305<br>(385 ± 19; 14.1 ± 0.8)</td><td>7576 ± 357<br>(&gt; 500 KPa; &gt; 6.8)</td></tr><tr><td>Dex-MA T70 DS=26%<br>1% m/m LAP</td><td>624 ± 38<br>(138 ± 8; 23.7 ± 1.1)</td><td>2465 ± 59<br>(390 ± 9; 17.6 ± 0.5)</td><td>6717 ± 342<br>(&gt; 500; &gt; 11.7)</td></tr><tr><td>Dex-MA T70 DS=13%<br>1% m/m LAP</td><td>238 ± 31<br>(57 ± 4; 25.0 ± 0.6)</td><td>597 ± 28<br>(104 ± 4; 22.3 ± 0.5)</td><td>1082 ± 171<br>(166 ± 17; 17.6 ± 1.6)</td></tr><tr><td>Dex-MA T70 DS=13%<br>0.5% m/m LAP</td><td>51 ± 2<br>(40 ± 3; 34.9 ± 1.8)</td><td>155 ± 22<br>(61 ± 7; 24.8 ± 2.9)</td><td>453 ± 50<br>(81 ± 9; 19.7 ± 2.8)</td></tr><tr><td>Dex-MA T70 DS=13%<br>0.1% m/m LAP</td><td>10 ± 2<br>(17 ± 2; 54.0 ± 5.8)</td><td>38 ± 4<br>(22 ± 3; 37.0 ± 2.5)</td><td>87 ± 8<br>(26 ± 3; 25.7 ± 1.9)</td></tr><tr><td>Dex-MA T20 DS=26%<br>1% m/m LAP</td><td>96 ± 15<br>(41 ± 4; 41.0 ± 2.9)</td><td>237 ± 25<br>(60 ± 6; 26.3 ± 2.1)</td><td>516 ± 13<br>(85 ± 5; 17.8 ± 0.3)</td></tr></tbody></table>

For the swollen hydrogels, significant differences in the mechanical properties were observed. First, a high decrease in the compression moduli values was observed for the wet state compared to the dry state, as expected. For example, for Dex-MA T70 with a DS of 13% at 10% m/m polymer and with 0.1% m/m LAP, the compression modulus was 34 ± 2 MPa at the dry state, whereas it was 10 ± 2 kPa after PBS absorption. For the swollen crosslinked Dex-MA hydrogels, an increase in the LAP concentration from 0.1% to 1% m/m resulted in an increase in the E and UCS values, and a decrease in the US. For example, values of E = 10 ± 2 kPa, UCS = 17 ± 2 kPa, and US = 54 ± 6% were determined for Dex-MA with DS=13% at 10% m/m with 0.1% m/m LAP, whereas they were 238 ± 31 kPa, 57 ± 4 kPa, and 25 ± 1 %, respectively,
for Dex-MA with DS=13% at 10% m/m with 1% m/m LAP. Similarly, an increase of polymer molar mass, concentration, or DS, also yielded more crosslinked swollen polymer networks that exhibited higher compression modulus, higher ultimate compression strength, and reduced ultimate strain. The maximum values of E and UCS were obtained for Dex-MA T70 with a DS of 62% at polymer and LAP concentrations of 30% and 1% m/m, respectively. This Dex-MA material exhibited an E of 8500 ± 120 kPa, a UCS higher than 500 kPa, and a US higher than 6.9%. Experimental compression modulus values were compared with the ones calculated based on the SR measurements according to the formula of Treloar [43] (Table S7). The calculated values were in the same range and showed a similar tendency (i.e. increase with the polymer concentration, polymer DS, and LAP concentration) compared to experimental ones. However, except for the low crosslinked 10% m/m materials with DS=13% and low LAP concentrations, the calculated values were below the experimental ones. This is certainly because of the very short distance between polymer crosslinks that made the model irrelevant (calculated mesh size below the size of one glycoside unit (around 15.5 Å)) for weakly swelling and dense materials.

## 3.4. Chemical and enzymatic degradation of photo-crosslinked Dex-MA materials

The stability of photo-crosslinked Dex-MA materials, prepared with different DS and different polymer and LAP concentrations, was studied in PBS at 37.5°C. The stability of the materials was assessed from day 7 to day 90 of immersion in PBS and the estimations did not account for the contribution of the insoluble fraction (Figure 7). A zero-order kinetic degradation was observed for the Dex-MA T70 polymers with a degradation rate of 0.033 ± 0.002 % per day for the very dense Dex-MA material (Dex-MA with DS=62% at 30% m/m with 1% m/m LAP). For the weakly crosslinked material issued from Dex-MA with DS=13% at 10% m/m with 0.1% m/m LAP, the degradation rate was faster with a value of 0.097 ± 0.007 % per day. Table S8 summarizes the chemical degradation rates that highlighted the effect of the different parameters on the hydrolysis of Dex-MA in PBS during a period of 3 months.
<figure><img src="image_13.png" /><figcaption>**Figure 7.** Long-term stability study of crosslinked Dex-MA T70 materials immersed in PBS at 37.5°C.</figcaption></figure>

Dextranase (1,6-α-d-glucan-6-glucanohydrolase, EC 3.2.1.11) is a very well-known enzyme that catalyzes the endo-hydrolysis of dextran α-1,6-linkages [43,44]. The enzyme was previously shown to be efficient to degrade dextran-based crosslinked hydrogels, provided they were not functionalized with a high DS [44]. The kinetics of the enzymatic degradation of equilibrium-swollen Dex-MA hydrogels at different DS, polymer and LAP concentrations, was also studied (Figure S21). A decrease of the enzymatic degradation activity was observed within 24 hours at 37.5°C. Negligible or very slow (hydrolysis) evolution of the material mass occurred after that time. An efficient enzymatic degradation of Dex-MA materials was obtained for materials where a low LAP concentration of 0.1% m/m was used. For Dex-MA at 30% m/m with DS=13%, a mass loss of 71% was achieved within 24 hours in the presence of dextranase. For Dex-MA at 10% m/m with DS=13%, complete degradation was observed within 24 hours. The enzymatic degradation efficiency decreased when higher LAP concentrations and polymers with a higher DS were used. These results suggested that the most crosslinked materials prevented the enzyme from effectively interacting with the polymer and/or were more structurally resistant and therefore difficult to breakdown. Hydrogels prepared with Dex-MA T20 and displaying shorter polymer chains were slightly more degraded than those with longer chains (Dex-MA
635

T70), suggesting that factors such as the entanglement played a role in the enzymatic
degradation process.

636

637

638

**3.7. Model drug release from Dex-MA hydrogels**

639

Drug release experiments were performed with two model compounds: ketorolac (255 g/mol), a nonsteroidal anti-inflammatory drug used to manage and treat acute moderate to severe pain, and fluorescein-labelled bovine serum albumin (BSA-FITC) (≈ 66,000 g/mol), as a model protein. For each drug, three Dex-MA T70 materials were compared, each displaying a different porosity and swelling ratio: a highly porous material with SR=238% issued from 10% m/m Dex-MA T70 with a low DS of 13% and LAP at 0.1% m/m (noted HP), an intermediate porosity material with SR=94% issued from 20% m/m Dex-MA with a low DS of 13% and LAP at 1% m/m (noted MP for 'medium porosity'), and the densest material with SR=40% issued from 30% m/m Dex-MA with a high DS of 62% and LAP at 1% m/m (noted LP for 'low porosity'). Control experiments were performed to check that the UV process did not degrade the active ingredients. Burst release was observed during the first 2 hours for ketorolac and BSA-FITC (Figure 8). The extent of burst release clearly depended on the hydrogel properties such as porosity, as well as the properties of the drug. The burst release was very important for low molecular mass ketorolac (> 85% of the drug was rapidly released, even for the low porosity (LP) hydrogel). In contrast, only 27.3 ± 1.2% of the BSA-FITC was released after two hours, even for the high porosity (HP) hydrogel. For both drugs, a significantly more progressive 'slow release' phase followed. Overall, a complete release of the ketorolac content was observed within 24 hours for all hydrogels (Figure 8a), whereas only 41.8 ± 1.3% of BSA-FITC diffused out of the high porosity material during the same period (Figure 8b). Finally, a very slow release of BSA-FITC was observed after day 2, which kinetics was compared to the hydrolytic degradation of the hydrogel matrix. BSA-FITC was released from the high porosity hydrogel (resp. from the low porosity hydrogel) at a rate of 0.477 ± 0.024%/day (resp. 0.059 ± 0.007%/day), whereas matrix degradation occurred at 0.097 ± 0.007%/day (resp. 0.033 ± 0.002%/day) (Table S8). It is hypothesized that during this late phase, BSA-FITC release was due probably to both hydrogel hydrolysis and diffusion of the protein from the matrix, facilitated by a more open porous network.

640

641

642

643

644

645

646

647

648

649

650

651

652

653

654

655

656

657

658

659

660

661

662

663

664

665

666
<figure><img src="image_14.png" /><figcaption>**Figure 8.** Kinetics of model drug release from photo-crosslinked Dex-MA materials: **a)** ketorolac; **b)** BSA-FITC. **HP:** high porosity hydrogel (10% m/m Dex-MA T70 DS=13%, LAP 0.1% m/m); **MP:** medium porosity hydrogel (20% m/m Dex-MA DS=13%, LAP 1% m/m); **LP:** low porosity hydrogel (30% m/m Dex-MA DS=62%, LAP 1% m/m).</figcaption></figure>

**3.6. Cytotoxicity assessment**

WST-1 assays were performed on murine fibroblasts incubated in cell culture medium. Crosslinked Dex-MA T70 DS=37% hydrogel samples were incubated for 1, 2 or 7 days. The results evidenced cell viabilities higher than 94.8 ± 1.9% (Figure S22).

**4. Discussion**

A new green Dex-MA synthesis protocol was implemented in aqueous medium that did not require the use of any toxic or carcinogenic species. The reaction presented unprecedented simplicity and efficiency, requiring a reaction time of only 1 hour. The study demonstrated a very high level of control over the new reaction that provided access to Dex-MA polymers with high purity, a wide range of DS values (13 – 62%) using a low number of equivalents of MAh (0.0625 – 0.3 eq.) with very good yields (85 – 92%). The methacrylation reaction was found to be limited by the solubility of the modified polymer in aqueous medium, with a DS threshold around 62%. Since biomedical applications most often require polymer solubility in water, this synthesis protocol limitation does not narrow its high interest to produce Dex-MA polymers with high purity, and free of solvent and toxic residues. The fast and efficient reaction reported here was quite remarkable considering the previously described Dex methacrylation reactions that required around 24 – 48 hours in organic solvents with
added organic bases. Van Dijk-Wolthuis et al. reported the reaction between Dex and glycidyl methacrylate that required DMAP as an organic catalyst, due to the low reactivity of Dex hydroxyl groups in organic solvents [36]. For the methacrylation of Dex using MAh reported by Kim et al. in DMF/LiCl, it was necessary to use a higher number of MAh equivalents to modify the polymer with a reaction completion time of 20 – 30 hours [37]. Our results provide compelling evidence for an improved reactivity of the secondary hydroxyl groups of Dex in aqueous solvent with carefully considered basic conditions, while providing a synthesis method with significantly improved environmental impact (aqueous solvent, under air, room temperature and short reaction time). Other groups have reported the methacrylation in water of other polysaccharides [13,15-16,18-26] but these reactions usually require 12 – 24 hours of reaction time using various equivalents of MAh. None of the reports showed a similar efficacy in terms of DS and reaction time as observed here, which could indicate an intrinsically remarkable reactivity of Dex in water compared to many other polysaccharides. Based on similar results obtained with Dex of different molar mass (T70 and T20), the synthesis protocol developed herein seems to be applicable and efficient to any dextran-based polymer.

The obtained Dex-MA biopolymers were then photo-crosslinked in the dry state in the presence of a cytocompatible photoinitiator (LAP). A dry photo-crosslinking process was selected because it was not possible to obtain non-fractured large pieces of materials after drying when using wet photo-crosslinking. Solid state polymer cross-linking of synthetic and natural polymers such as chitosan and cellulosics using ionizing radiations (i.e. γ-irradiation) has previously been described [45-48]. The direct action of irradiation results in radical formation on the molecular chains. These radicals are then responsible for their cross-linking. However, this direct process can also lead to the degradation of the polymer chains, creating smaller polymer fragments that are able to diffuse out of the material and/or leading to materials with high polydispersity. We report here low soluble polymer mass fractions (i.e. high photo-crosslinking efficiency, > 85%), provided sufficient LAP concentration (1% m/m) was used. Huang et al. studied wet photo-crosslinking of similar Dex-MA hydrogels in the presence of 0.5% m/m LAP [41]. They reported insoluble mass fractions of only about 10% higher than the values reported here. The crosslinking of Dex-MA hydrogels at the solid state appeared slightly less efficient than in water, due to the lower initiator diffusion during the polymerization process; nevertheless, the dry-state crosslinking was still largely
satisfactory and especially when a high concentration of LAP was used, e.g. 1% m/m.
Radical generation and the use of a photosensitizer could also induce some
cytotoxicity of the materials prepared by photo-crosslinking. Dextran-based hydrogels
in general are well acknowledged for their good biocompatibility [27,29]. We selected
LAP as a water-soluble photoinitiator since it has been reported to yield photo-
crosslinked biopolymer-based materials with good biocompatibility while also being a
highly reactive photoinitiator [10-12]. In this work, photo-crosslinked 20% m/m Dex-MA
DS=37% with 1% m/m LAP displayed cell viability > 94% for seven days. This value
was similar to that reported by Liu *et al.* for a 10% m/m Dex-MA (DS=37%) hydrogel
[30], and slightly higher than that reported by Huang *et al.* [41] for a 20% m/m Dex-MA
DS=20% hydrogel crosslinked with 0.2% m/m LAP: we used a higher LAP
concentration, but also a higher methacrylate concentration capable of consuming
photo-produced radicals.

Dried materials obtained after photo-crosslinking appeared very dense, and
mechanically very resistant (compression moduli of 40-340 MPa for Dex-MA T70, 3-
30 MPa for Dex-MA T20). In the dried state, the compression moduli of Dex-MA
depended only on the polymer concentration and molecular mass. The compression
moduli values were independent of the polymer DS and LAP concentration. More
specifically, a decrease in the compression modulus was observed when decreasing
the polymer molecular mass and the concentration of Dex-MA. Thus, the mechanical
properties of the dried materials were intrinsic to the polymers and therefore to the
entanglements and interactions between chains, with a negligible contribution from
inter-chain covalent bonds. The shorter chains of T20 dextran polymers were less
entangled than their longer T70 counterparts, and therefore yielded materials with
lower compression modulus values. This suggested that the mechanical properties of
the materials in the dry state were mainly governed by their enthalpy elasticity,
considering the materials were in an amorphous state and the very high glass transition
temperatures of Dex and Dex-MA [49]. In the swollen state, hydrogel compression
modulus and resistance to compression increased with polymer molar mass, polymer
DS, and polymer and LAP concentrations. The increase of compression modulus and
ultimate compression strength with polymer DS and LAP concentrations evidenced a
contribution of inter-chain covalent bonds (i.e. related to the crosslink density which
increases with these parameters) in the material mechanical properties, contrary to the
dry state. Therefore, swollen hydrogel mechanical properties depended on both
polymer chain entanglements (impacted by polymer molar mass and concentration) and crosslink density, increasing with the probability that a LAP photo-excited state encounters a methacrylate group during photo-crosslinking (i.e. with LAP concentration and polymer DS). The compressive modulus of 30% m/m Dex-MA T70 DS=62% photo-crosslinked with 1% m/m LAP was as high as 8.5 MPa, whereas values of 10 to 200 kPa were achieved with softer hydrogels obtained with a DS of 13%. These later values were similar to those previously reported for crosslinked dextran-based hydrogels. For instance, the 30% m/m Dex-MA (DS=20%) hydrogel reported by Huang et al. displayed a compression modulus of 150 kPa [41], while the 10% m/m Dex-MA (DS=37%) hydrogel reported by Liu et al. exhibited a modulus of 30 kPa [30]. Interestingly, values in the 30-100 kPa range were reported for dextran-based composite hydrogels developed for cartilage [50] and bone [51] tissue repair.
Dried photo-crosslinked materials showed their ability to rapidly absorb aqueous fluids, with swelling equilibrium reached in 1 to 2 hours. Highly tunable SR, from 40 to 238 % for T70 hydrogels, were obtained by changing the polymer and LAP concentrations. The SR values reported here were more than 3 times lower than those obtained for Dex-MA hydrogels photo-crosslinked in the wet state under similar conditions (10 to 30% m/m Dex-MA DS=26%, 0.2% m/m LAP) [41]. This could signify that dry-state crosslinking induced the production of denser materials since the polymer chains in the dry-state were more intimately intertwined and/or stacked in the absence of solvent molecules during the crosslinking process. For high polymer DS, the dense material structure was maintained after swelling: no large macropores were created during liquid uptake (SR = 40% for the more crosslinked hydrogel prepared from 30% m/m Dex-MA DS=62% and LAP 1% m/m). The high DS probably also induced an important hydrophobicity of the Dex-MA polymer chains, and reduced their affinity with water. For lower polymer DS, the polymer chains could expand in contact with water (SR = 238% for hydrogel prepared from 10% m/m Dex-MA DS=13% and LAP 0.1% m/m), creating a porous structure after freeze-drying. Values calculated for the mesh size (16.7 ± 1.0 Å) and crosslinking density (9.45 ± 0.95 mmol cm⁻³) for the hydrogel prepared from 30% m/m Dex-MA DS=13% photo-crosslinked with 1% m/m LAP were in adequation with a previous report describing dextran hydrogels crosslinked in the liquid state (mesh size of 13.2 ± 0.73 Å and crosslinking density of 7.76 ± 0.479 mmol cm⁻³) [52].
The large range of porosity and swelling properties observed for the Dex-MA photo-
crosslinked hydrogels was expected to impact their ability to release drugs entrapped
in the material matrix, through matrix degradation and/or drug diffusion. Dextran is
highly soluble in water and chemically stable, but the ester bonds of the methacrylate
crosslinks between the Dex polymer chains are prone to hydrolysis [44,53]. The
hydrolytic degradation of methacrylate ester crosslinks in aqueous buffer has been
reported in the literature for different methacrylated polysaccharides such as
hyaluronan [15], chondroitin sulfate [15], and dextran derivatives [54]. The release of
two model molecules with low molecular mass (ketorolac) and high molecular mass
(BSA-FITC) was investigated to assess the permeability of the hydrogels and their
release abilities. Ketorolac diffused in a few hours out of the scaffolds. On the contrary,
after a burst phase, BSA-FITC was shown to be released with a prolonged timeline,
both due to the hydrolysis of the matrix and the network-density (i.e. DS) dependent
molecule diffusion.

In conclusion, the Dex-MA polymers obtained in this work, thanks to an original and
easy-to-implement protocol, free of toxic organic solvent and catalyst, yield hydrogel-
based materials highly relevant for biomedical applications. In particular, hydrogels
with strong mechanical properties could be interesting for cartilage and bone tissue
engineering [50-51], or for the design of microneedle devices able to perforate the skin,
then diffuse drugs or uptake interstitial fluid [55].

814 Credit authors

815 **BD:** Conceptualization, Investigation, Data curation, Writing- Original draft. **AG:** Conceptualization, Validation, Writing - review & editing, Supervision, Funding acquisition, Project administration. **AH:** Investigation. **IT:** Conceptualization, Validation, Visualization, Writing- Original draft, Writing - review & editing, Supervision, Funding acquisition, Project administration.

821 Data availability

Data will be made available on request.

824 Acknowledgements

This work was funded by the French National Research Agency in the framework of
the LabEx Arcane (grant ANR-17-EURE-0003) (B. Darmau PhD grant) and supported
827 by Glyco@Alps (ANR-15-IDEX-02) programs. We greatly acknowledge the help of
828 Marta Sacchi (CEA Grenoble/LETI/DTBS and CEA Fontenay/JACOB/SEPIA) for
829 cytotoxicity experiments, Eric Bayma, and Laurine Buon (CERMAV-CNRS) for the
830 GPC analysis, Christine Lancelon-Pin for the SEM images.

831

832

833 References

834 [1] R. Rebelo, M. Fernandes, R. Fangueiro, Biopolymers in Medical Implants: A
835 Brief Review, Procedia Engineering 200 (2017) 236-243.
836 https://doi.org/10.1016/j.proeng.2017.07.034.

837 [2] N. H. Thang, T. B. Chien, D. X. Cuong, Polymer-Based Hydrogels Applied in
838 Drug Delivery: An Overview, Gels 9 (2023) 523.
839 https://doi.org/10.3390/gels9070523.

840 [3] Q. Yang, J. Peng, H. Xiao, X. Xu, Z. Qian, Polysaccharide hydrogels: 841 Functionalization, construction and served as scaffold for tissue engineering, 842 Carbohydrate Polymers 278 (2022) 118952. 843 https://doi.org/10.1016/j.carbpol.2021.118952.

844 [4] J. Nie, B. Pei, Z. Wang, Q. Hu, Construction of ordered structure in 845 polysaccharide hydrogel: A review, Carbohydrate Polymers 205 (2019) 225- 846 235. https://doi.org/10.1016/j.carbpol.2018.10.033.

847 [5] S. Trombino, R. Sole, M. L. Di Gioia, D. Procopio, F. Curcio, R. Cassano, Green 848 Chemistry Principles for Nano- and Micro-Sized Hydrogel Synthesis, Molecules 849 28 (2023) 2107. https://doi.org/10.3390/molecules28052107.

850 [6] C. M. Chakole, P. K. Sahoo, J. Pandey, M. K. Chauhan, A green chemistry 851 approach towards synthesizing hydrogel for sustained ocular delivery of 852 brinzolamide: In vitro and ex vivo evaluation, Journal of the Indian Chemical 853 Society 99 (2022) 100323. https://doi.org/10.1016/j.jics.2021.100323.

854 [7] B. Kaczmarek-Szczepańska, M. Wekwejt, A. Pałubicka, A. Michno, L. Zasada, 855 A. M. Alsharabasy, Cold plasma treatment of tannic acid as a green technology 856 for the fabrication of advanced cross-linkers for bioactive collagen/gelatin 857 hydrogels, International Journal of Biological Macromolecules 258 (2024) 858 128870. https://doi.org/10.1016/j.ijbiomac.2023.128870.

859 [8] S. H. Moon, H. J. Hwang, H. R. Jeon, S. J. Park, I. S. Bae, Y. J. Yang, 860 Photocrosslinkable natural polymers in tissue engineering, Frontiers in 861 Bioengineering and Biotechnology 11 (2023) 1127757. 862 https://doi.org/10.3389/fbioe.2023.1127757.

863 [9] J. P. Fouassier, J. Lalevée, in Photoinitiators for Polymer Synthesis, 2012, pp. 864 3-9. https://doi.org/10.1002/9783527648245.ch1.

865 [10] B. D. Fairbanks, M. P. Schwartz, C. N. Bowman, K. S. Anseth, Photoinitiated 866 polymerization of PEG-diacrylate with lithium phenyl-2,4,6- 867 trimethylbenzoylphosphinate: polymerization rate and cytocompatibility, 868 Biomaterials 30 (2009) 6702-6707. 869 https://doi.org/10.1016/j.biomaterials.2009.08.055.

870 [11] A. K. Nguyen, P. L. Goering, V. Reipa, R. J. Narayan, Toxicity and 871 photosensitizing assessment of gelatin methacryloyl-based hydrogels 872 photoinitiated with lithium phenyl-2,4,6-trimethylbenzoylphosphinate in human
873

874

875

876

877

878

879

880

881

882

883

884

885

886

887

888

889

890

891

892

893

894

895

896

897

898

899

900

901

902

903

904

905

906

907

908

909

910

911

912

913

914

915

916

917

918

919

920

921

922

primary renal proximal tubule epithelial cells, Biointerphases 14 (2019) 021007.
https://doi.org/10.1116/1.5095886.

[12]

H. Xu, J. Casillas, S. Krishnamoorthy, C. Xu, Effects of Irgacure 2959 and lithium phenyl-2,4,6-trimethylbenzoylphosphinate on cell viability, physical properties, and microstructure in 3D bioprinting of vascular-like constructs, Biomed. Mater. 15 (2020) 055021. https://doi.org/10.1088/1748-605X/ab954e.

[13]

F. Araiza-Verduzco, E. Rodríguez-Velázquez, H. Cruz, I. A. Rivero, D. R. Acosta-Martínez, G. Pina-Luis, M. Alatorre-Meda, Photocrosslinked Alginate-Methacrylate Hydrogels with Modulable Mechanical Properties: Effect of the Molecular Conformation and Electron Density of the Methacrylate Reactive Group, Materials 13 (2020) 534. https://doi.org/10.3390/ma13030534.

[14]

S. S. Prado, J. M. Weaver, B. J. Love, Gelation of photopolymerized hyaluronic acid grafted with glycidyl methacrylate, Materials Science and Engineering: C 31 (2011) 1767-1771. https://doi.org/10.1016/j.msec.2011.08.008.

[15]

C. C. L. Schuurmans, A. J. Brouwer, J. A. W. Jong, G.-J. P. H. Boons, W. E. Hennink, T. Vermonden, Hydrolytic (In)stability of Methacrylate Esters in Covalently Cross-Linked Hydrogels Based on Chondroitin Sulfate and Hyaluronic Acid Methacrylate, ACS Omega 6 (2021) 26302-26310. https://doi.org/10.1021/acsomega.1c03395.

[16]

B. S. Spearman, N. K. Agrawal, A. Rubiano, C. S. Simmons, S. Mobini, C. E. Schmidt, Tunable methacrylated hyaluronic acid-based hydrogels as scaffolds for soft tissue engineering applications, Journal of Biomedical Materials Research Part A 108 (2020) 279-291. https://doi.org/10.1002/jbm.a.36814.

[17]

C. Yang, X. Wang, X. Yao, Y. Zhang, W. Wu, X. Jiang, Hyaluronic acid nanogels with enzyme-sensitive cross-linking group for drug delivery, Journal of Controlled Release 205 (2015) 206-217. https://doi.org/10.1016/j.jconrel.2015.02.008.

[18]

H. Qian, X. Wang, K. Yuan, C. Xie, W. Wu, X. Jiang, L. Hu, Delivery of doxorubicin in vitro and in vivo using bio-reductive cellulose nanogels, Biomaterials Science 2 (2014) 220-232. https://doi.org/10.1039/c3bm60176e.

[19]

L. Soullard, P.-A. Bayle, C. Lancelon-Pin, S. Rolere, I. Texier, B. Jean, G. Nonglaton, Optimization of the methacrylation of carboxymethylcellulose and use for the design of hydrogels and cryogels with controlled structure and properties, Cellulose 30 (2023) 6203-6217. https://doi.org/10.1007/s10570-023-05266-w.

[20]

L.-F. Wang, S.-S. Shen, S.-C. Lu, Synthesis and characterization of chondroitin sulfate-methacrylate hydrogels, Carbohydrate Polymers 52 (2003) 389-396. https://doi.org/10.1016/S0144-8617(02)00328-4.

[21]

H. Chen, R. Li, J. Xie, X. Xu, S. Chen, Facile methacrylation of cellulose via alkaline aqueous esterification for thiol-ene click functionalization, Materials Letters 245 (2019) 18-21. https://doi.org/10.1016/j.matlet.2019.02.083.

[22]

P. V. Kelly, P. Cheng, D. J. Gardner, W. M. Gramlich, Aqueous Polymer Modification of Cellulose Nanofibrils by Grafting-Through a Reactive Methacrylate Group, Macromolecular Rapid Communications 42 (2021) 2000531. https://doi.org/10.1002/marc.202000531.

[23]

L. Soullard, F. Pradalié, B. Labat, C. Lancelon-Pin, G. Nonglaton, S. Rolere, I. Texier, B. Jean, Methacrylated Cellulose Nanocrystals as Fillers for the Development of Photo-Cross-Linkable Cytocompatible Biosourced Formulations Targeting 3D Printing, Biomacromolecules 24 (2023) 6009-6024. https://doi.org/10.1021/acs.biomac.3c01090.
923

924

925

926

927

928

929

930

931

932

933

934

935

936

937

938

939

940

941

942

943

944

945

946

947

948

949

950

951

952

953

954

955

956

957

958

959

960

961

962

963

964

965

966

967

968

969

970

971

[24]

C. Noè, C. Tonda-Turo, A. Chiappone, M. Sangermano, M. Hakkarainen, Light Processable Starch Hydrogels, Polymers 12 (2020) 1359. https://doi.org/10.3390/polym12061359.

[25]

P. Joshi, M. S. U. Ahmed, K. Vig, I. B. Vega Erramuspe, M. L. Auad, Synthesis and characterization of chemically crosslinked gelatin and chitosan to produce hydrogels for biomedical applications, Polymers for Advanced Technologies 32 (2021) 2229-2239. https://doi.org/10.1002/pat.5257.

[26]

O. M. Kolawole, W. M. Lau, V. V. Khutoryanskiy, Methacrylated chitosan as a polymer with enhanced mucoadhesive properties for transmucosal drug delivery, International Journal of Pharmaceutics 550 (2018) 123-129. https://doi.org/10.1016/j.ijpharm.2018.08.034.

[27]

A. Luanda, V. Badalamoole, Past, present and future of biomedical applications of dextran-based hydrogels: A review, International Journal of Biological Macromolecules 228 (2023) 794-807. https://doi.org/10.1016/j.ijbiomac.2022.12.129.

[28]

A. Delvart, C. Moreau, B. Cathala, Dextrans and dextran derivatives as polyelectrolytes in layer-by-layer processing materials – A review, Carbohydrate Polymers 293 (2022) 119700. https://doi.org/10.1016/j.carbpol.2022.119700.

[29]

E. Díaz-Montes, Dextran: Sources, Structures, and Properties, Polysaccharides 2 (2021) 554-565. https://doi.org/10.3390/polysaccharides2030033.

[30]

Z. Q. Liu, Z. Wei, X. L. Zhu, G. Y. Huang, F. Xu, J. H. Yang, Y. Osada, M. Zrínyi, J. H. Li, Y. M. Chen, Dextran-based hydrogel formed by thiol-Michael addition reaction for 3D cell encapsulation, Colloids and Surfaces B: Biointerfaces 128 (2015) 140-148. https://doi.org/10.1016/j.col surfb.2015.02.005.

[31]

D. Li, J. Han, J. Ding, L. Chen, X. Chen, Acid-sensitive dextran prodrug: A higher molecular weight makes a better efficacy, Carbohydrate Polymers 161 (2017) 33-41. https://doi.org/10.1016/j.carbpol.2016.12.070.

[32]

N. Yadav, A. P. Francis, V. V. Priya, S. Patil, S. Mustaq, S. S. Khan, K. J. Alzahrani, H. J. Banjer, S. K. Mohan, U. Mony, R. Rajagopalan, Polysaccharide-Drug Conjugates: A Tool for Enhanced Cancer Therapy, Polymers 14 (2022) 950. https://doi.org/10.3390/polym14050950.

[33]

O. Franssen, R. D. van Ooijen, D. de Boer, R. A. A. Maes, W. E. Hennink, Enzymatic Degradation of Cross-Linked Dextrans, Macromolecules 32 (1999) 2896-2902. https://doi.org/10.1021/ma981759m.

[34]

P. Edman, B. Ekman, I. Sjöholm, Immobilization of proteins in microspheres of biodegradable polyacryldextran, Journal of Pharmaceutical Sciences 69 (1980) 838-842. https://doi.org/10.1002/jps.2600690725.

[35]

W. N. E. van Dijk-Wolthuis, O. Franssen, H. Talsma, M. J. van Steenbergen, J. J. Kettenes-van den Bosch, W. E. Hennink, Synthesis, Characterization, and Polymerization of Glycidyl Methacrylate Derivatized Dextran, Macromolecules 28 (1995) 6317-6322. https://doi.org/10.1021/ma00122a044.

[36]

W. N. E. van Dijk-Wolthuis, J. J. Kettenes-van den Bosch, A. van der Kerk-van Hoof, W. E. Hennink, Reaction of Dextran with Glycidyl Methacrylate: An Unexpected Tranesterification, Macromolecules 30 (1997) 3411-3413. https://doi.org/10.1021/ma961764v.

[37]

S.-H. Kim, C.-C. Chu, Synthesis and characterization of dextran–methacrylate hydrogels and structural study by SEM, Journal of Biomedical Materials Research 49 (2000) 517-527. https://doi.org/10.1002/(SICI)1097-4636(20000315)49:4<517::AID-JBM10>3.0.CO;2-8.
972 [38] K. Malzahn, W. D. Jamieson, M. Dröge, V. Mailänder, A. T. A. Jenkins, C. K. Weiss, K. Landfester, Advanced dextran based nanogels for fighting Staphylococcus aureus infections by sustained zinc release, Journal of Materials Chemistry B 2 (2014) 2175-2183. https://doi.org/10.1039/c3tb21335h.

975
976 [39] W. N. E. van Dijk-Wolthuls, S. K. Y. Tsang, J. J. Kettenes-van den Bosch W. E. Hennink, A new class of polymerizable dextrans with hydrolyzable groups: hydroxyethyl methacrylated dextran with and without oligolactate spacer, Polymer 38 (1997) 6235-6242. https://doi.org/10.1016/S0032-3861(97)00189-4.

981 [40] J. Hoque, J. Haldar, Direct Synthesis of Dextran-Based Antibacterial Hydrogels for Extended Release of Biocides and Eradication of Topical Biofilms, ACS Applied Materials & Interfaces 9 (2017) 15975-15985. https://doi.org/10.1021/acsami.7b03208.

982
983 [41] S. Huang, H. Liu, S. Huang, T. Fu, W. Xue, R. Guo, Dextran methacrylate hydrogel microneedles loaded with doxorubicin and trametinib for continuous transdermal administration of melanoma, Carbohydrate Polymers 246 (2020) 116650. https://doi.org/10.1016/j.carbpol.2020.116650.

984
985 [42] N. A. Peppas, S. L. Wright, Solute Diffusion in Poly(vinyl alcohol)/Poly(acrylic acid) Interpenetrating Networks, Macromolecules 29 (1996) 8798-8804. https://doi.org/10.1021/ma9613392.

989 [43] L. R. G. Treloar, The elasticity and related properties of rubbers, Reports on Progress in Physics 36 (1973) 755-826. https://doi.org/10.1088/0034-4885/36/7/001.

992
993 [44] T. K. L. Meyvis, S. C. De Smedt, J. Demeester, W. E. Hennink, Influence of the Degradation Mechanism of Hydrogels on Their Elastic and Swelling Properties during Degradation, Macromolecules 33 (2000) 4717-4725. https://doi.org/10.1021/ma992131u.

998
999 [45] R. Gottlieb, C. Kaiser, U. Gohs, K.-F. Arndt, Temperature Sensitive Hydrogels Based on Hydroxypropylcellulose by High Energy Irradiation, Macromolecular Symposia 254 (2007) 361-369. https://doi.org/10.1002/masy.200750852.

1000
1001 [46] M. Monier, Y. Wei, A. A. Sarhan, D. M. Ayad, Synthesis and characterization of photo-crosslinkable hydrogel membranes based on modified chitosan, Polymer 51 (2010) 1002-1009. https://doi.org/10.1016/j.polymer.2010.01.001.

1005 [47] I. Silvestro, R. Sergi, A. Scotto D'Abusco, A. Mariano, A. Martinelli, A. Piozzi, I. Francolini, Chitosan scaffolds with enhanced mechanical strength and elastic response by combination of freeze gelation, photo-crosslinking and freeze-drying, Carbohydrate Polymers 267 (2021) 118156. https://doi.org/10.1016/j.carbpol.2021.118156.

1010 [48] R. A. Wach, H. Mitomo, F. Yoshii, ESR investigation on gamma-irradiated methylcellulose and hydroxyethylcellulose in dry state and in aqueous solution, Journal of Radioanalytical and Nuclear Chemistry 261 (2004) 113-118. https://doi.org/10.1023/b:jrnc.0000030943.48622.33.

1014 [49] C. Asgreen, M. M. Knopp, J. Skytte, K. Löbmann, Influence of the Polymer Glass Transition Temperature and Molecular Weight on Drug Amorphization Kinetics Using Ball Milling, Pharmaceutics 12 (2020) 483. https://doi.org/10.3390/pharmaceutics12060483.

1018 [50] Y. Fu, B. Zoetebier, S. Both, P. J. Dijkstra, M. Karperien, Engineering of Optimized Hydrogel Formulations for Cartilage Repair, Polymers 13 (2021) 1526. https://doi.org/10.3390/polym13091526.
[51] X. Zhang, K. Nan, Y. Zhang, K. Song, Z. Geng, D. Shang, X. Guan, L. Fan, A novel injectable hydrogel prepared from phenylboronic acid modified gelatin and oxidized-dextran for bone tissue engineering, International Journal of Biological Macromolecules 261 (2024) 129666.
https://doi.org/10.1016/j.ijbiomac.2024.129666.

[52] N. Paradee, J. Thanokiang, A. Sirivat, Conductive poly(2-ethylaniline) dextran-based hydrogels for electrically controlled diclofenac release, Materials Science and Engineering: C 118 (2021) 111346.
https://doi.org/10.1016/j.msec.2020.111346.

[53] B. G. Stubbe, W. E. Hennink, S. C. De Smedt, J. Demeester, Swelling Pressure of Hydrogels That Degrade through Different Mechanisms, Macromolecules 37 (2004) 8739-8744.
https://doi.org/10.1021/ma040103s.

[54] W. N. E. van Dijk-Wolthuis, M. J. van Steenbergen, W. J. M. Underberg, W. E. Hennink, Degradation Kinetics of Methacrylated Dextrans in Aqueous Solution, Journal of Pharmaceutical Sciences 86 (1997) 413-417.
https://doi.org/10.1021/js9604220.

[55] B. Darmau, M. Sacchi, A. Gross, I. Texier, Self-Extracting Dextran-Based Hydrogel Microneedle Arrays with an Interpenetrating Bioelectro catalytic Sensor for Transdermal Monitoring with Matrix Protection, Advanced Healthcare Materials, in revision.
1043 **Figure legends**

1044

1045 **Figure 1.** Dextran methacrylation in water at room temperature in air. **a)** Reaction scheme. **b)** Time-evolution of Dex T70 DS methacrylation in water using different eq. of MAh. **c)** Time-evolution of unreacted (circles, red line) and hydrolyzed (triangles, black line) MAh during the methacrylation of Dex T70 using 0.125 eq of MAh.

1049

1050 **Figure 2.** ¹H NMR spectra in D₂O at 300 MHz of dextran T70 (**a**), and Dex-MA T70 with DS=13% (**b**), DS=26% (**c**), DS=52% (**d**), DS=62% (**e**).

1051

1052

1053 **Figure 3.** Preparation of Dex-MA hydrogels. **a)** Photo-chemical crosslinking reaction at the dried state; **b)** Scheme of the crosslinked Dex-MA hydrogel preparation process and **c)** photographs of material samples prepared from Dex-MA DS=13% 10% m/m solutions with 0.1% m/m LAP before (i) and after (ii) photo-crosslinking at the dried state, and after 1 hour swelling in PBS (iii).

1054

1055

1056

1057

1058

1059 **Figure 4.** Insoluble mass fractions (IM) at the equilibrium for Dex-MA (T70) materials: **a)** influence of the polymer DS and concentration at a LAP concentration of 1% m/m; **b)** influence of the LAP concentration (m/m) for Dex-MA (T70) materials with DS=13%.

1060

1061

1062

1063 **Figure 5.** Swelling ratios (SR) at the equilibrium recorded after 24 hours in PBS for Dex-MA (T70) materials: **a)** influence of the polymer DS and concentration at a LAP concentration of 1% m/m; **b)** influence of the polymer and LAP concentration for Dex-MA DS=13%.

1064

1065

1066

1067

1068 **Figure 6.** Cross-section SEM micrographs of dried Dex-MA materials after UV-crosslinking (**a**), **b**), scale bar 200 µm), and after PBS swelling then freeze-drying (swollen state) (**c**), **d**), scale bar 200 µm). (**a**), (**c**) Materials issued from 10% m/m solutions of Dex-MA T70 with DS=13% and 0.1% m/m of LAP. (**b**), **d**) Materials issued from 30% m/m solutions of Dex-MA T70 with DS = 62% and 1% m/m of LAP.

1069

1070 **1070** Solutions of Dex-MA T70 with DS=13% and 0.1% m/m of LAP. (b), (d)) Materials issued from 30% m/m solutions of Dex-MA T70 with DS = 62% and 1% m/m of LAP.

1071

1072

1073

1074 **Figure 7.** Long-term stability study of crosslinked Dex-MA T70 materials immersed in PBS at 37.5°C.

1075

1076
**Figure 8.** Kinetics of model drug release from photo-crosslinked Dex-MA materials: **a)** ketorolac; **b)** BSA-FITC. **HP:** high porosity hydrogel (10% m/m Dex-MA T70 DS=13%, LAP 0.1% m/m); **MP:** medium porosity hydrogel (20% m/m Dex-MA DS=13%, LAP 1% m/m); **LP:** low porosity hydrogel (30% m/m Dex-MA DS=62%, LAP 1% m/m).

1081
1082