A review on the construction of hydrogel scaffolds by various chemically techniques for tissue engineering

European Polymer Journal 117 (2019) 64–76

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A review on the construction of hydrogel scaffolds by various chemically techniques for tissue engineering Parinaz Nezhad-Mokhtaria,b,1, Marjan Ghorbanic,

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, Leila Roshangarc,d, , Jafar Soleimani Radc,d,

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a

Polymer Research Laboratory, Department of Organic and Biochemistry, Faculty of Chemistry, University of Tabriz, Tabriz, Iran Stem Cell And Regenerative Medicine Institute, Tabriz University of Medical Sciences, Tabriz, Iran c Stem Cell Research Center, Tabriz University of Medical Sciences, Tabriz, Iran d Department of Anatomical Sciences, Faculty of Medicine, Tabriz University of Medical Sciences, Tabriz, Iran b

A R T I C LE I N FO

A B S T R A C T

Keywords: Hydrogels Click chemistry Michael-type addition reaction Schiff base-crosslinking Tissue engineering

The field of tissue engineering places complex demands on the natural and synthetic biomaterials that have been applied for organ/tissue development. The materials selected to support the complicated processes of tissue development and repairs which assist to the structural and mechanical requirements of the target tissue. Hydrogels have in situ formability which permits an actual and homogeneous encapsulation of drugs/cells, and useful in vivo surgical operation in a minimally invasive way, causing less discomfort for patients. This review provides an overview on the current chemically gelling techniques, with an explanation of the state of the art on the various synthesis methods and biomedical applications of chemically gelling hydrogels, contrasting and comparing the many chemistries available for biomedical hydrogel generation using both biologic and synthetic base materials. So, we discuss about the versatile methods to design the hydrogels including click chemistry, Michael-type addition reaction and Schiff base reaction.

1. Introduction Scaffolds, including hydrogel networks, are typically produced and then seeded by cells and attach to the inner walls of scaffolds’ cavities. Hydrogels can offer mechanical reinforcement as well as enhancement of biological activity of hydrogels via extracellular matrix (ECM) mimicking. Hydrogels are 3 dimensional (3D) constructs designed with hydrophilic polymers, in which the polymer chains were cross-linked with a chemical or physical technique to form a network. The most commonly used classifications of hydrogels are associated with the nature of chemical cross-linking, because the resulting hydrogel is generally more resistant to mechanical forces but undergoes greater volume changes than physically crosslinked networks. Generally, this junctions are permanent as a result of several chemical reactions, such as Michael addition [1], Schiff-base reaction [2], and azide-alkyne cycloaddition reactions [3] (click chemistry) (Fig. 1). Schiff base reaction between (aldehyde) AD groups and amine groups was also applied to produce reversible chemically crosslinked hydrogels for synthesizing injectable hydrogels [4–7]. The hydrogels have been interested owing to the high reaction rate and mild reaction situations, as well as the capability to form imine bonds between AD

and amino groups without any additional reagents or external stimuli. Chitosan (CS) is one of excellent natural polymers for preparing hydrogels through the Schiff base crosslinking, due to the presence of amino groups on its backbone. For example, Ye et al. described an in situ biodegradable hydrogel by self-cross-linking of water-soluble CS and oxidized sodium Alginate (Alg). Schiff base reaction was performed between the amino groups of CS and oxidized Algs which could potentially function as biodegradable and non-toxic crosslinking agents in the preparation of hydrogels under physiological condition [8]. In another example, Liu et al. utilized methacrylate- and AD-bifunctionalized dextran (Dex–MA–AD) and revealed a novel cell-encapsulating hydrogel family based on the interpenetrating network of gelatin and Dex–MA–AD [9]. The hydrogel is formed via UV-crosslinking between pendant methacrylate groups on Dex–MA–AD and Schiff base reaction between the AD groups of Dex-MA-AD and the amino groups of gelatin. In addition, other hydrogels coupled by Schiff base cross-linking have been mostly examined. In recent times, Honglin et al, introduced the active groups (AD groups) on the surface of bacterial cellulose (BC) nanofibers with 2,2,6,6-tetramethylpyperidine-1-oxy radical (TEMPO)mediated oxidation and evaluated the potential of the TEMPO-oxidized BC as tissue engineering scaffolds [10]. Click chemistry is one of chemo-

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Corresponding authors. E-mail addresses: [email protected] (M. Ghorbani), [email protected] (L. Roshangar), [email protected] (J. Soleimani Rad). 1 These authors contributed equally to this work. https://doi.org/10.1016/j.eurpolymj.2019.05.004 Received 26 March 2019; Received in revised form 25 April 2019; Accepted 5 May 2019 Available online 06 May 2019 0014-3057/ © 2019 Published by Elsevier Ltd.

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Fig. 1. Different synthesis methods to construct the hydrogels by various chemically techniqe.

stem cells to differentiate and the formation of cartilage ECM before the degradation of gel, the gels could be stable for at least 3 weeks. For this purpose, they could influence the swelling behavior and degradation time of the Dex-PEG hydrogel by varying the DS of Dex, resulting in the degradation times of 3–7 weeks for Dex-based hydrogels with increasing thiol groups. Jin et al. reported the synthesis of hydrogels through the Michael-type addition of thiolated hyaluronic acid (HA-SH) and 4-arm PEGvinylsulfone (PEG-4VS) [27]. The authors motivated on significant factors including enzymatic degradation profile, mechanical moduli, and gelation rate. In this content, the researchers observed that the degradation time enhanced from 3 to 15 days when the polymer concentration was enhanced in three folds. After 21 days, collagen II (Col II) and glycosaminoglycans (GAG) were initiated to have accumulated in hydrogels. Another kind of hydrogel constructed by Dithiothreitol (DTT) as a crosslinker/extender was synthesized towards the self-assembly of α-cyclodextrins (CDs) with acryloyl end capped 3arm PEG [28]. Due to the increasing in the feeding number of α-CDs per end-acrylated 3-arm PEG, the longer time required for the conjugate addition. The partial crosslink resulting from the rapid Michael-type addition reaction likely also gave rise to the lower gel content values as long as more α-CDs were added. Another example of in situ cross-linked hydrogel was shown by Teng et al. [29]. In this work, a novel hydrogel was prepared by thiol-modified CS (CS–NAC) and PEG diacrylate (PEGDA). They found that the gelation time depended on the content of free thiols in CS–NAC, concentration of CS-NAC and PEGDA, and temperature. Rheological studies showed that the cross-linking density and elasticity of hydrogel were raised by increasing the content of CS–NAC and PEGDA. Moreover, the result showed that the swelling was extremely temperature-dependent and was directly related to the amount of cross-linking. Biological activities of the hydrogels were assessed and the results specified that the hydrogel was biocompatible and degradable under physiological conditions which present great potential in biomedical applications such as drug delivery. The cell viability in hydrogel constructed between CS-NAC and PEGDA was a little lower than that of hydrogel fabricated by disulfide-crosslinked hydrogel which might be due to the rapider degradation of hydrogel.

selective crosslinking strategies to produce controlled crosslinking of hydrogels in situ without toxic additives under physiological environments [11,12]. These reactions have exposed great potential for the construction of hydrogels, because of their fast polymerization kinetics and low reactivity with cellular components. For example, Feng et al, developed a multi-functional hydrogel with self-healing, good structure integrity, and tissue-adhesive property formed by combining acylhydrazone bond and DA click reaction. The hydrogel physicochemical properties displayed outstanding elasticity, shape recovery and antifatigue property. Another commonly used approach to formulate hydrogels is the Michael addition reaction, which is the nucleophilic addition of a carbanion to an α,β-unsaturated carbonyl compound [13–15]. For example, Mariarosaria et al. have reported the nucleophilic thiol-ene addition between the thiol-capped polyvinyl alcohol (PVA) chains and the vinyl moiety enclosed in a grafted methacrylic PVA as a probable candidate for surgical applications [16]. Although several journal reviews and articles on hydrogels have been published for tissue engineering, this is the first review that particularly focuses on various chemically gelling for developing hydrogels, for use in tissue engineering. 2. Various chemical synthesis methods of hydrogels for tissue engineering 2.1. Hydrogels constructed by Michael-type addition reaction Michael-type addition in the meaning of tissue engineering refers to the addition of a nucleophile (typically a deprotonated thiol) onto a double bond conjugated to an electron withdrawing group including a carbonyl (typically vinyl sulfone, acrylate, or maleimide) [17–21]. It is also known as conjugate addition or 1,4-nucleophilic addition [22–25]. In this content, Jukes et al. explored the use of thiol-functionalized Dex with a tetra-acrylated polyethylene glycol (PEG) for the production of Dex-based hydrogels [26]. They demonstrated that the degradation time of synthesized hydrogels could be changed from 3 to 7 weeks via altering the degree of substitution of thiol groups. To allow less devoted 65

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growth factor released from MG5 was faster than that released from MG10. This result was achieved in a low concentration of BMP-2 remaining in MG5, while MG15 reserved BMP-2 as a sustained release in the MGs environment; due to the high crosslinking density maintained their structures (Fig. 2C). Silva et al. designed the gellan gum (GG) hydrogels through the functionalization of divinyl sulfone (DVS) and then formed by a dithiol peptide crosslinker sensitive to matrix metalloproteinases (T1 and C16, both existing in the ECM) [32]. The results showed that the mechanical properties, swelling and MMP-1-driven degradation were adjusted by increasing the polymer amount and crosslinking density. For all the hydrogels the amount of viable cells increased from day 3, representing that cells were proliferating. In another example, Jain et al. focused on the thiol-functionalized crosslinkers on PEG hydrogel gelation and degradation kinetics [33]. The fast-gelling crosslinkers had gelation times in the following order (from fast to slow): Glyceryl dithioglycolate (GDT) > Glycol di mercaptoacetate (GDMA) > PEG-diesterdithiol-1 (PEGDD-1) > (S)-2 aminobutane 1,4 dithiol (DTBA). Between slowgelling crosslinkers, PEG-based dithiol crosslinkers were the slowest to form gel. Gelation time depended on crosslinker MW and ester presence. Moreover, mesh size could be controlled via varying the dithiol crosslinker MW and chemical structure. Among hydrogels constructed with ester-containing crosslinkers, hydrogels prepared by crosslinkers through 1 methylene spacer degraded in the following order: GDT < GDMA < PEGDD-1. It should be noted, when the methylene spacer increased, the time of degradation was also improved. The cell viability studies showed the maximum viability when encapsulated in hydrogels with high MW and the lowest viability of when encapsulated in hydrogels with low MW crosslinkers. Rensburg et al. reported the synthesis of a degradable (Type D) and non-degradable (Type N) Heparin (HEP)- and HEP sulphate (HS)-containing PEG hydrogels [34]. For this purpose, HEP and HS were modified with acrylation incorporated into PEG gels via nucleophilic addition mechanism. Gels were assessed to determine the Hep/HS release kinetics from the gels, and the capability of the gels to release VEGF. Sustained release of covalently conjugated (CC) HEP from Type D (21 days) and Type N (56 days) was reached, while the noncovalently incorporated (NI) HEP that eluted within a

Chen et al. reported the facile formation of microporous hydrogels based on self-crosslinking reactions between the maleimide and amine on the modified CS. This method produced no toxic byproducts, and did not need any toxic initiator/catalyst, or irradiation. Even though the click chemistry between thiol and malemido groups is too fast to be controlled for the formation of hydrogels, so Chen et al. were the first group to make use of the Michael-type addition reaction between amino and malemido groups which was more tunable. They also tuned the mechanical property of scaffolds by altering the concentration of CS and/or the degree of maleimidyl modification and their swelling behaviors were in a good agreement. In this research, the modified CS presented the good biocompatibility and immunocytochemistry showed that the porous structures could promote osteogenesis of hAMSCs. In another work, Yu et al. developed glycol CS-based (GCS) hydrogels by the reaction between thiolated glycol CS (GCS-SH) and oligo (acryloyl carbonate)-b-PEG-b-oligo(acryloyl carbonate) (OAC-PEG-OAC) copolymers [30]. By DS of thiolated GCS (GCS-SH), polymer concentration, pH, the gelation times and storage moduli of hydrogels could be organized. These GCS hydrogels had slow hydrolytic degradation, and microporous structures which is steady for over 6 months. The in vivo degradation of GCS-SH/OAC-PEG-OAC hydrogels showed several weeks to several months required to complete degradation. pH-degradable hydrogel was designed with a microfluidic technology for bone tissue engineering applications [31]. For this purpose, the PVAbased microgels (MGs) were fabricated from vinyl ether acrylate-functionalized PVA (PVA-VEA) and thiolated PVA-VEA (PVA-VEA-SH). Then, it could be co-loaded by bone marrow human mesenchymal stem cells (hMSCs) and growth factor to protect the stem cells. Moreover, the long-term sustained release of bone morphogenetic protein-2 (BMP-2) by PVA MGs increased osteogenic variation in vitro. By increasing the initial polymer concentrations of MGs (MG5, MG10 and MG15), the MGs presented slight swelling for MG10 and MG15 whereas for a concentration of 5 wt% (MG5), the MGs swelled up more (Fig. 2A). Moreover, the young’s modulus of the microdroplets was in agreement with the results of SEM images. The PVA-MGs demonstrated that the viability of hMSCs encapsulated in the MGs was all above 80% after 7day incubation period for all the three types of MGs (Fig. 2B). BMP-2

Fig. 2. SEM images of poly vinyl ether acrylate (PVA) microgels formed with different polymer concentrations after lyophilization (A), Effects of the crosslinking density of PVA microgels on cell viability and proliferation (B), effect of the crosslinking density on in vitro bone morphogenetic protein-2 (BMP-2) release from PVA microspheres (mean ± SD, n = 3) (C), Reproduced from [39]. Copyright (2018) Elsevier Ltd. 66

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Fig. 3. The dependence of gelation time on temperature in different Diels–Alder (DA) click reaction formed hyaluronic acid/polyethylene glycol (HA/PEG) hydrogels (a), the compressive modulus of the hydrogels (b), and fluorescence microscope images of live (green), dead (red) and merged images of cells encapsulated in HA/ PEG hydrogels. (A–C) Cells encapsulated in hydrogel DS3 3:1 which had undergone a 6 h gelation time (GT 6 h). (C–E) Cells encapsulated in hydrogel DS2 3:1 which had undergone a 3 h gelation time (GT 3 h). (G–I) Cells encapsulated in hydrogel DS2 1:1 which had undergone a 1 h gelation time (GT 1 h). Degree of subsituation (DS). Reproduced from [49]. Copyright (2014) Royal Society of Chemistry. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

burst release in the first two days. Furthermore, 2.7% of the incorporated growth factor was released from control gels during the first day, while, at day 12, 11.3% of the total amount of incorporated growth factor was released. The release rate of VEGF from PEG-Acrylate (PEGAc)/SH/Hep-Ac and PEG-Ac/SH/HS-Ac was lower and higher than the control during 1–12 days. To end, when the growth factor was delivered from covalently heparinzed gels, the highest levels of vascularization was achieved. Shikanov et al. designed trifunctional cross-linking peptides used to 4-arm PEG hydrogels [35]. The PEG macromer was a 4-arm star with 4 acrylate groups, whereas the peptide in the form of cysteine residues had three cross-linking sites, creating a 4:3 network. The 2:3 network formed by linear PEG-acrylate precursor (Mw: 3.5 kDa) formed hydrogels in low concentration. While, a linear PEGacrylate molecule by a greater Mw of 10 kDa did not produce a hydrogel under any conditions. By comparing the 4:3 with 4:2 network, the swelling was decreased as the PEG concentration was enhanced. The 4:3 network permitted the successful encapsulation of the ovarian follicles and follicles survived in 10 days of culture. The formation of

hydrogels via the step-growth polymerization of thiol-and maleimideterminated PEG macromers designed by Fu et al. [36]. In this work, thiol-maleimide (SH-Mal) was used in the formation of a physically entangled PEG network. The gel formation was not due to the formation of cross-links but reasonably due to the physical chain entanglements, which makes it an interesting system different from classical crosslinked PEG hydrogels. The results of swelling tests also presented that these hydrogels are highly elastic. Moreover, by increasing the polymer concentration, the higher G’ obtained. In another works, catecholfunctionalized CS (CHIC) was cross-linked by terminally thiolated Pluronic F-127 (Plu-SH) triblock copolymer to make hydrogels. The adhesive CS/Plu hydrogels exposed strong adhesiveness to mucous layers and soft tissues [37]. Unlike the unmodified CS, the CHIC was soluble in neutral, basic, and acidic solutions due to the presence of multiple catechol groups changed the overall pKa of primary amine groups. The change in elastic modulus values of CHIC and CHIC/Plu-SH specified that the addition of Plu could not only decreased the gelation time, but also improved the gel strength because of the intermolecularly 67

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Fig. 4. Fabrication of bio/synthetic interpenetrating network (BioSIN) of gelatin methacrylamide and poly(ethylene glycol) via simultaneous photoinitiation of thiolene and thiolyne coupling (A), and 2D endothelial cell adherence on hydrogel monoliths of different macromer composition and crosslinking modality (B). Reproduced from [55]. Copyright (2014) Elsevier Ltd.

reactions, by emphasis on tunable properties, and biomedical applications. In addition, DA click chemistry for hydrogel formation has recently considered, because of the efficient chemical bonding and the mild biological reaction conditions [39]. However, the long gelation time of DA designed hydrogel is a big problem for cell encapsulation and limits its application in the cytobiology field. Since HA had been verified to preserve chondrocyte viability, promoting and supporting the chondrogenic differentiation of MSCs, so HA was selected as a main appropriate component to encapsulate the cells. In this content, Yu et al. designed HA/PEG hydrogels with high elasticity for articular cartilage tissue repair [40]. The series of HA/PEG hydrogels with molar ratios of 1:1 and 3:1 showed suitable gelation times for proliferation and survival. As shown in Fig. 3b, by simply tuning the molar ratio of functional group, the value of the compressive modulus was controlled and the gelation time could be tuned (Fig. 3a). As revealed in Fig. 3cA and 3cC, because of the lack of nutrient supply and the long gelation time, all the cells died (Fig. 3cB) after 6 h, while some dead cells and some living cells could be seen in the gelation time 3 h hydrogels (Fig. 3cD and cE). After 1 h (Fig. 3cH), the number of dead cells was almost minor compared with the gelation time 3 h and 6 h. Therefore, a gelation time of around 1 h was appropriate for cell viability, chondrogenesis and proliferation. In another work, Wang et al. studied injectable Dex-based hydrogels prepared via metal-free click chemistry [41]. The non-toxic azide-alkyne cycloaddition was used for the first time with azide-modified Dex (Dex-N3) and azadibenzocyclooctyne-modified Dex (Dex-ADIBO). Rheological analysis showed that the Dex hydrogels were elastic and

crosslinked CS and Plu molecules. Moreover, CHI-C/Plu-SH hydrogels displayed good stabilities in comparison to CHIC alone. The hydrogels was injected into the subcutaneous region of mice and the results showed that the CHI-C/Plu-SH hydrogels after 25 maintained 75% of their original dry weight, respectively. Kim et al. explored that enhancing PEG functionality from 4 to 8-arms affected the kinetics of hydrogel formation [38]. They evaluated 4 and 8-arm PEG vinyl sulfone (PEG-VS) hydrogels with YKNR peptide and modified by changing concentration of L-Cys. The degrees of G’ and swelling of 8-arm PEG were less affected in contrast to 4-arm PEG. The gel formation rate of the 8-arm PEG hydrogel was higher than those of 4-arm PEG hydrogel. In PEG hydrogels by L-Cys functionalization, 4-arm PEG hydrogels had a swelling ratio 1.5 times more than 8-arm PEG hydrogels, but when no L-Cys was added, there was no important alteration in the swelling ratio in 8-arm PEG hydrogels. PVA-based hydrogel was constructed by Tortora et al. [16]. In this work, the reaction between thiol-capped PVA chains and the vinyl moiety contained in a grafted methacrylic PVA (PVA-MA) was studied. The conjugation of thiol groups by PVA aldehydes was obtained via a Schiff reaction by the amine group of cysteine. 2.2. Hydrogels constructed by click chemistry reaction Click reactions that comprise of selective reactions under mild conditions have provided real stimulus not only in manufacturing elegant bioactive materials of choice but also in forming the leap to industrial scale build-up of multifunctional products. In this section, we review several types of polymeric scaffolds fabricated through “click” 68

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Fig. 5. Schematic illustration of Cys-Ala-Gly-peptide (CAG)-peptide functionalized polycarbonate urethane (PCU)-film surfaces (A), and endothelial cells (EC) adhesion and proliferation on different functionalized PCU-film surfaces for 1, 3 and 7 day/s (B). Reproduced from [64]. Copyright (2015) Elsevier Ltd.

superficial zone and resist compressive forces in the deep zone for cell adhesion and differentiation. Daniele et al. described an efficient and simple approach for the fabrication of a bio/synthetic interpenetrating network (BioSINx) [45]. The BioSINx was comprised of a gelatin methacrylamide (gelatin-MA) polymerized within a PEG framework to form a mechanically robust network (Fig. 4A). The aim of this project was to increase the mechanical properties of bio/synthetic scaffolds and for use the encapsulating the cells. So, a bio/synthetic interpenetrating network showed ability in the incorporation of other ECM macromolecules and proteins, e.g. hyaluronan and HEP. The BioSINx adhered Cells displayed extensive cytoplasmic dispersion, and the GelMA: PEG ratio intensely affected the phenotype of the cells (Fig. 4B). Cells also adhered and proliferated over the physically-incorporated gelatin

had storage moduli with enhancing degrees of substitution (DS) of ADIBO. The chondrocyte spheroids in the Dex hydrogel provided significantly higher matrix content than individual chondrocytes due to the 3D structure of the spheroids. Moreover, Yu et al. fabricated an interpenetrating hydrogel to improve the mechanical properties and relatively poor degradation of the natural hydrogels [42]. In this content, a new hydrogel was obtained from the gelatin, HA and chondroitin sulfate (Chs) to mimic the natural cartilage ECM [11]. As known, HA has application in hydrogel systems owing to its good biodegradability, biocompatibility, as well as good gel-forming properties [43]. Furthermore, HA can be oxidized to develop chemical crosslinking action (Aldehyde groups) with amino functions by Schiff’s base linkage [44]. Besides, gelatin is a natural biopolymer that resist shear stresses in the 69

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after 3th day, ECs seemed noticeable changes in morphology and shape which could deliver a capability of EC growth and proliferation on the surface. Besides, after 7th day, ECs were well adhered to the surfaces except PCU-g-PEGMA surface, due to the absence of CAG peptide. Elastic polyurethane (PU) bearing pendant peptide was synthesized and its ability to retain and release of TGF-β1 was tested by Wu et al. [53]. In this work, polycaprolactone and polytetrahydrofuran tri-blocks enhanced elasticity of PU. The pendant peptide sequence had a high affinity for TGF-β1, even when the surface was protected by other proteins. He et al. synthesized polyvinylidene fluoride (PVDF) film tethered by Arginylglycylaspartic acid (RGD)-click-poly(glycidyl methacrylate) brushes (PGMA) [54]. To improve the cytocompatibility of PVDF films, PGMA were grafted onto the PVDF surface. Followed by introducing the azide groups on the PGMA side chains with sodium azide, the PVDF surfaces were grafted with PGMA using the secondary fluorine atoms. The cell adhesive RGD peptide was conjugated onto the PGMA brushes through alkyneazide click reaction. It should be noted that, by enhancing the pendant azide groups concentration, the amount of immobilized RGD peptide on the PGMA was increased. In other work, Liu et al. fabricated the hyperbranched PCL (hyPCL) formulation via the strain promoted alkyne-azide cycloaddition reaction between a BCNmodified 32-arm hyPCL dendrimer and an azide-modified 32-arm hyPCL dendrimer component [55]. The injectable hydrogels was synthesized after two polymers were mixed with together by syringe to produce the in situ scaffold. A tetra-hydroxy terminated 4-arm PEG was functionalized with acetylene and reacted with peptide azide/diazide and/or PEG diazide to produce hydrogels [56]. The gelation was very efficient and fast with very high swelling degree and good mechanical properties. Hydrogels containing peptide diazide delivered higher G’ in contrast to those containing peptide azide. Moreover, the proliferation of cells on the peptide-containing hydrogels was higher than the cells on PEG hydrogels due to the protein resistance properties of PEG. Moreover, by increasing the RGD peptide concentration, proliferation of primary human dermal fibroblasts was enhanced. In another example, polymers with clickable functional end-groups such as PEG-poly (trimethylene carbonate) (PTMC) could be used for the click reactions [57]. In this study, the cross-link density of the hydrogels could be optimized by varying the molar mass. The SEM images were shown that the gels are heterogeneous with interconnected pores. The degradation time was found to increase with increasing crosslink density or by decreasing MW between cross-links.

(BioSINP), but they formed spheroidal clusters which dashed into the matrix. By increasing gelatin content, the number of adhered cells on BioSINx was expressively higher than the number of adhered cells on BioSINP (Fig. 4Ba and Bb). The BioSINx provided a robust support for spreading and cell adherence; whereas, BioSINP caused cell proliferation into spheroidal structures (Fig. 4Bc and Bf). Another example for application of click chemistry was suing the water-soluble CS crosslinked through propionic acid ester-functional PEG [46]. This scaffold was constructed with crosslinking an azidefunctionalised CS by a water soluble 3-arm PEG containing an activated ester alkyne group. The G’ value at the gel formation time indicated a robust gel network structure. Moreover, the observed fast gelation time was due to the: (a) elevated temperature (37 °C) used for crosslinking; and (b) the CS-azide high functionality. The results of cell viability showed that MSCs on the surface of PEG-CS gels were grown for 1 week and seeded cells revealed high viability after 1 week in culture. Ziane et al. described the biological properties of a thermosensitive hydrogels based on low-MW glycosyl-nucleosyl-fluorinated (GNF) [47]. This compound was degraded slowly and could induce moderate chronic inflammation. The mechanical tests showed that even at short times (< 200 s), the elastic modulus was higher than the viscous modulus and the final elastic modulus was ∼5 times the viscous modulus. Moreover, when the culture medium containing cells incubated for 48 and 72 h with (GNF)-hydrogels, no significant cytotoxicity was detected for 48 h. The GNF hydrogel enabled ASC survival in vitro and in vivo and maintained them in the implantation site with uniformly distributed and round-shape. Grover et al. developed a novel methodology to synthesize hydrogels that encapsulate stem cells (3-D) or support adherence of the cells (2-D) by oxime bond formation [48]. 8-Armed aminooxy PEG was linked with glutaraldehyde (GTA) to form hydrogels. The mechanical properties of this system could be tuned by adjusting the weight percent of the aminooxy-PEG. By increasing the PEG chains length, the water absorption and the swelling ratios was increased. The encapsulated MSCs were viable for 1 week indicating the biocompatibility of this materials for cell encapsulation. Alg hydrogels are biologically inert materials that are used in the delivery of drugs, proteins, and cells [49]. Tetrazine and norbornene groups were grafted to Alg polymer chains for construction of hydrogels, allowing for covalent crosslinking. The swelling and mechanical properties of hydrogels could be tuned via the total polymer concentration and the stoichiometric ratio of the functional groups of polymer chains. In another work, thermosensitive hydrogels based on PEG-block-poly(γ-propargylL-glutamate) (PEG-PPLG) were prepared [50]. The in situ forming hydrogels were prepared by ring opening polymerization of γ-propargyl-Lglutamate N-carboxyanhydride (PLG NCA) in the presence of aminoterminated mPEG. Interestingly, to improve cell adhesion, the galactose group’s incorporation into the hydrogels was done owing to the fibronectin (FN) adsorption in cell-ECM. Moreover, the mass loss of the hydrogel was speeded in the presence of proteinase K, and the situ gels formed in the subcutaneous layer of rats duration of 3 weeks. Huynh et al. reported PEG-based hydrogel systems formed via thiol-epoxy [51]. The hydrogels were prepared by mixing solutions of 8-arm thiolated PEG (PEG(-SH)8) and PEG-diepoxy (PEG-DE) at basic pH. Cells could be encapsulated into the hydrogels and exhibited high cell viability over 4 weeks. In addition, a pro-osteogenic siNoggin was loaded into the hydrogels and the result showed that a pro-osteogenic siRNA, significantly enhanced the osteogenic differentiation of hMSCs. Khan et al. modified polycarbonate urethane (PCU) surface by both Cys-AlaGly-peptide (CAG) and PEG [52] (Fig. 5A). For this purpose, the hydrophilic macromonomer PEGMA was grafted onto PCU-film surface which showed an essential role in suppressing platelet adhesion. While the immobilized peptide had no direct inhibitory influence on platelet bonding, it combined by antifouling PEG to restrain platelet adhesion. The adhered endothelial cells (ECs) presented green color and PCU-gPEGMA and PCU-blank surfaces were used as controls (Fig. 5B). EC adhesion was uniform on all surfaces during the first day, while

2.3. Hydrogels constructed by Schiff base-reaction Schiff base is provided from amine and carbonyl groups to generate an imine group. It is biodegradable via hydrolysis and pH decreasing. So it is very suitable for use in the field of biomedical and tissue engineering [58–61]. Hydrogels resulting from natural polysaccharides are ideal scaffolds as they similar to the ECM of tissues. Tan et al. constructed an injectable hydrogels based on biodegradable CS-HA [62]. In the Schiff base reaction, the gelation was attributed to the reaction between AD and amino groups of polymers. So, in the current study, N-succinyl-CS (S-CS) and AD groups (A-HA) were used for preparation of the composite hydrogels. The S-CS/A-HA ratios did not significantly influence the gelation time and this time was about 1–4 min. Besides, the SEM images of hydrogels showed that by increasing the content of S-CS, the smaller pore diameters and tighter network structure were achieved and the compressive modulus of the composite hydrogels enhanced. With the less crosslinking in hydrogels, faster weight losing rate was observed and the hydrogels were dissolved in 1 day, while the other hydrogels lost their weight up to 1 month. In another work, melatonin conjugated CS microparticle loaded with methylprednisolone (MCC-MP) was prepared by Naghizadeh et al. [63]. The hydrogel was synthesized with the connection between the amine group of CMC and the AD group of Alg oxide. As a result, high release rates of melatonin (> 70%) and MP (nearly 100%) from microparticles 70

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Fig. 6. SEM image of chondrocytes attached to hydrogel pore walls (scale bar 100 µm) (A), calcein AM stained live cells encapsulated within hydrogel alone and with bioactive agents (B), and variation in DNA content (C). Reproduced from [1 0 0]. Copyright (2014) Elsevier Ltd.

Chondrocytes showed an increase around 2.53-fold in DNA content with time during 7 days (Fig. 6C). GAG deposition also increased 1.5- and 1.6-fold at 7 and 14 days, respectively. Yan et al. described the hydrogels fabricated by selfcrosslinking of hydrazide-modified poly(l-glutamic acid) (PLGA-ADH) and aldehyde-modified Alg (Alg-CHO) [65]. The mixing of PLGA-ADH and Alg-CHO solutions led to rapid formation of hydrogels at physiological condition. Hydrogels with high concentration displayed higher magnitude of G’ in contrast to low concentration hydrogels, proving the hydrogel was stiffer at higher solid content. In similar work, the PLGAADH and PLGA-CHO were prepared by using 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) activation and NaIO4 oxidation [66] (Fig. 7A). When the oxidation degree of PLGA-CHO increased, the solid content increased 3 ∼ 5 times (Fig. 7B and C). When the solid content or the oxidation degree of PLGA-CHO was high, it could be concluded that the crosslinking degree was high. The pore size of the hydrogel decreased rapidly when the solid content increased (Fig. 7D–G). After subcutaneously injection (Fig. 7Ha-Hd), the histological evaluation by toluidine blue and safranin-O staining staining showed positive blue staining, representing strongly deposition of GAG (Figs. 7Hc1–8Hd3).

were occurred after 24 h. While in the hydrogel/microparticle system, both drugs revealed low release rate (about 60% and 90% of melatonin and MP, respectively) during 7 days. Furthermore, the amount of released proteoglycan with the cells in hydrogel/microparticle was high, representing the properties of the hydrogel/microparticle in stimulating the cells for secretion of further GAGs. In another work, the authors illustrated that oxidized Alg (ADA) and gelatin undergo self-crosslinking in the presence of borax to form hydrogels [64]. These hydrogels revealed complete degradation within six weeks in vitro and 1 month in vivo. After filling of cartilage defect by the hydrogel, the SEM analysis confirmed fine integration of hydrogel with the cartilage. Additionally, frequency sweep analysis showed only a slight decrease in the G’ with no break, demonstrating the stability of the gel under shear force conditions. SEM image of a cross section of chondrocyte-incorporated hydrogel was shown after 1 week (Fig. 6A). Chondrocytes preserved their phenotype with expression of Col II and Acan. GAG deposition also exhibited a steady increase with time, which was significantly higher than that achieved for gel (Fig. 6B). After the incorporation of 3 bioactive molecules, including dexamethasone, PDGFBB and Chs, the cells are viable in the presence of these agents. 71

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Fig. 7. Images of the in situ formation of poly(l-glutamic acid) (PLGA) hydrogel using dual syringes (A), Frequency dependence of storage modulus/loss modulus (G′/ G″) for PLGA hydrogels with different solid contents (B) and oxidation degrees of aldehyde-modified PLGA (PLGA-CHO) (C). SEM images of the hydrogels with different solid contents: (D) 3, (E) 4, (F) 5, and (G) PLGA hydrogel as cell carriers and bare cell suspension for cartilage regeneration in vivo (H). Subcutaneous injection of hydrogels into athymic nude mice: (a) during injection, (b) after 12 weeks. (c, d) generated cartilageous tissues, (e) tissue mass, (c1, d1) safranin-O staining, (c2, d2) toluidine blue staining, and (c3, d3) immunohistochemical staining, respectively (bar scales: 200 µm). Reproduced from [66]. Copyright (2016) Royal Society of Chemistry.

example for Schiff-base reaction, Xu et al. used natural vanillin, as a crosslinking agent to construct CS hydrogels through the AD and the hydroxyl groups of vanillin, leading to form amide and hydrogen bond [69]. The point of note in this article was the presence of vaniline as a stimuli responsive crosslinker. When the vanillin was increased, an improvement was obtained in G’ due to the formation of a compact network. The amino-containing polyphosphazene and AD functionalized-Dex could react together [70]. Tyr (an antioxidant) and ferulic acid (hydrophilic monomer)-based side groups of polyphosphazene were coupled to AD functionalized-Dexs to form hydrogels. Hydrogel A with 25% oxidized Ald-Dex and hydrogel B with 98% oxidized Ald-Dex were compared and the storage moduli was much reduced for hydrogel A than hydrogel B. Cryo-SEM showed the low-density crosslink network with an average pore size of ∼114 µm for hydrogel A, while hydrogel B revealed a high-density crosslink network (∼26 µm). Glucose-responsive Dex-based hydrogels to deliver adipogenic factor was prepared by Tan et al. [71]. The gelation was corresponded to the cross-linking between aldehydic and aminated Dex (Fig. 8a). The adipogenic factor of insulin was integrated in the lectin concanavalin A (ConA) immobilized hydrogels to enhance adipogenesis. The competitive

Nguyen et al. developed a CS-HA based hydrogels that have potential to use as a bio glue for many medical applications [67]. In this work, by introducing carboxymethyl groups to the N-position and Oposition of the glucosamine and N-acetylglucosamine units of CS, the solubility of N,O-Carboxymethyl CS (NOCC) was improved. Additionally, HA was oxidized and the amino groups in NOCC were used to linkage to AD groups in A-HA. The hydrogels combined with hydroxy apatite (HAp) nanoparticles and calcium carbonate microspheres (CMs) were synthesized by Rena et al. [68]. The oxidized Alg (OAlg) consisting of AD and CMC consisting of amine (eNH2) groups were crosslinked. In the next step, tetracycline hydrochloride (TH) loaded CMs and HAp dispersion were blended by the hydrogel to improve mechanical strength of scaffolds. Due to the addition of HAp, the average pore sizes were decreased from 800 (pure gel) to 500 μm. When the HAp concentration increased, more than 4-fold improve was detected in stress analysis. Moreover, by adding the HAp, the gelation timewas reduced from 6 min to 3 min as the concentration of CMs enhanced which was associated to the hydrogen bonds of CMs and HAp. It should be noted that by increasing the content of HAp, the swelling ratio decreased by comparing with pure hydrogels. As an another 72

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Fig. 8. Synthetic route of aldehydic dextran (Dex) (a), aminated Dex (b), and Dex gel (c) (A), schematic illustration of glucose-responsive Dex hydrogel for selfregulating insulin delivery. Insulin would be released form gel by swelling of the hydrogel network (B), Gelation time of hydrogels (C), Weight loss of insulin/Dex and insulin/concanavalin A (ConA)/Dex hydrogels in PBS at 37 °C (D).

swelling was less in the CS hydrogels compared with HA and CS-HA hydrogels due to the smaller porosity perceived in CS hydrogel than HA and CS-HA hydrogels. Due to the interactions between cells and HA via CD44, CD44 can play as an important role in the migration of cells. Thus, the cells with high CD44 such as NIH3T3 and MG63 had higher migration when seeded onto the scaffolds compared with the control group. Additionally, matrix metalloproteinases (MMPs) show an important role in periodontal disease and the interaction between HA and CD44 can prevent the induction of MMPs in chondrocytes. In another research, controlling the elastic modulus of cellulose nanofibril (CNF) hydrogels-scaffolds was explored by Syverud et al. [75]. The carboxyl and AD groups were introduced on the CNF surfaces and the ethylenediamine (EDA) and hexamethylenediamine (HMDA) were used as crosslinker molecules. HMDA with four carbon atoms and longer chain was more flexible in the chain than EDA. Additionally, HMDA were able to connect to two AD groups, so gave stronger gels compared to EDA, constructing a functional crosslink between two fibrils. The elastic moduli of the gels were also adjusted by varying the concentration and length of the crosslinker molecules and by increasing the length of crosslinker molecules, the elastic moduli was decreased. Bhattacharjee et al. reported the presentation of CS on the in vitro deposition [76]. The presence of CS led to an increased construct stiffness and crosslinking could not induce alterations in the measured Col or GAG content. A 9fold increase in compressive modulus with cells was exhibited for silkbased scaffolds with time. Besides, the amount of GAG/DNA had increased up to 5.1-fold, but this amount were not significantly different in constructs based on silk. In another work, Dex was covalently grafted on HA-CS hydrogels by Sun et al. [77]. For this purpose, N-succinyl-CS was grafted with Dex by carbodiimide and the obtained results showed the better characteristics for hydrogels grafted with Dex compared to

displacement from the lectin receptor sites of Dex by glucose triggered by the ConA, resulting in decreasing in the cross-linking density (Fig. 8b). A shorter gelation time obtained by complexation between ConA and glucose moieties of Dex molecules, while the gelation time slightly increased after the addition of insulin in the hydrogels. The ConA/Dex and insulin/ConA/Dex hydrogels had a superior compressive modulus than the control Dex and insulin/Dex hydrogels (Fig. 8c). Thus, with translocation of ConA, the compressive modulus was improved. The insulin/ConA/Dex hydrogel showed weight loss up to 14 days and a slower weight losing rate than the insulin/Dex hydrogel was observed (Fig. 8d). Also, the in vitro insulin delivery demonstrated the delivery of insulin in response to physiologically relevant glucose concentrations. A tri-component hydrogel was reported by Khorshidi et al. [72]. The hydrogel was constructed by gelatin (G), oxidized Alg (ADA) and HA diAD (HDA). Swelling degree reduced along with enhance of total concentration of gel components. The hydrogel with equal concentration of polysaccharides and gelatin showed the highest degree of crosslinking and the lowest value of pore size. Cross-linked glycol CS with multibenzaldehyde functionalized PEG analogues reported as an injectable hydrogel by Cao at al [73]. Poly(ethylene oxide-co-glycidol)-CHO (poly (EO-co-Gly)-CHO) was synthesized and applied as a cross-linker. By enhancing the poly(EO-co-Gly)-CHO concentration about 10 fold, the gelation time and the water uptake of hydrogels could be reduced around 3.5 fold. While, with increasing the poly(EO-co-Gly)-CHO concentration (∼10 fold), G' gave rise to a 7-fold enhance. The hydrogels with higher cross-linker concentration showed a lower weight enhance, confirming the effect of poly(EO-co-Gly)-CHO concentration on the degradation rate of hydrogels. CS-HA hydrogel scaffold for periodontal tissue regeneration was introduced by Miranda et al. [74]. Water 73

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crosslinked samples in comparison with the other crosslinked samples [86]. Furthermore, the denaturation temperature of scaffolds enhanced with greater degree of crosslinking. Thus, the scaffolds crosslinked with GTA and GP showed the higher temperature denaturation in contrast to the 2 other crosslinking agents. The water absorption ability of scaffolds reduced owing to the reduction in the number of free amines acted as sites to form the hydrogen bonds. In this content, the swelling ratio of the scaffolds crosslinked with HMDI and epoxy was higher than that crosslinked with GTA and GP because of the elasticity of the structure decreased [83].

hydrogels alone. The grafting of Dex was an essential factor in hydrogel bioactivities, because of the AD-HA-N-succinyl-CS-Dex (AHA-SCS-Dex) hydrogel displayed higher water uptake, a slightly longer gelation time and faster weight loss than those of hydrogels without Dex. The degradation of AHA-SCS and AHA-SCS-Dex hydrogels enhanced by incubation time. This fact was due to the more water molecules attacked to the network chains because of higher swelling ratio in AHA-SCS-Dex hydrogel. Furthermore, after the cell seeding, the DNA content of the AHA-SCS-DEX hydrogel was improved after 5 days, while for the hydrogels without Dex, there was no important change in DNA number during 5 days. Tan et al. reported the synthesis of double-crosslinked HA hydrogel via AD and amino groups of polysaccharide derivatives, and the following crosslinking by genepin (GP) [78]. Opening the sugar ring was occurred to form di-AD derivatives and HA-NH2 was synthesized by coupling EDA to HA. It should be noted that addition of GP did not effect on the gelation time because of the reaction time of GP was slower than that of Schiff-base. The HA-HA/GP hydrogels confirmed a higher compressive modulus about 2.5 folds than the HA-HA hydrogel. The additional GP crosslinking led to increase in crosslinking density, decrease water content, dense microstructure, and slow mass loss. The thermoreversible laminin-functionalized hydrogel was fabricated by Stabenfeldt et al. [79]. For this purpose, by functionalization of methylcellulose (MC) with laminin-1 (LN), a bioactive scaffold was fabricated and the gelation mechanism for MC achieved from an intra- and inter-molecular hydrophobic interactions as the temperature enhanced. Due to the disruptions of hydrophobic interactions, the LCST of oxidized MC chain (OXMC) increased after the functionalization with LN (OXMC–LN (33.2 °C). Additionally, OXMC–LN supported higher cell survival in comparing to OXMC, MC, and MC/LN, demonstrating the important cellular response within LN-functionalized hydrogels. Fan et al. reported the CS-Chs hydrogels embedded with CS microspheres for tissue engineering and drug delivery [80]. CMC and oxidized Chs (OChs) were prepared and then, bovine serum albumin (BSA) loaded CS-based microspheres (CMs) were fabricated. Because of the amphiphilic nature of BSA protein acts as an emulsifying agent, the diameter of BSA-loaded CMs (3.8–61.6 μm) was slightly narrower than that of CMs (7.2–94.6 μm). It should be noted that adding CMs into hydrogels had reduce the gelation time because the excess amino groups in CMs may crosslink with the AD groups in OChs. Because of CS is not watersoluble occupied the porous of hydrogels, the CMs/gel showed minor swelling rate than the gel. Moreover, release study showed that of BSA was released less from the CMs/gel less than those of CMs and gel, proving that the incorporation of BSA-loaded CMs into hydrogels. In another work, an injectable hydrogel (Gel-CDH/HA-mAD) composed of carbohydrazide-modified gelatin (Gel-CDH) and HA mono-AD (HAmAD) was developed [81]. HA-mAD showed a greater stability of hydrogels than hydrogels containing conventional di-AD HA (HA-dAD). The MW of HA-dAD reduced compared with HA and HA-mAD, representing that HA-mAD was much more stable against hydrolysis than HA-dAD. Additionally, by varying AD modified HA from HA-dAD, to HA-mAD, the fabricated hydrogel became more stable against hydrolytic degradation and the lowest degradation rate around 3 weeks was observed. So, it was estimated that Gel-CDH/HA-mAD hydrogel could be applied as scaffold for a longer period. After embedding the dissected thoracic aorta of rat in Gel-CDH/HA-mAD hydrogels, cells migrated and microvascular effect did not happen in Low Gel-ADH/HA-dAD and Low Gel-CDH/HA-dAD hydrogels because they degraded within 1 week. Poursamar et al. used the 4 crosslinking agents that was appropriate for in-situ gas foaming method [82]. For this study, hexamethylene diisocyanate (HMDI), PEG diglycidyl ether (epoxy), GTA, and genipin were used as crosslinkers. It is worth mentioning that the chain length of crosslinking agents can influence on the elasticity [83], brittleness [84], and crosslinking density (index) [85] of structures. On the contrary with the longer crosslinker molecules, shorter crosslinkers could take the polymeric network closer together [84] and led to less flexibility in the hydrogels. Also, lower tensile strength was obtained via epoxy-

3. Conclusions and future perspectives The most widely used classifications of hydrogels are associated with the nature of chemical cross-linking. Because the prepared scaffold provides wide control over resultant hydrogel properties while physically crosslinked networks lose their stability in biological environments. In this review, we explained the most recent progress in the field of chemically gelling hydrogels as a tissue regeneration and highlighted examples of synthesized hydrogel scaffolds via the three varied crosslinking mechanisms including Click chemistry, Michael-type addition reaction and Schiff base reaction. Although several efforts have been organized to improve hydrogels for the development of functional engineered tissues, an important future success of hydrogel scaffolds is extremely dependent on the development of these gels with the physical, chemical, and biological properties in order to better mimic the structural. Therefore, there is also a continuing requirement to progress new strategies to control the incorporation and release of cellular biofactors. Declaration of Competing Interest No potential conflict of interests was reported by the authors. References [1] C. Hiemstra, L.J. van der Aa, Z. Zhong, P.J. Dijkstra, J. Feijen, Rapidly in situforming degradable hydrogels from dextram triols through Michael addition, Biomacromolecules 8 (2007) 1548–1556, https://doi.org/10.1021/bm061191m. [2] I. Jun, K.M. Park, D.Y. Lee, K.D. Park, H. Shin, Control of adhesion, focal adhesion assembly, and differentiation of myoblasts by enzymatically crosslinked cell-interactive hydrogels, Macromol. Res. 19 (2011) 911–920, https://doi.org/10.1007/ s13233-011-0909-6. [3] A. Fernández-Colino, F. Wolf, H. Keijdener, S. Rütten, T. Schmitz-Rode, S. Jockenhoevel, et al., Macroporous click-elastin-like hydrogels for tissue engineering applications, Mater. Sci. Eng. C 88 (2018) 140–147, https://doi.org/10. 1016/j.msec.2018.03.013. [4] J. Joy, J. Pereira, R. Aid-Launais, G. Pavon-Djavid, A.R. Ray, D. Letourneur, et al., Gelatin- Oxidized carboxymethyl cellulose blend based tubular electrospun scaffold for vascular tissue engineering, Int. J. Biol. Macromol. 107 (2018) 1922–1935, https://doi.org/10.1016/j.ijbiomac.2017.10.071. [5] H. Ju, F. Zhu, H. Xing, Z.L. Wu, F. Huang, Ultrastiff hydrogels prepared by Schiff’s base reaction of bis(p-formylphenyl) sebacate and pillar[5]arene appended with multiple hydrazides, Macromol. Rapid Commun. 38 (2017), https://doi.org/10. 1002/marc.201700232. [6] B.G. Molina, L. Cianga, A.D. Bendrea, I. Cianga, L.J. Del Valle, F. Estrany, et al., Amphiphilic polypyrrole-poly(Schiff base) copolymers with poly(ethylene glycol) side chains: Synthesis, properties and applications, Polym. Chem. 9 (2018) 4218–4232, https://doi.org/10.1039/c8py00762d. [7] A.I. Cañas, J.P. Delgado, C. Gartner, Biocompatible scaffolds composed of chemically crosslinked chitosan and gelatin for tissue engineering, J. Appl. Polym. Sci. 133 (2016) 1–10, https://doi.org/10.1002/app.43814. [8] Y. Liang, W. Liu, B. Han, C. Yang, Q. Ma, F. Song, et al., An in situ formed biodegradable hydrogel for reconstruction of the corneal endothelium, Colloids Surf. B Biointerfaces 82 (2011) 1–7, https://doi.org/10.1016/j.colsurfb.2010.07.043. [9] Y. Liu, M.B. Chan-Park, Hydrogel based on interpenetrating polymer networks of dextran and gelatin for vascular tissue engineering, Biomaterials 30 (2009) 196–207, https://doi.org/10.1016/j.biomaterials.2008.09.041. [10] H. Luo, G. Xiong, D. Hu, K. Ren, F. Yao, Y. Zhu, et al., Characterization of TEMPOoxidized bacterial cellulose scaffolds for tissue engineering applications, Mater. Chem. Phys. 143 (2013) 373–379, https://doi.org/10.1016/j.matchemphys.2013. 09.012. [11] X. Hu, D. Li, F. Zhou, C. Gao, Biological hydrogel synthesized from hyaluronic acid, gelatin and chondroitin sulfate by click chemistry, Acta Biomater. 7 (2011) 1618–1626, https://doi.org/10.1016/j.actbio.2010.12.005.

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