Modularly engineered alginate bioconjugate hydrogel as biocompatible injectable scaffold for in situ biomineralization

Journal Pre-proof Modularly engineered alginate bioconjugate hydrogel as biocompatible injectable scaffold for in situ biomineralization Seong Han Kim, Thavasyappan Thambi, V.H. Giang Phan, Doo Sung Lee

PII:

S0144-8617(20)30006-0

DOI:

https://doi.org/10.1016/j.carbpol.2020.115832

Reference:

CARP 115832

To appear in:

Carbohydrate Polymers

Received Date:

9 August 2019

Revised Date:

2 January 2020

Accepted Date:

3 January 2020

Please cite this article as: Kim SH, Thambi T, Giang Phan VH, Lee DS, Modularly engineered alginate bioconjugate hydrogel as biocompatible injectable scaffold for in situ biomineralization, Carbohydrate Polymers (2020), doi: https://doi.org/10.1016/j.carbpol.2020.115832

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(Revised Manuscript for the Carbohydrate Polymers)

Modularly engineered alginate bioconjugate hydrogel as biocompatible injectable scaffold for in situ biomineralization

Seong Han Kim,a,1 Thavasyappan Thambi,a,1 V.H. Giang Phan b and

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Doo Sung Lee a,*

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School of Chemical Engineering, Theranostic Macromolecules Research Center, Sungkyunkwan University, Suwon 16419, Republic of Korea b Biomaterials and Nanotechnology Research Group, Faculty of Applied Sciences, Ton Duc Thang University, Ho Chi Minh City 70000, Vietnam

These authors contributed equally to this work.

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*Corresponding author: Doo Sung Lee, Ph.D.

Tel.: +82-31-299-6851; Fax: +82-31-299-6857; e-mail: [email protected]

Highlights 1

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Injectable bioconjugates were synthesized by the post modification of alginate.

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Bioconjugates exhibited sol-gel transition between room and body temperature.

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Bioconjugate gels in SBF showed the nucleation and growth of hydroxyapatite.

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BMP 2-loaded bioconjugate hydrogel exhibited in situ biomineralization in vivo.

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Abstract In the present study, a type of bioconjugate was synthesized by post modification of alginate by

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conjugating temperature-responsive poly(ε-caprolactone-co-lactide)-b-poly(ethylene glycol)-b-

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poly(ε-caprolactone-co-lactide) and O-phosphorylethanolamine as phosphorylation functional groups. Freely flowing bioconjugate sols at low temperature can transform to stable viscoelastic

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gels at the physiological temperature (37 °C). Subcutaneous administration of temperatureresponsive bioconjugate sols into the dorsal region of Sprague-Dawley rats formed in situ hydrogel.

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In situ formation of bioconjugate gels in stimulated body fluids at 37 °C showed nucleation and hydroxyapatite mineral growth. Furthermore, hydroxyapatite growth was also found in in vivo gels,

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which suggested the potential of alginate-based bioconjugate gels as a scaffold for bone engineering. Bone morphogenetic protein 2 (BMP-2)-loaded bioconjugate formed stable gel in vivo, and demonstrated sustained release. BMP-2-loaded bioconjugates exhibited in situ biomineralization in vivo. These results imply that the in situ formation of injectable biomimetic

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materials has potential for bone tissue engineering applications. Key words: controlled release; alginate; temperature-responsive; BMP-2; biomineralization; bone repair

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1. Introduction

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Over the past few decades, smart polymeric materials that can transform from free-flowing liquid to non-flowing stable gels at the body condition have paid great attention in biomedical

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applications (L. Yu & Ding, 2008). The water-based phase transition induced by the physiological

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stimuli, such as pH, enzyme, and temperature, enables the minimal invasive therapy, which allows the entrapment of therapeutic agents, proteins, and cells, and forms a controlled delivery depot at

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the site of interest (Du, Zhou, Shi, & Xu, 2015; K. Y. Lee & Mooney, 2001; Vermonden, Censi, & Hennink, 2012). In particular, the in situ transforming hydrogels prepared using the physical cross-

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linking of amphiphilic copolymers were found to be a perfect means to adapt to irregularly shaped wounds or surfaces, and offers benefit for practical applications that include bone regeneration (Balakrishnan & Banerjee, 2011). Generally, during intrinsic bone repair, the extracellular matrix proteins (e.g., collagen, fibronectin, and proteoglycans) accelerate the conversion of amorphous

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calcium phosphate into crystalline calcium phosphate, mainly apatite (He, Dahl, Veis, & George, 2003; J. Li et al., 2013; D. Wu et al., 2013). This is mainly due to the presence of acidic functional groups, which could acts as binding domains for calcium ions, and their subsequent alignment led to apatite crystal lattice (He et al., 2003). Therefore, it is of the utmost importance to develop polymeric hydrogel precursors with in situ sol-to-gel phase transition. Such hydrogels could adapt to irregular defect surfaces by minimal invasive administration, and introduction of phosphate 3

functional group in the biopolymer mimic extracellular matrix proteins that could accelerate in situ mineralization for bone repair. Numerous hydrogel systems, derived from synthetic and natural polymeric materials, have been used for biomedical application (Calvert, 2009; Gu & Mooney, 2015; Han et al., 2017; Han et al., 2018; Nguyen & Alsberg, 2014; S. Y. Wang et al., 2018; S. Yu et al., 2018). In particular,

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thermo-responsive polymers showing lower critical solution temperatures (LCSTs) with ideal phase transitions between room and body temperatures are attractive polymers (Bordat, Boissenot,

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Nicolas, & Tsapis, 2019). Such thermo-responsive polymers are free-flowing sols at low

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temperature and transform to stable viscoelastic gel upon raising the temperature to the physiological condition. For example, pluronic, poly(phosphazene), and poly(N-isopropyl

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acrylamide) exist sol to gel phase transition between room temperature and body temperature (K. M. Park et al., 2009; M.-R. Park, Seo, & Song, 2013; Sanoj Rejinold, Sreerekha, Chennazhi, Nair,

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& Jayakumar, 2011). However, the biomedical applications of these polymers have been limited due to nondegradability and limited biocompatibility. To overcome these limitations, poly(ester)-

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based copolymers, generally regarded as safe (GRAS) materials, are found to be biocompatible and biodegradable when they implanted into the body. Among them, poly(lactide), poly(εcaprolactone), poly(ethylene glycol) (PEG), poly(ε-caprolactone-co-lactide)-b-PEG-b-poly(ε-

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caprolactone-co-lactide) (PCLA), and poly(lactic-co-glycolic acid) (PLGA) have advantages in tuning molecular weight, gelation window, in vivo gelation, and biodegradation (Bae, Wang, & Kurisawa, 2013; T. T. Lee et al., 2014; Phan, Thambi, Gil, & Lee, 2017; Thambi, Li, & Lee, 2017). Furthermore, PEG and poly(ester)-based have been extensively used as localized drug carriers, which have been approved by the FDA that proved to be safe for in vivo implantation (Stewart, Domínguez-Robles, Donnelly, & Larrañeta, 2018). The poor efficacy of the proteins that are 4

delivered via oral or nasal routes usually necessitates frequent injections, which is inconvenient for patients who require prolonged therapeutic regimens (Ghadiri, Young, & Traini, 2019). Therefore, proteins and peptides have been loaded into hydrogel depot for controlled delivery (Buwalda, Vermonden, & Hennink, 2017; Gil et al., 2017; Kim, Park, & Song, 2012; Phan et al., 2017; Turabee, Thambi, Lym, & Lee, 2017). When designing the hydrogel systems for prolonged

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release, a number of issues should be considered beforehand. For instance, the drug needs to be released for a prolonged time without an initial burst within the required time (Ci, Chen, Yu, &

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Ding, 2014). In addition, the interaction between proteins and hydrogel materials is an important

which can lead to burst release (Thambi et al., 2017).

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issue, because many proteins are charged, and they are not able to bind in uncharged formulation,

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The rate of biodegradation of bone regenerating biomaterials should match the optimal

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bone regeneration; and the delivery of growth factors and minerals should persist for a long time, to ensure the bone regeneration at the site of implantation (Y. Li, Rodrigues, & Tomás, 2012; M. Liu et al., 2017; Y. Liu et al., 2017; Z. Wang et al., 2017; Wei, Yuan, Cai, Mao, & Yang, 2018).

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The general class of hydrogels used in bone regeneration is based on physically cross-linked nondegradable gels (Tozzi, De Mori, Oliveira, & Roldo, 2016). However, the lack of acidic functional groups in these gels limits their biomineralization in situ. Therefore, the rational design of

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biomaterials based on polysaccharide-polymer conjugates may stimulate the biomimetic strategy. Fusion of synthetic and natural polymers can utilize the advantages of both components (Giang Phan et al., 2019; Thambi, Phan, & Lee, 2016). Therefore, appropriate fusion of synthetic and natural polymers may provide useful injectable hydrogel platform (Le et al., 2018; Y. Li et al., 2012). In particular, it is known that during natural bone formation, the extracellular matrix proteins in collagens can accelerate conversion of amorphous calcium phosphates to aligned 5

apatite crystals (Schröder et al., 2015) (Lin, Cao, Chen, Wu, & Li, 2013; W. Wu et al., 2017). To realize our objective, in this study, we synthesized a set of bioconjugates based on alginate finely grafted with PCLA and O-phosphorylethanolamine as an in situ implantable device for in situ biomineralization and subsequent bone regeneration (Scheme 1). Alginate is a non-toxic, biocompatible, biodegradable, and non-immunogenic anionic polymer, composed of anionic

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polymer residues of guluronic acid (G) and mannuronic acid (M) (Bidarra, Barrias, & Granja, 2014; Diaz-Rodriguez, Garcia-Triñanes, Echezarreta López, Santoveña, & Landin, 2018; Shih et al.,

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2018; Zhang et al., 2018). The bioconjugates were administered to Sprague-Dawley (SD) rats, and

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the biodegradation pattern was assessed. In addition, the bone morphogenetic protein 2-loaded bioconjugates formed stable gel in vivo, and the release pattern was examined. Finally, we injected

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bioconjugates subcutaneously into dorsal region of SD rats, and the hydroxyapatite mineral formation

2. Materials and method

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2.1. Materials

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was examined by Masson’s trichrome (MT) and alizarin red S staining.

α,ω-Bis-hydroxy-poly(ethylene glycol) (PEG, Mn = 1650 g/mol) was purchased from Pharmicell Co. Ltd. (Gyeonggi-do, Korea). D,L-lactide (LA), alginate sodium salt (Mn =120,000-190,000

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g/mol; product number: 180947), 3-(boc-amino)-1-propanol, tin-2-ethylhexanoate (Sn(Oct)2), hydrogen chloride solution (4.0 M in dioxane), ninhydrin, N-hydroxysuccinimide (NHS), and Ophosphorylethanolamine were supplied by Sigma-Aldrich Co. (St. Louis, MO, USA). εCaprolactone (CL) and 1-(3-Dimethylaminopropyl)-3-ethylcarbodiimide (EDC) were purchased from the Tokyo Chemical Industry Co., Ltd. Bone morphogenetic protein-2 (BMP-2, human, recombinant) was supplied by Sino Biological, Inc. (Beijing, China). The PCLA copolymer was 6

prepared using previously reported procedure (Giang Phan et al., 2019). 2.2. Preparation of Alg-PCLA bioconjugates Scheme 2 shows that the Alg-PCLA bioconjugates are prepared by the individual synthesis of amine-functionalized PCLA copolymer (PCLA-NH2) and phosphate-modified alginate (Alg-PEA),

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followed by the conjugation of both polymers to alginate through amide bond formation. 2.2.1. Synthesis of PCLA-NH2

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PCLA-NH2 was synthesized via the bulk polymerization of CL and LA with 3-(Boc-amino)-1propanol initiator, and Sn(Oct)2 as catalyst. Briefly, 3-(Boc-amino)-1-propanol (0.625 mL),

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Sn(Oct)2 (0.06 g), CL (4 mL), and LA (1.217 g) in a flask were dried at 55 °C under vacuum for 1

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h. Next, nitrogen was passed and the temperature was increased to 110 °C, and stirring continued for 24 h. Then, the Boc-protected PCLA copolymer (PCLA-NH-Boc) residue was dissolved in

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chloroform, and precipitated in a mixture of n-hexane and diethyl ether (50/50 (v/v)). The obtained product was filtered and vacuum dried for 72 h at room temperature (RT).

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The PCLA-NH-Boc (2.5 g) was dissolved in 4 M HCl in dioxane (20 mL). The flask was filled with nitrogen, and the reaction was continued at RT for overnight. The reaction mixture was concentrated using vacuum evaporation to obtain the product as yellow solid, and the product was used in the next step without further purification. Boc deprotection reaction was monitored by

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ninhydrin assay.

2.2.2. Synthesis of Alg-PCLA bioconjugates For the preparation of Alg-PCLA bioconjugates, the alginate was first functionalized with phosphate groups, followed by conjugation with PCLA-NH2. Phosphate-functionalized alginate 7

(Alg-PEA) was synthesized via EDC/NHS reaction between the amine group on Ophosphorylethanolamine, and the carboxylic acid group on the backbone of alginate ([Amine]: [Carboxylic acid] = 1: 10). Briefly, alginate sodium salt (3 g) was dissolved in DI water (150 mL). Thereafter, EDC (0.87 g) and NHS (1.305 g) were added and pH was tuned to 6.0, and stirred for 30 min at RT. O-phosphorylethanolamine (0.32 g) was then added, and continued the reaction at

non-conjugated reagents, and lyophilized to obtain Alg-PEA conjugates.

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RT for 24 h. Finally, the product was dialyzed (MWCO: 12-14 kDa) against DI water to remove

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The Alg-PCLA graft copolymer was synthesized via the EDC/NHS reaction between the

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amine group on PCLA-NH2 and the carboxylic acid group on the backbone of Alg-PEA ([Amine]: [Carboxylic acid] = 1: 10). Alg-PEA (0.5 g) was dissolved in DI water (25 mL). Thereafter, EDC

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(0.145 g) and NHS (0.217g) were added and the pH was adjusted to 6.0. Subsequently, PCLANH2 (0.963 g) dissolved in 10 mL of DMSO was dropwise added to the solution, and stirred for

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24 h. Finally, the product dialyzed (MWCO: 12-14 kDa) against DI water to remove the DMSO

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and unreacted reagents and lyophilized to obtain Alg-PCLA bioconjugates.

2.3. Characterization

2.3.1. 1H NMR. The chemical structure of copolymers and modified reagents was measured using H NMR measurements (Varian Unity Inova, 500 MHz). The sample was prepared in CDCl3 and

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D2O NMR solvents. 2.3.2. Gel permeation chromatography (GPC). The Mn of copolymers was measured using gel permeation chromatography (GPC). Agilent 1100 system was used, with CHCl3 and H2O as an eluent. PEG standards were used for molecular weight calibration. 2.3.3. Zeta potential (ζ). The ζ of copolymers and alginate conjugates was measured using dynamic 8

light scattering (DLS) (Zetasizer, Nano-ZS90). Briefly, polymer solution with 10 mg/mL was prepared, stabilized for 30 min at RT and measured. 2.3.4. Rheological study. Rheological properties of Alg-PCLA and PCLA hydrogels were examined using dynamic mechanical analyzer (Bohlin Rotational Rheometer) in the oscillation mode. For the rheological measurement, Alg-PCLA or PCLA hydrogels (22.5 wt%) were placed

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between the parallel plates with a gap of 250 µm, and the measurements were conducted under

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2.4. In vitro gelation and sol-gel phase transition

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controlled shear stress (0.4 Pa) with a frequency of 1 rad s-1.

The in vitro gelation and sol-gel transition pattern of PCLA copolymer and Alg-PCLA conjugate

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solutions were investigated by tube-inverting method (Phan et al., 2017). In short, either

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copolymers or bioconjugates were suspended in PBS with various given concentrations. The prepared sample kept at 2 °C for 12 h. Thereafter, all the vials with copolymer or bioconjugate

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solutions were kept in temperature-controllable water bath. The vials were heated from 10 °C to 65 °C with 2 °C of temperature interval, and sample vials were inverted to observe flowability and determine the sol-gel transition.

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2.5. Cytotoxicity in vitro

The cytotoxicity of bioconjugate and copolymer was investigated via MTT assay by using 293T cells (KCLB Inc., Korea). Briefly, cells were cultured using DMEM consist 10% (v/v) FBS and 1% (w/v) penicillin-streptomycin under humidified atmosphere at 37 °C. After 24 h of growth, cells were directly exposed to various concentration of hydrogel solution. One day after incubation, each well received 20 μL of MTT solution, and incubated further 3 h. Then, media were removed, 9

and DMSO (200 μL) was added. After 1 h incubation, microplate reader was used to measure absorbance at 490 nm. Cells treated with only RPMI 1640 medium were used as a control (0 µg/mL) and the viabilities were considered 100%.

2.6. In vitro mineralization

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To investigate the mineralization property in vitro, Alg-PCLA bioconjugate hydrogel (22.5 wt.%) in a dialysis bag (MWCO = 1 kDa) was immersed stimulated body fluid (SBF, 50 mL) at 37 °C.

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Every 24 h, the SBF was replenished with fresh solution. After 4 weeks, the dialysis bag with

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mineralized hydrogel was washed by soaking in distilled water for 12 h. Thereafter, mineralized hydrogel was obtained by being lyophilized for 3 days. X-ray diffraction (XRD, D8 ADVANCE,

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Bruker) analysis was used to examine the formed crystal inside the mineralized hydrogel.

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2.7. In vivo gel formation and biodegradability

All animal experiments were performed in compliance with the relevant laws and institutional

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guidelines of Sungkyunkwan University, and were approved by institutional committees. To evaluate the in vivo gel formation and biodegradability of PCLA and Alg-PCLA hydrogel, each hydrogel solution (22.5 wt.%) was prepared. Each 300 µL of samples was subcutaneously injected

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into male SD rats (Hanlim Experiment Animal Research Institute, Korea). After predetermined time point, the rats humanly sacrificed, and gels formed were harvested, lyophilized for 3 days, and weighed to investigate the rate of biodegradation.

2.8. Release of BMP-2 protein in vivo To study in vivo release of BMP-2 protein, rats were split into 3 groups: (i) control (free BMP-2 10

solution), (ii) BMP-2-loaded PCLA hydrogel, and (iii) BMP-2-loaded Alg-PCLA hydrogel (n = 3 for each group). Briefly, hydrogel solution (22.5 wt.%) that contained 10 µg/mL of BMP-2 protein was prepared, and 300 µL of samples in each group were injected subcutaneously into rats. At predetermined time point, blood samples were obtained from tail vein. Collected blood samples were centrifuged to obtain serum, and the concentration of BMP-2 protein in serum was analyzed

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by using BMP-2 ELISA Kit (Sino Biological, Inc., China). The BMP-2 loading in the hydrogels was optimized according to the previously reported

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procedure (Kissling et al., 2016). To optimize BMP-2 loading into the hydrogels, different

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concentration of BMP-2 (5, 10 and 20 µg/mL) containing hydrogel solution (22.5 wt.%) were prepared. Among them, we chose 10 µg/mL of BMP-2 protein containing formulation was we used

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for our study because its shows good dispersion and optimal amount form biomineralization.

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2.9. In vivo gel morphology and mineralization

SEM-EDS (JSM-7500F, JEOL, Tokyo, Japan) was used to investigate the morphology and

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mineralization property of in vivo gel. The BMP-2-loaded PCLA and Alg-PCLA hydrogel solutions (300 µL, 22.5 wt.%) were injected, collected at predetermined time points, and lyophilized for 3 days, as previously mentioned. The lyophilized hydrogels were cut cross-

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sectioned to observe the microporous structure. Furthermore, energy dispersive spectroscopy (EDS) analysis was used to investigate the Ca/P ratio of mineralized hydrogel.

2.10. Histological analysis For the histological analysis, BMP-2-loaded hydrogel precursors (300 µL, 22.5 wt.%) were injected subcutaneously into rats. The hydrogels with surrounding tissues were collected at 11

predetermined time point, fixed with 10% formalin solution, embedded using paraffin, and sectioned (4 µm each). Paraffin embedded sections were MT and Alizarin red S stained, according to the previous report. Stained sections were visualized using optical microscope. The quantification of calcium deposition and collagen fibers was done using ImageJ software

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(National Institute of Health, Bethesda, MD, USA).

2.11. Statistical analysis

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Statistical significance between experimental group and control group was conducted using the

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one-way ANOVA test. The values are presented as mean ± standard deviation. The values are

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considered statistically significant if p < 0.05.

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3. Results and discussion

3.1. Preparation and characterization of phosphorylated Alg-PCLA bioconjugate The phosphorylated Alg-PCLA bioconjugates were prepared by the sequential conjugation of

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PCLA-NH2 and O-phosphorylethanolamine to alginate via amide bond formation via EDC/NHS chemistry, as shown in Scheme 2. Firstly, PCLA-NH2 was prepared using the polymerization of LA and CL with N-Boc protected butanol initiator. The 1H NMR of PCLA-NH-Boc shows the

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characteristics peaks of ethylene glycol units (-CH2CH2O-) at 3.65 ppm, methine protons of lactide units (-CH-) at 5.12 ppm, and caprolactone units at 2.31 ppm, indicating effective polymerization (Fig. 1). The molar composition of CL and LA in the copolymer was determined by comparing the integration peaks with PEG (Table S1). Acidic treatment of PCLA-NH-Boc results the formation of PCLA-NH2, which was confirmed by the ninhydrin assay.

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For effective phosphorylation, phosphates groups were introduced into the backbone of alginates. For the introduction of phosphates groups, O-phosphorylethanolamine was conjugated to alginate. The 1H NMR spectra show that new characteristics peaks at 3.84 ppm and 3.46 ppm correspond to the methylene protons of O-phosphorylethanolamine, which implied the effective conjugation of phosphorylated functional groups (Fig. 1). Finally, the PCLA-NH2 was conjugated

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to Alg-PEG to obtain Alg-PCLA bioconjugates. 1H NMR and water GPC were used to examine the successful conjugation of PCLA to alginate. The GPC trace shows a remarkable peak shift to

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lower retention time, which suggested the conjugation of PCLA-NH2 units, and increased the

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molecular weight of the bioconjugates (Fig. 2A). The GPC trace found that 6.6 units of PCLA were conjugated to alginate polymer (Table S2). Furthermore, ζ analysis demonstrated that

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nucleophilic substitution of alginate carboxylic acid groups by O-phosphorylethanolamine and

3.2. Sol-gel phase transition

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PCLA-NH2 increased the zeta potentials (Fig. 2B).

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The sol-gel transition patterns of PCLA copolymers and hybrid copolymers (mixture of PCLA copolymer and Alg-PCLA bioconjugate with a weight ratio of 1:8, hereafter the mixtures were called as Alg-PCLA bioconjugate hydrogel) in an aqueous solution was investigated using the vial inversion technique. Fig. 2C shows that the sol-gel transition of bioconjugates and copolymers

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were influenced by temperature and concentration. At low concentration and temperature, both PCLA copolymers and Alg-PCLA bioconjugates were in a sol state, but at the physiological temperature transformed to a gel state. The critical gelation concentrations (CGC) of PCLA copolymers and Alg-PCLA bioconjugates were found to be 17 wt.% and 16.8 wt.%, respectively. In aqueous solutions, copolymers assembled into flower-like micelles with PCLA blocks and PEG blocks as cores and shells, respectively. Concentrated clusters of inter-connected flower-like 13

micelles induce gel formation at the body condition. Interestingly, Alg-PCLA bioconjugates form grafted flowers on alginate surface. Aqueous copolymer solutions, at low temperature, exist as sols containing free-flowing clusters of micelles. This is mainly due the hydrogen bonding between hydrophilic copolymers and water. However, at the physiological condition, the hydrophobicity increases due to the weakening of hydrogen bonds, resulting micelle aggregation and form rigid

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bridges between hydrophobic PCLA segments led to the gel formation. When the temperature increased further, hydrophobic blocks shrunk, and the PEG blocks undergo dehydration, which led

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to phase separation.

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The mechanical properties of the hydrogels influence to certain degree the capacity of the hydrogelators to mix with therapeutic proteins or peptides (Nguyen & Alsberg, 2014). In addition,

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these properties will also affect the injectability of hydrogel precursor when using a hypodermic needle. Fig. 2D shows that at 22.5 wt% concentrations (23 °C), PCLA copolymer and Alg-PCLA

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bioconjugate showed low viscosity. This indicated a huge benefit in terms minimal invasive administration. Remarkably, viscosity changes abruptly to three orders of magnitude for PCLA

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and five orders of magnitude for the Alg-PCLA bioconjugate at the physiological temperature (37 °C). These results indicated the effective in situ sol-to-gel phase transition.

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3.3. Biocompatibility test

To examine the gelation of bioconjugates on warm-blooded animals, in vitro compatibility of materials was tested first on human embryonic kidney 293T cells. To examine the effect of cytotoxicity on 293T cells, cells incubated with solutions of PCLA copolymers and Alg-PCLA bioconjugates for 48 h (Fig. 3A). The 293T cells incubated with cell culture media alone were considered as a control group. Interestingly, cell proliferation was not affected up to a 14

concentration of 1,000 µg/mL of bioconjugates. The statistical significance analysis shows the no significance difference between control group (0 µg/mL) and1000 µg/mL, indicating the biocompatibility of PCLA and Alg-PCLA bioconjugates. The high cell viability of the bioconjugates group may be due to natural polysaccharides in the bioconjugates and the biocompatible PCLA copolymers. It should be noted that free alginate and phosphorylated

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alginates found to be non-toxic to cells (Fig. S1A). Furthermore, PCLA and Alg-PCLA hydrogel releasates not exhibited toxicity, suggesting that degradation fragments of hydrogels were also

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compatible with cells (Fig. S1B). Therefore, the surface functionalization of synthetic copolymers

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using naturally occurring biopolymers could improve the biocompatibility profile of a hydrogel network; also, the functionalization of synthetic polymers could be an advantage in biomedical

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applications.

3.4. In situ gelation and biodegradation in vivo

The in situ gelation of bioconjugates and copolymers was investigated by dorsal subcutaneous

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injection of copolymers (200 µL, 22.5 wt.%) into SD rats using hypodermic needles (26G). Fig. 3B shows that the thermo-responsive properties of the bioconjugates or copolymers resulted in stable gel formation underneath the skin. The in situ gels were mainly formed by physical cross-

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linking of molecular self-assembly, which showed biochemical and biophysical stability, following administration to SD rats. Thus, the bioconjugate hydrogel is suitable depot to control the delivery of therapeutic agents.

The biodegradation behavior and biocompatibility of biomaterials are important factors in therapeutics delivery. The biodegradation pattern of hydrogels in vivo was investigated by cutaneous injection of 22.5 wt.% copolymers, and examined as a function of time. Fig. 4A and 4B 15

shows that the PCLA hydrogel slowly degraded. The presence of cells and enzymes in the subcutaneous layers initially eroded the surface of the biogel networks. Thereafter, the biodegradation rate was accelerated by the hydrolysis of polyesters in the PCLA copolymers. In contrast, the complete degradation of Alg-PCLA bioconjugate hydrogels lasted eight weeks. During the first week, the Alg-PCLA hydrogels degraded slowly. After four weeks, approximately

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50% of hydrogels remained. During the experimental period, no hemorrhaging or tissue necrosis was observed. These results imply that the bioconjugate hydrogels are not-toxic and biodegradable

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in vivo, and suggest that bioconjugates fit for in vivo implantation.

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To observe the biodegradability of hydrogels, they were incubated under different pH conditions (e.g., pH 4.0, pH 7.4, and pH 10.0) (Fig. S2). In addition, alginate lyase, which is known

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to hydrolyze glycosyl linkages, has been chosen as an enzyme to test the enzymatic hydrolysis. Addition of alginate lyase enzyme to the hydrogels exhibited sustained degradation and was found

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that around 20 to 30% of hydrogels were degraded after 10 days. On the other hand, hydrogels incubated under acidic and basic conditions degraded rapidly. Interestingly, the hydrogels

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incubated under the physiological condition found to be stable and exhibited negligible degradation.

3.5. In vitro biomineralization

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In situ growth of hydroxyapatite crystals from SBF is a process that involves nucleation and growth of apatite crystal. Due to the presence of calcium phosphate, the SBF is saturated with Ca2+ and HPO42-; the ion concentration is significantly low in critical nucleation concentration, and not induced crystal nucleation (Jiang et al., 2017). It is expected that the interaction between rich phosphates on bioconjugates and Ca2+ ion in the SBF can induce heterogeneous nucleation, and accelerate growth of crystal on the bioconjugates (Qi, Musetti, Fu, Zhu, & Huang, 2019). 16

To examine the capability of bioconjugates to induce in situ biomineralization, Alg-PCLA bioconjugate and PCLA copolymer hydrogel (22.5 wt.%) were exposed with SBF for 4 weeks. XRD analysis was used to examine the newly formed layer (Fig. 5). The diffraction peaks at 2θ = (25.96, 31.87, 32.97, 39.96, 46.86 and 49.74)° correspond to the lattice planes of (002), (211), (300), (130), (222), and (213) of hydroxyl apatite. Other diffraction peaks at 2θ ranging from (18

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to 35)° are attributed to amorphous structure of Alg-PCLA bioconjugates. Four weeks after biomineralization, new diffraction fraction peaks at 2θ = 25.97° and a broad peak at 32°

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corresponds to lattice planes of (002) and (211), respectively. The XRD results demonstrate that

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the Alg-PCLA bioconjugate hydrogel induced the growth of crystalline hydroxyl apatite.

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3.6. In vivo BMP-2 release

To investigate the ability of Alg-PCLA bioconjugate hydrogel and PCLA hydrogel to

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sustain the delivery of mitogenic factors, BMP-2 was loaded into the hydrogel precursors by simple mixing (Fig. 6A). The in vivo release of BMP-2 was followed by subcutaneous

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administration of BMP-2-loaded hydrogel precursors in male SD rats (Fig. 6B). Furthermore, subcutaneously administered free BMP-2 solution group was considered as control group. The hydrogel formulations were able to sustain the release of BMP-2 for at least 3 days. The PCLA

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hydrogel formulation showed an initial burst for the first 16 h, which after 24 h eventually plateaued. Interestingly, the initial burst release was remarkably inhibited, and the release pattern levelled off for the remaining period tested. The strong complexation between Alg-PCLA bioconjugate hydrogel and BMP-2, and the sustained biodegradation of hydrogel networks, effectively inhibit initial burst, and allow the release of BMP-2 for a long period. Overall, the temperature-responsive Alg-PCLA hydrogels prepared in this study reduced initial burst, and 17

sustained the release of BMP-2. Furthermore, in vitro release test demonstrated that hydrogels exhibited sustained release of BMP-2 (Fig. S3). The BMP-2 was easily mixed with hydrogel solution (sol) through hydrophobic and ionic interactions, and then the mixture is directly injected into the body at specific site, forming an BMP-2-loaded in situ hydrogel. Then, the BMP-2 was released in a sustained manner due to the

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diffusion of medium into the matrix, which loosens the ionic interaction and induced the protein

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release. 3.7. In vivo biomineralization

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To investigate in vivo biomineralization, Alg-PCLA bioconjugate solution and PCLA copolymer

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were injected subcutaneously into SD rats. The minimally invasively injected solution transformed into hydrogel depot at subcutaneous tissues. After 4 weeks, the rats were sacrificed, the recovered

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hydrogel showed fibrous ribbon coating without damage. Before mineralization, SEM images of hydrogels in Fig. 7 demonstrate the microporous structure of hydrogels with smooth surfaces. The

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SEM images show the smooth surface of the hydrogels. After 4 weeks of biomineralization, ribbon-like layers are evident on the surface of the bioconjugate hydrogel. The elemental deposition on hydrogels was determined using the EDS. The results show that the bioconjugate hydrogel contained predominantly Ca and P, suggesting biomineralization. In contrast, PCLA

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hydrogel shows no significant changes in the SEM images. These results indicated that biomimicking acidic extracellular matrix-like fragments present in the bioconjugates facilitated the biomineralization. The strong electrostatic affinity between acid groups of bioconjugate hydrogels and calcium ions accelerated the nucleation and mineral growth.

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The calcium deposition during biomineralization was examined using alizarin red S staining immunohistochemical analysis, because it can specifically bind with calcium ions at low pH. Fig. 8A shows that the areas of subcutaneous tissue (above yellow line) and hydrogel (below yellow line). After 4 weeks, it was interesting to observe dark red dots and dark red ribbons, which indicated calcium deposition inside of hydrogel. In contrast, the PCLA exhibited poor calcium

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deposition. In addition, BMP-2 loaded PCLA hydrogels also showed the similar pattern. Interestingly, the BMP-2 containing Alg-PCLA bioconjugate hydrogel group showed relatively

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high amount of calcium deposition compared to other groups. The Alg-PCLA bioconjugate

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hydrogel containing BMP-2 showed significantly high amount of calcium deposition and interestingly, Alg-PCLA bioconjugate hydrogel without BMP-2 also showed biomineralization to

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some extent.

Furthermore, MT staining was performed to examine collagen protein in newly generated bone. Fig. 8B shows that hydrogels were filled with fibrous tissue after 4 weeks and collagen fibers

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were widely detected by MT staining. In the Alg-PCLA bioconjugate hydrogel group containing BMP-2, the thicker layer composed of fibrous tissue with rich collagen expression (indicated as blue ribbons) inside of hydrogel was compared to other groups. Both the alizarin red S and MT

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staining results declared that Alg-PCLA bioconjugate hydrogel has potential in bone tissue engineering and regeneration.

4. Conclusions We developed an in situ biomineralizable Alg-PCLA bioconjugate for bone repair applications. The bioconjugate exhibited free-flowing sols, and transformed to a viscoelastic gel at the 19

subcutaneous tissue of warm-blooded animals. After incubation of bioconjugate with SBF formed a mineral layers composed of fibrous hydroxyapatite crystals. The MTT assay demonstrated that Alg-PCLA bioconjugate showed low toxicity to 293T cells, with a bioconjugate concentration of up to 1000 µg/mL. The in vivo tests ensure that Alg-PCLA bioconjugate hydrogel can induce in situ biomineralization. Therefore, Alg-PCLA bioconjugate hydrogel could biologically mimic the

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function of extracellular matrix proteins and have promising potential for in situ bone repair.

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Author contributions

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DSL, TT and KSH designed the research. KSH mainly performed all the experiments. VHGP mainly assisted the synthesis of the materials. TT and DSL conducted the analysis and wrote the manuscript.

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Acknowledgments

This research was supported by the Basic Science Research Program through a National Research

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Foundation of Korea grant funded by the Korean Government (MEST) (20100027955), and the National Research Foundation of Korea (NRF) funded by The Ministry of Science, ICT & Future Planning (NRF-2017R1D1A1B03028061).

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Conflict of Interest

The authors declare no conflict of interest.

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Fig. 1. 1H NMR spectra of (A) Alg-PCLA conjugate, (B) Alg-PEA, (C) Alginate, (D) PCLA-NH2, (E) PCLA-NH-Boc, (F) PCLA.

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Fig. 2. (A) GPC trace, (B) ζ of Alg-PCLA conjugates and PCLA copolymers. (C) Viscosity and (D) Sol-to-gel phase transition of Alg-PCLA conjugates and PCLA copolymers.

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Fig. 3. (A) Cell viability of Alg-PCLA conjugates and PCLA copolymers after incubation with 293T cells for 48 h (n = 8). NS indicates no statistical difference between control (0 µg/mL) and 1,000 µg/mL polymer samples. (B) In situ formation of Alg-PCLA conjugate and PCLA copolymer hydrogels. Rats were subcutaneously injected with 22.5 wt.% of either conjugate or copolymer, and the gels were retrieved 10 min after subcutaneous implantation.

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Fig. 4. (A) In vivo biodegradation of Alg-PCLA conjugate and PCLA copolymer hydrogels. (B and C) Hydrogels were withdrawn and lyophilized for visual examination and the extent of degradation was measured using mass loss method. Rats were subcutaneously injected with 22.5 wt.% of either conjugate or copolymer using 26G hypodermic needle.

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Fig. 5. XRD analysis biomineralized Alg-PCLA conjugates after 4 weeks. Hydroxyapatite and Alg-PCLA conjugates without phosphate groups were used for comparison.

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Osteoblasts

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Hydroxyapatite

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Fig. 6. (A) Illustration of in situ gel formation and sustained delivery of BMP-2 from Alg-PCLA bioconjugate hydrogels. (B) Concentration of BMP-2 in the serum of SD rats at different time points. The sampling was done after subcutaneous administration of 300 µL of BMP-2 and BMP2 loaded Alg-PCLA bioconjugate hydrogel into the back of SD rats (n = 3). Asterisks denotes the statistical significant differences (*p<0.05, **p<0.01).

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Fig. 7. SEM images of Alg-PCLA conjugate and PCLA copolymer hydrogels after in vivo biomineralization for 0 and 4 weeks. Enlarged images are for SEM-associated EDS demonstrating the mineralization.

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Fig. 8. (A) Alizarin red S, and (B) MT staining of hydrogel slices collected 4 weeks after subcutaneous administration. (C and D) Quantitative amount of calcium deposition and collagen fibers calculated using ImageJ software analysis. Asterisks denotes the statistical significant differences (*p<0.05).

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Scheme 1. Schematic diagram of extracellular mimicking in situ forming injectable bioconjugate hydrogels loaded with BMP-2 and its sustained release to repair bone defect. Alg-PCLA bioconjugate sols were subcutaneously administered into the back of SD rats, which formed physically cross-linked depot and controlled the release of BMP-2. The extracellular matrix mimicked hydrogel induced in situ biomineralization for tissue repair.

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EDC/NHS DMSO/water

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Scheme 2. Synthesis of Alg-PCLA bioconjugate and PCLA-NH2 copolymer.

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