Injectable, Biomolecule-Responsive Polypeptide Hydrogels for Cell

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Injectable, biomolecule-responsive polypeptide hydrogels for cell encapsulation and facile cell recovery through triggered degradation Qinghua Xu, Chaoliang He, Zhen Zhang, Kaixuan Ren, and Xuesi Chen ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.6b08292 • Publication Date (Web): 20 Oct 2016 Downloaded from http://pubs.acs.org on October 25, 2016

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ACS Applied Materials & Interfaces

Injectable, biomolecule-responsive polypeptide hydrogels for cell encapsulation and facile cell recovery through triggered degradation Qinghua Xu,†, ‡ Chaoliang He,*, † Zhen Zhang,†, ‡ Kaixuan Ren,† Xuesi Chen*, † †

Key Laboratory of Polymer Ecomaterials, Changchun Institute of Applied

Chemistry, Chinese Academy of Sciences, Changchun 130022, P. R. China ‡

University of Chinese Academy of Sciences, Beijing 100039, P. R. China

1

ACS Paragon Plus Environment


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Abstract: Injectable hydrogels have been widely investigated in biomedical applications, and increasing demand has been proposed to achieve the dynamic regulation of physiological properties of hydrogels. Herein, a new type of injectable and biomolecule-responsive hydrogel based on poly(L-glutamic acid) (PLG) grafted with disulfide bond modified phloretic acid (denoted as PLG-g-CPA) was developed. The hydrogels formed in situ via enzymatic crosslinking under physiological condition in the presence of horseradish peroxidase and hydrogen peroxide. The physiochemical properties, including gelation time and rheological property, of the hydrogels were measured. Particularly, the triggered degradation of the hydrogel in response to a reductive biomolecule, glutathione (GSH), was investigated in detail. The mechanical strength and inner porous structure of the hydrogel were influenced by the addition of GSH. The polypeptide hydrogel was used as a three-dimentional (3D) platform for cell encapsulation, which could release the cells through the triggered disruption of the hydrogel in response to the addition of GSH. The cells released from the hydrogel were found to maintained high viability. Moreover, after subcutaneous injection into rats, the PLG-g-CPA hydrogels with disulfide-containing crosslinks exhibited a markedly faster degradation behavior in vivo compared to the PLG hydrogels without disulfide crosslinks, implying an interesting accelerated degradation process of the disulfide-containing polypeptide hydrogels in the physiological environment in vivo. Overall, the injectable and biomolecule-responsive polypeptide hydrogels may serve as potential platform for 3D cell culture and easy cell collection. 2


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Keywords: Biomolecule-responsive hydrogel, polypeptide hydrogel, triggered degradation, enzymatic crosslinking, cell recovery

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1. Introduction In the past few decades, the injectable polymeric hydrogels have received much attention in biomedical applications for their unique properties.1-4 A variety of polymeric materials have been used to construct hydrogels and they can be mainly divided into natural and synthetic polymers according to their origin.5-7 As an emerging synthetic polymer, the synthetic polypeptides exhibit unique advantages, such as good biocompatibility, biodegradability and structures mimicking natural proteins.8-11 Many peptide based hydrogels have been developed and investigated for biomedical applications.12-15 The three-dimensional (3D) polymeric networks of hydrogels are commonly fabricated by chemical or physical crosslinking.16-17 Increasing attention has been paid to enzyme-mediated hydrogelation in recent years due to its mild reaction conditions and fast gelation process.18-25 Horseradish peroxidase (HRP) is an enzyme with good biocompatibility, which is generally used for triggering the intermolecular crosslinking through catalyzing the radical coupling of aniline and phenol moieties in presence of hydrogen peroxide. It has been reported that HRP and H2O2 mediated coupling of phloretic acids facilitated the hydrogelation of different polysaccharide hydrogels.26-29 Moreover, the physiochemical properties of enzymatically-crosslinked hydrogels, such as gelation time and strength, are found to be well controllable. Attributed to the unique properties of injectable hydrogels, the local depot systems containing cells, drugs or bioactive molecules may be facilely fabricated by mixing the precursor solution with cells, drugs or bioactive molecules, followed by delivering to desired sites through minimally invasive injection procedure 4


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and subsequent in-situ sol-gel phase transition. Hydrogels are commonly used as 3D scaffolds for cell culture in vitro and carriers for cell delivery in tissue engineering, due to their physical properties similar to natural extracellular matrix (ECM).30-33 The interconnected 3D network and high water content of hydrogels are beneficial for the diffusion of oxygen and nutrients. It is important to note that controllable degradation rate of hydrogels are highly desirable for precisely matching the tissue regeneration process.34-36 Additionally, after growth and proliferation in vitro, the facile recovery and collection of the amplified cells can also be benefited from a triggered hydrogel degradation in response to a specific stimulus.37 The degradation of hydrogels can be achieved by incorporating various degradable moieties, including degradable polymer backbones, cleavable crosslinks and degradable pendent groups. Hydrolysis and enzymolysis are two kinds of the most widely investigated mechanisms for hydrogel degradation. However, hydrolysis usually follows a passive control process and enzymolysis by specific enzymes secreted by cells may be not ideal when a fast degradation is needed at given time schedule. Recently, the hydrogels that exhibit controllable degradation rates in response to specific stimuli, such as reactive oxygen species (ROS), light and reductants, have received considerable attention.36,

38-43

The biomolecule-triggered degradation of

hydrogels is of special interest for biomedical applications owing to its good biocompatibility. Reduction-responsive scaffolds has been fabricated based on disulfide-crosslinked poly(γ-glutamic acid) and used in tissue engineering 5



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applications.44-46 Disulfide bond is known to be cleavable in the presence of a thiol-containing reducing reagent, e.g., glutathione (GSH).41, 47-49 GSH is found to be the most abundant nonprotein thiol compound in living organisms and plays important roles in cell metabolism.50-51 The production of GSH increases in highly metabolically active cells and the concentration may reach as high as 10 mM, in contrast to only about 10 µM in extracellular environment. As a result, the degradation of hydrogels crosslinked by disulfide bonds may be achieved in an on-demand manner in response to a biological trigger, which is sufficiently mild and biocompatible. In the present study, a novel biomolecule-responsive, enzymatically-crosslinked hydrogel was developed based on poly(L-glutamic acid) (PLG) grafted with disulfide-modified phloretic acid (denoted as PLG-g-CPA). The hydrogels were prepared under physiological condition with the addition of horseradish peroxidase (HRP) and hydrogen peroxide (H2O2). The physiochemical properties of the hydrogels were characterized, including gelation time, rheological property and swelling behavior. Furthermore, the degradation profiles of the PLG-g-CPA hydrogel with the supplement of GSH were emphasized. The hydrogel was found to be quite stable when incubated in phosphate buffer saline (PBS), while the hydrogel degradation was significantly accelerated in the presence of GSH. A continuous decrease in modulus was observed along with the gel degradation in the reductive microenvironment. L929 cells were encapsulated in the hydrogel and maintained high viability. For proof-of-concept, the degradation of the hydrogels was triggered by 6


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addition of GSH, and the encapsulated cells were released and collected. Moreover, the recovered cells were reseeded and the viability was investigated. Additionally, to reveal the degradation of the hydrogel in the physiological environment in vivo, the degradation of the hydrogel in the subcutaneous layer of rats were observed, and the PLG hydrogel without disulfide crosslinks was used as comparison. 2. Materials and Methods 2.1 Synthesis of Cystamine-Conjugated Phloretic Acid (CPA) The synthesis route of CPA is shown in Scheme 1A.52-53 First, one amine group of cystamine

was

protected

by

t-butoxycarbonyl

group.

Briefly,

Cystamine

dihydrochloride (10 g, 0.044 mol) was suspended in 500 mL methanol and triethylamine (20 mL, 0.145 mol) was added. Then di-tert-butyldicarbonate was added to the solution under stirring, and the mixture reacted for 20 min at room temperature. After that, the solvent was evaporated to obtain white solid which was treated with 1 M NaH2PO4 (200 mL, pH 4.2) and extracted with ether for 3 times to remove the residual di-t-Boc-cystamine. The aqueous solution was basified to pH 9 using NaOH and extracted with 100 mL ethyl acetate for 5 times. The organic phase was collected and dried with anhydrous MgSO4, filtered and evaporated to obtain a yellowish oil-like product (compound 1, 3.6 g, 35%). The structure of the product was characterized by 1H NMR and 13C NMR spectra as shown in Figure S1 of supporting information. Compound 1 was coupled with phloretic acid via EDC-NHS activated amidation reaction. Phloretic acid (1.98 g, 0.012 mol), EDC (2.3 g, 0.012 mol) and NHS (1.38 g, 7



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0.012 mol) was dissolved in 100 mL DMF, and the mixture was stirred for 1 h to activate the carboxyl group of phloretic acid. Then t-Boc-cystamine (3 g, 0.012 mol) was added and the resulting solution was reacted at room temperature for 48 h. The solvent was evaporated after the reaction was completed and the product was purified by silica column chromatography using ethyl acetate and petroleum ether (v:v = 2:1) as eluent. The purified product (compound 2, 2.2 g, 46%) was a lightly yellowish viscous solid and measured by 1H NMR spectrum (Figure S1 of supporting information). Compound 2 (2.2 g, 0.0055 mol) was dissolved in the mixture of trifluoroacetic acid (TFA, 20 mL) and dichloromethane (CH2Cl2, 30 mL) to remove the protection t-Boc group. The solution was stirred for 30 min at room temperature and the solvent was evaporated. The residue was dissolved in CH2Cl2, washed with distilled water for 3 times and the aqueous solutions were collected. Then the pH was adjusted to neutral with dilute NaOH solution, followed by extracted with ethyl acetate for 5 times. The organic phases were dried with anhydrous MgSO4, filtered, evaporated and dried under vacuum to get the product (CPA, 1.4 g, 85%，a pale orange solid). The structure of CPA was confirmed by 1H NMR，13C NMR and ES+ MS spectra as shown in Figure 1A, and Figure S2 of supporting information. 2.2

Synthesis

of

poly(L-glutamic

acid)-graft-CPA

(PLG-g-CPA)

and

poly(L-glutamic acid)-graft-tyramine (PLG-g-TA) First of all, poly(L-glutamic acid) (PLG) was synthesized according to the previously reported method, and the structure of the synthesized PLG was characterized by 1H 8


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NMR spectrum (Figure S3 of supporting information)55 The PLG-g-CPA copolymer was prepared by coupling PLG with CPA, as shown in Scheme 1A. PLG (2 g, 15.5 mmol of COOH groups) was dissolved in 30 mL of DMF, and PyBop (0.565 g, 15.5 mmol), DIPEA (189 µL, 31 mmol) as well as CPA (0.326 g, 1.09 mmol) were added to the solution. The resulting mixture was stirred for 2 days at room temperature and the product was purified by dialysis against distilled water for three days (MWCO 3500 Da) which then collected by lyophilization. The final product was white solid with a yield of 90 %. The detailed characterization data of PLG-g-CPA are demonstrated in Figure 1. The PLG-g-TA was synthesized in a similar procedure except for using tyramine (TA) instead of CPA. The feeding ratio of TA to COOH groups in PLG was 10 % and the yield of the product was 85 %. The synthetic route of PLG-g-TA is illustrated in Scheme S1, and the 1H NMR spectrum of the copolymer is shown in Figure S3 of supporting information. 2.3 Dynamic Mechanical Analysis and Interior Morphology of the PLG-g-CPA Hydrogels The storage modulus (G’) of the hydrogels was measured according to the testing procedure of rheological properties of the PLG-g-CPA hydrogels . Then the hydrogel samples were immersed in 1.5 mL PBS or PBS containing 1 mM of GSH, and the G’ values of the hydrogels were recorded after incubated for different time. To investigate the interior morphology of the hydrogel and the hydrogel after culturing with 1 mM GSH for 30 min, the corresponding samples were frozen rapidly 9



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through plunging them into liquid nitrogen and freeze-dried. The obtained specimens were cross-sectioned followed by coating with gold. The morphologies were observed by scanning electron microscopy (SEM, Micrion FEI PHILIPS). 2.4 In Vitro Hydrogel Duration PLG-g-CPA (final concentration and volume: 4 % (w/v), 600µL) and PLG-g-TA (final concentration and volume: 4 % (w/v), 600µL) hydrogels were prepared according to the above method in vials with the inner diameter of 16 mm. Tris-HCl buffer solution (0.05 M, pH 7.4) containing elastase with the concentration of 5 units/mL, 10 mM CaCl2, and 0.2 wt% NaN3 was used as degradation medium. Hydrogels incubated in Tris-HCl buffer solution were used as control. 3 mL of the media were placed on the top of the samples at 37 °C and the remaining hydrogels were weighted at predetermined time intervals. Degradation media were replaced every two days and the experiments were performed in triplicate. 2.5 Cell Encapsulation and Recovery from the PLG-g-CPA Hydrogel L929 cells and rabbit bone marrow-mesenchymal stem cells (rMSCs) were encapsulated into the PLG-g-CPA hydrogels respectively, using the above procedure and cultured for 24 h. Subsequently, the culture medium was replaced by 1 mL of fresh DMEM containing 10 mM of GSH. The hydrogels were disintegrated after incubated for 1 h and the resulting solutions were centrifuged to collect the released cells. Afterward, the cells were reseeded to 96-well plate and cell activity was measured by CCK-8 method. Normal cells seeded on cell culture plate was used as the control. Spreading and proliferation of the released L929 cells and rMSCs were 10


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also observed by microscope and live/dead staining. rMSCs were harvested from one-month-old New Zealand White rabbits.55 The isolated rMSCs were cultured with low-glucose DMEM containing 10% fetal bovine serum, 50 U/mL streptomycin and 50 U/mL penicillin.

3. Results and Discussion Synthesis and Characterization of PLG-g-CPA Copolymer The synthetic strategy of PLG-g-CPA is illustrated in Scheme 1. First of all, in order to prepare the enzymatically-crosslinked hydrogel with cleavable crosslinks, phloretic acid was conjugated with cystamine and the resulting cystamine-conjugated phloretic acid (CPA) containing a disulfide bond was obtained (Scheme 1A). The successful synthesis of CPA was suggested by 1H NMR, as seen in Figure 1A. The signals at δ 6.64 and 6.96 ppm are attributed to the benzene group of phloretic acid, and the signals at δ 3.32 and δ 3.05 ppm belong to the methylene group next to the amine group of cystamine. The signals at δ 6.64, 6.96, 3.32 and 3.05 ppm had an intensity ratio close to 1:1:1:1, conﬁrming the equivalent coupling of phloretic acid and cystamine. In addition, the structure of CPA was also confirmed by 13C NMR and ES MS spectra (Figure S2 of supporting information). The poly(L-glutamic) (PLG) backbone of the hydrogel was synthesized by the ring-opening polymerization of BLG NCA with n-hexylamine as an initiator. The degree of polymerization of PLG was calculated to be ~ 75 according to 1H NMR characterization (Figure S3A of supporting information).

11



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Scheme 1. Synthetic route of (A) CPA (B) PLG-g-CPA copolymer.

The PLG-g-CPA copolymer was then prepared by readily grafting CPA to the carboxyl pendants of PLG via amidation reaction (Scheme 1B). Typical 1H NMR spectrum of the PLG-g-CPA is demonstrated in Figure 1B, and all peaks have been well assigned. The grafting ratio of CPA residues was calculated to be 6% by comparing the integration of peaks for phenyl protons (δ 6.96 and 6.77) with that of the methine peak of PLG (δ 4.69). Additionally, the intensity ratio of peaks at δ 6.96 and 6.77 to peak at δ 3.62 (-CH2NH- on the side chain) was 1:1:2, suggesting that the disulfide bond remained intact during the synthesis process. The number-average molecular weight (Mn) of the copolymer determined by GPC was 8.2 × 103 with a polydispersity (PDI) of 1.57 (Table 1). Conjugation of CPA with PLG was also determined by UV-Vis spectrophotometry. The aqueous solution of PLG-g-CPA showed a characteristic absorbance at 275 nm, corresponding to the typical absorbance of phenol moieties (Figure 1C).20 The grafting ratio of CPA calculated by 12


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the UV test was 6.5%, which was consistent with the value based on 1H NMR. Furthermore, the conformation of PLG-g-CPA in distilled water at pH 7.4 was measured by circular dichroism spectroscopy. The copolymer predominantly adopted a random conformation indicated by a positive maximum at 217 nm and a minimum at 203 nm in Figure 1D. This should be attributed to the fact that most of the carboxyl groups on the side chain remained an ionized status at pH 7.4. In addition, PLG grafted with tyramine (TA) was also prepared, and the synthetic route was shown in Scheme S1 and the 1H NMR spectrum of PLG-g-TA was supplied in Figure S3 in the supporting information.

Table 1. Characterizations of the synthesized PLG-g-CPA and PLG-g-TA copolymers. Code

Copolymers

DP of PLG a

P1

PLG-g-CPA

75

6%

P2

PLG-g-TA

75

9%

a

1

Grafting Grafting ratio a ratio b

b

Mw a

Mn c

PDI c

6.5%

11200

8150

1.57

8.7%

10500

7400

1.4

c

Determined by H NMR, Determined by UV-vis, Determined by GPC

13



DMSO

A g b a

f e

7

6

5

4

h

c

d

3

2

Chemical Shift (ppm) d

B a, b

c

i j

f, g, h

e

7

6

5

4

3

2

Chemical Shift (ppm) 25

C

0.45

0.30

0.15

0.00 260

280

300

320

340

D

0

Ellipticity (mdeg)

0.60

Absorbance

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

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-25 -50 -75 -100 190

200

210

220

230

240

250

Wavelength (nm)

Wavelength (nm)

Figure 1. 1H NMR spectra of (A) CPA in DMSO-d6, and (B) PLG-g-CPA copolymer in CF3COOD. (C) UV-Vis absorbance spectrum of PLG-g-CPA solution. (D) Circular dichroism (CD) spectrum of PLG-g-CPA solution (0.1 mg/mL).

Preparation and Characterization of the Hydrogel The PLG-g-CPA hydrogels were fabricated by the enzyme-mediated crosslinking reaction of PLG-g-CPA through the radical coupling of phenol moieties to form carbon-carbon or carbon-oxygen bonds in the presence of HRP and H2O2. The generation of intermolecular covalent linkages led to the rapid formation of PLG-g-CPA hydrogels. It is established that exogenous H2O2 was used to activate HRP; however, excess H2O2 can inactivate HPR due to the formation of inactive intermediate.21 The PLG-g-CPA solution could not transfer to hydrogel with the 14


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addition of HRP or H2O2 only (Figure S5 of supporting information). In this study, the mole ratio of H2O2 to the phenol group was fixed at 0.3. Accordingly, the gelation time was closely related to the concentration of HRP in the aqueous solution as seen in Figure 2A. The gelation time of the 4 % (w/v) polymer solution (final concentration) decreased from 30 s to 8 s as the concentration of HRP increased from 1 unit/mL to 2 units/mL. The gelation time also depended on the concentration of PLG-g-CPA, and it was obviously decreased from 130 s to 30 s with increasing the PLG-g-CPA concentration from 1 % to 4 % (w/v) (HRP, 1 unit/mL). In consideration of the feasibility of the injection process, the HPR concentration was fixed to be 1 unit/mL in the following study. The rheological properties of the PLG-g-CPA hydrogels were examined by monitoring the variations of loss modulus (G’’) and storage modulus (G’) as a function of time in the process of gel formation. The G’ values rapidly increased during the initial stage, and then reached a plateau (Figure 2B), suggesting almost completion of the crosslinking reaction under the test condition. The mechanical strength of the hydrogels progressively strengthened with the increase of PLG-g-CPA concentration. The G’ of hydrogel consisting of 1% (w/v) polymer was about 40 Pa, while the value markedly increased to 4300 Pa as the polymer concentration increased to 4 % (w/v). This should be attributed to the increase of the concentration of crosslinking units. It demonstrated that the mechanical property of the PLG-g-CPA hydrogels can be easily controlled within a relatively wide range by adjusting the polymer concentration. Moreover, the storage modulus of the PLG-g-TA hydrogel (4 % 15



(w/v), HRP: 1 unit/mL) was nearly 4000 Pa (Figure S6 of supporting information), which was comparable to that of PLG-g-CPA hydrogel.

A

HRP: 1 U/mL, H2O2 / TA: 0.3

125

HRP: 2 U/mL, H2O2 / TA: 0.3

Time (s)

100 75 50 25 0 1%

2%

3%

4%

Concentration (wt%)

B

G', 1% G', 3%

10000

Modulus (Pa)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

G'', 1% G'', 3%

G', 2% G', 4%

G'', 2% G'', 4%

1000 100 10 1 0.1 0

400

800

1200

1600

Time (s)

Figure 2. (A) Dependence of gelation time for PLG-g-CPA hydrogels on polymer and HRP concentrations. (B) Storage modulus (G’) and lose modulus (G’’) of PLG-g-CPA hydrogels with different polymer concentrations at fixed HRP (1 unit/mL) and H2O2 (H2O2/TA = 0.3) concentrations.

In Vitro Degradation of the Hydrogel The swelling ratios of the PLG-g-CPA and PLG-g-TA hydrogels were determined by immersing the freeze-dried hydrogels in 0.01 M PBS at 37 oC, and the remaining hydrogels were weighed at predetermined time intervals. As seen in Figure 3A, the dried gel samples absorbed water rapidly and reached equilibrium swelling in a few 16


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hours. The equilibrium swelling ratio of PLG-g-CPA hydrogels decreased with increasing the polymer concentration of precursor solution which led to an increase in the crosslinking density.

Scheme 2. (A) Schematic illustration of the crosslinking network of PLG-g-CPA hydrogel and triggered degradation of the hydrogel. (B) Photographs for the formation and GSH triggered degradation of the PLG-g-CPA hydrogel.

The structure of the PLG-g-CPA hydrogels and potential breakdown mechanism of the network responding to a reducing agent, e.g., GSH, were schematically illustrated in Scheme 2. The crosslinks containing a disulfide bond can be easily cleaved in the presence of GSH through the thiol-disulfide exchange reaction. Eventually, it would lead to the rupture of the polymeric network and the dissociation of the hydrogel. GSH is a biosynthetic reducing agent and it was used to investigate the degradation dynamics of the hydrogels in this study. Freeze-dried hydrogels were incubated in PBS with different concentrations of GSH, and the swelling ratio as a function of time was recorded. The swelling ratio curve of PLG-g-CPA hydrogels showed a rising phase at the initial stage, and then a descending phase until complete disruption of the 17



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crosslinking network, as demonstrated in Figure 3B. In addition, the degradation process was susceptible to the concentration of GSH. When treated with 0.2 mM GSH, the swelling ratio of 2% (w/v) PLG-g-CPA hydrogel reached the maximum after 2.5 h while it reduced to 0 in 7 h, indicating the complete decomposition of the hydrogel. The increase in the polymer concentration of the PLG-g-CPA hydrogel to 4% (w/v) resulted in the prolongation of the gel degradation time to 11 h. However, as the GSH concentration was increased to 1 mM, both the 2% and 4% (w/v) PLG-g-CPA hydrogels totally decomposed in 2 h, implying an elevated cleavage rate of the disulfide-containing crosslinks. In comparison, the PLG-g-TA hydrogel (4% (w/v)) without cleavable disulfide bond exhibited no obvious degradation in PBS containing 1 mM GSH in the experimental period. In addition to the change in polymer concentration, the control of hydrogel degradation dynamics can also be realized by varying the fraction of disulfide moiety. A series of hydrogels were prepared by mixing PLG-g-CPA and PLG-g-TA together with different ratios (3:1, 2:2, 1:3, w/w), and the total polymer concentration was fixed at 4 % (w/v). The degradation dynamics of these hydrogels in 1 mM of GSH were shown in Figure 3D. All hydrogels kept steady in PBS (Figure 3C), while the situations were totally different in the presence of GSH. As the decrease of disulfide bond content, the time for disruption of hydrogels was prolonged obviously.

18


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A

B

32

75

PLG-g-CPA, 2%, 0.2 mM GSH PLG-g-CPA, 4%, 0.2 mM GSH PLG-g-CPA, 2%, 1 mM GSH PLG-g-CPA, 4%, 1 mM GSH PLG-g-TA, 4%, 1 mM GSH

24

Swelling ratio

Swelling ratio

60

16 PLG-g-CPA, 2%, PBS PLG-g-CPA, 4%, PBS PLG-g-TA, 4%, PBS

8

45 30 15

0

0 0

15

30

45

60

75

0

2

4

Time (hours)

6

8

10

Time (hours)

D

C

50

Swelling ratio

32

Swelling ratio

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


24

16

(PLG-g-CPA:PLG-g-TA) 4:0 3:1 2:2 1:3

8

40 30 (PLG-g-CPA:PLG-g-TA) 4:0 3:1 2:2 1:3

20 10 0

0 0

5

10

15

20

25

0

5

10

15

20

25

Time (hours)

Time (hours)

Figure 3. The swelling ratios of PLG-g-CPA and PLG-g-TA hydrogels (A) in PBS and (B) in PBS containing different concentrations of GSH. The swelling ratios of hydrogels formed by mixed PLG-g-CPA and PLG-g-TA with different ratios (total polymer concentration: 4% (w/v), HRP:1 unit/mL , H2O2: 5 mM) (C) in PBS and (D) in PBS containing 1 mM GSH.

The mechanical property of hydrogels plays an important role in clinical applications, such as cell delivery, stem cell differentiation and tissue regeneration.36, 57-58

Tunable strength of PLG-g-CPA hydrogel could be achieved via the controllable

degradation of hydrogel in various extents. Hydrogels incorporating disulfide linkages are of particular interest, because the degradation process can be tailored by the addition of GSH under physiological condition. More importantly, the degradation can also be autonomously and partially controlled by the cells encapsulated within the 19



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hydrogels, which are able to secrete and release the reducing agents. As shown in Figure 4A, the storage modulus of PLG-g-CPA hydrogel (4% (w/v)) continuously decreased by treating with GSH. However, the hydrogel was found to be stable in PBS, and the slight decrease in storage modulus after incubating in PBS should be resulted from the swelling of the hydrogel. These data also confirmed that GSH aided the cleavage of the crosslinking network and speeded up the degradation of the hydrogels. The interior morphologies of the hydrogels treating with PBS with 0 or 1 mM GSH for 30 min were observed via scanning electron microscopy (SEM) (Figure 4B). The PLG-g-CPA hydrogel demonstrated a porous three-dimensional network structure. It was found that much more pores distributed on the surface of the hydrogel network after immersing in PBS containing 1 mM GSH, indicating the erosion of the hydrogel. Except for the GSH-responsive property, the PLG-g-CPA hydrogels containing a PLG backbone can also be broken down by protease.59 To investigate the in vitro duration behavior, the PLG-g-CPA and PLG-g-TA hydrogels were incubated in Tris-HCl

buffer

solution

containing

elastase

(5

units/mL).

The

enzymatically-crosslinked hydrogels were very stable in buffer solution without elastase and the mass of the hydrogels almost remained unchanged for 3 weeks. Whereas, both of the hydrogels were gradually degraded by treating with elastase, which was due to the breakage of the peptide bonds on the main chains and finally totally disintegration of the polymer network (Figure 5).

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Figure 4. (A) Storage modulus (G’) of the PLG-g-CPA (4% (w/v)) hydrogels after treating with PBS and PBS with 1 mM GSH for different time. (B) SEM images of the PLG-g-CPA (4% (w/v)) hydrogels after treating with PBS and PBS with 1 mM GSH for 30 min. 200

PEG-g-CPA, Tris-HCl PEG-g-TA, Tris-HCl

PEG-g-CPA, Elastase PEG-g-TA, Elastase

160

Mass remaining (%)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


120

80

40

0 0

5

10

15

20

Tiem (days)

Figure 5. In vitro mass loss profiles of the in situ formed PLG-g-CPA and PLG-g-TA hydrogels (polymer concentration: 4% (w/v)) in 0.05 M Tris-HCl buffer (pH 7.4) and Tris-HCl buffer containing 5 units/mL elastase.

21



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Cytocompatibility of the Copolymer and Recovery of Cells from the Hydrogel The cytocompatibility of the PLG-g-CPA copolymer was evaluated using MTT assay against L929 cells. L929 mouse fibroblasts were incubated with PLG-g-CPA copolymer at concentrations ranging from 0 to 2 mg/mL for 24 h. The relative cell viabilities were demonstrated in Figure 6A. All experiment groups showed no detectable cytotoxicity toward L929 cells (cell viability > 85%), suggesting the good biocompatibility of the synthesized PLG-g-CPA copolymer. Additionally, the supernatants collected after the PLG-g-CPA hydrogels were extracted for 1, 3 and 5 days all showed good cytocompatibility (cell viability > 85%, Figure 6B). To assess the feasibility of using the enzymatically-crosslinked PLG-g-CPA hydrogel as the scaffold for cell survival, L929 cells were encapsulated into the in situ formed hydrogel for a 3D cell culture test. The cell viability was determined by cell counting kit-8 method. It was found that nearly all the cells remained viable after 24 h culturing (Figure 6C). The cell survival status was further evaluated by a live-dead cell staining kit. As seen in Figure 6D, most of the cells within the hydrogel were stained green with calcein-AM, which confirmed the high viability of the L929 cells and the desirable cytocompatibility of the hydrogel. Moreover, cells were stained with DAPI after cultured in the hydrogel for 2 h, and observed by the confocal laser scanning microscope (CLSM). Cells were distributed uniformly within the hydrogel as shown in both the stacked 3D image and the longitudinal section image of the cell loaded hydrogel, which suggested that the PLG-g-CPA hydrogel could act as 3D-culture platform (Figure S8 of supporting information). 22


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Figure 6. In vitro cytocompatibility of (A) PLG-g-CPA copolymer, (B) eluants of PLG-g-CPA hydrogels at 1, 3 and 5 days against L929 cells measured via MTT assay (n=3). Cells were incubated with the copolymer or the eluants for 24 h. (C) Viability of L929 cells in the PLG-g-CPA hydrogel analyzed by the CCK-8 method (n=3). (D) Live-dead assay of L929 cells cultured in the PLG-g-CPA hydrogel for 24 h. Live cells stained with calcein-AM show green fluorescence and dead cells stained with PI show red fluorescence.

Next, we investigated whether it was feasible to recover viable cells from the hydrogel with disulfide-containing crosslinks.37, 60 The phase and internal structure of the PLG-g-CPA hydrogel underwent quick transition by treating with GSH according to the above characterizations. After culturing for 24 h in normal culture media, the cell-laden hydrogel was then incubated with fresh DMEM containing 10 mM GSH for another 1 h, and the entire hydrogel completely dissolved during this process. The mixtures were transferred to centrifuge tube and cells released from the hydrogel were 23



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collected by centrifuging. The obtained cells were resuspended in cell culture media and then seeded onto cell culture plate to measure the cell viability. It was observed that, the cells adhered and spread on the plate after re-incubated for 4 h, and the cells nearly proliferated into a confluent layer after 24 h (Figure 7C). Moreover, the high activity of the L929 cells released from the PLG-CPA hydrogel was also verified by compared to normal L929 cells via CCK-8 assay (Figure 7A). rMSCs were also encapsulated into the PLG-g-CPA hydrogels and released after 24 h of culture. The recovered rMSCs had activity comparable to the normal rMSCs cells as well (Figure 7B and 7D). These results clearly indicated that GSH-triggered decomposition process of the PLG-g-CPA hydrogel did not affect the cell viability, and the cells recovered from the 3D matrix could be used for further biological applications. It should be noted that the conditions employed for cell recovery had marked influence on the cell activity. Both the concentrations of thiol component and the exposure time for gel dissolving should be taken into consideration. Based on the results in this study, the PLG-g-CPA hydrogel was found to be sensitive enough and the fast GSH-triggered degradation process displayed good biocompatibility.

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Figure 7. L929 cells and rMSCs were released and reseeded onto cell culture plates after encapsulated into PLG-g-CPA hydrogels for 24 h. The proliferation ability of normal control cells and released cells, (A) L929 cells and (B) rMSCs. Spreading and proliferation of the released (C) L929 cells and (D) rMSCs after reseeding for different period.

In Vivo Hydrogel Degradation and Biocompatibility In order to determine the in vivo degradation rates and biocompatibility of the PLG-g-CPA and PLG-g-TA hydrogels, 500 µL of copolymer solutions in PBS containing H2O2 and HRP were injected subcutaneously at the back of rats. Hydrogels were rapidly formed after injection (Figure 8A). The rats were sacrificed at various time intervals, and it was observed that the PLG-g-CPA hydrogel exhibited a much faster degradation rate compared to the PLG-g-TA hydrogels without disulfide crosslinks (Figure 8A and Figure S9 of supporting information). It was found that 25



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the in situ formed PLG-g-CPA hydrogels in the subcutaneous layer completely degraded in 12 days. However, the complete degradation of the PLG-g-TA counterpart took more than one month in vivo. The intracellular GSH concentrations are in the range of 0.5 to 10 mM. Meanwhile, GSH as an intermediate of metabolism widely distributes in fluids and extracellular matrices with relatively low concentrations. The GSH concentration ranges from 2 to 20 µM in plasma, and it is about 4 µM in subcutaneous tissue of rats as reported.61 Thus, the results of in vivo degradation suggested that the PLG-g-CPA hydrogel could respond to the low GSH concentration in subcutaneous layer, leading to the significant difference in the degradation rate between PLG-g-CPA and PLG-g-TA hydrogels. Additionally, the skins attached to the PLG-g-CPA and PLG-g-TA hydrogels were studied by hematoxylin and eosin (H&E) staining to investigate the biocompatibility of the hydrogels in vivo (Figure 8 and Figure S10 of supporting information). As seen in Figure 8B, elevated numbers of inflammatory cells were observed during the degradation process (e.g., day 3 and day 6), however, the slight inflammatory reaction was eliminated with the disappearance of the hydrogels. At the same time, no obvious edema, hyperemia or tissue necrosis was observed during the experimental process. The results indicated that the PLG-g-CPA hydrogel demonstrated acceptable biocompatibility in vivo, and this confirmed its potential applications in biomedical and biotechnology fields.

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Figure 8. (A) In vivo hydrogel status at different time intervals. PLG-g-CPA hydrogels (0.5 mL, 4% (w/v)) were injected into the back of rats. Photos were taken at 15 min (0 day) and 3, 6, 12 days after injection. (B) Tissue biocompatibility of the in situ formed hydrogels (H&E staining of tissues surrounding the injection sites).

4. Conclusion In this study, a novel type of enzymatically-crosslinked and biomolecule-responsive polypeptide hydrogel was developed. The injectable hydrogel based on the synthesized PLG-g-CPA copolymer demonstrated a triggered degradation behavior, and degradation dynamics of the hydrogel can be regulated effectively. The mechanical property and inner structure of the hydrogels can be well controlled by adjusting the hydrogel degradation. Meanwhile, the hydrogels showed good cytocompatiblity when used as a platform for 3D cell culture. The cells encapsulated within the hydrogel can be released in a facile and triggered manner, which kept a 27



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high viability. Furthermore, the in vivo degradation tests suggested that the hydrogel displayed good biocompatible in vivo and was capable of responding to the reducing agents in vivo at physiologically relevant concentration levels. In summary, this biomolecule-responsive hydrogel may have potential applications as platforms for 3D cell culture, tissue repair and regeneration medicine.

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Associated Content Supporting Information Materials, reagents and characterizations; the testing methods of gelation time, rheological properties, swelling behavior, in vitro cytocompatibility, and in vivo degradation and biocompitibility of the PLG-g-CPA hydrogels; synthetic route of PLG-g-TA copolymer; 1H NMR spectra of two intermediate products during the synthesis of CPA, PLG and PLG-g-TA, and and intermediate;

13

13

C NMR spectrum of the first

C NMR and ES+ MS spectra of CPA; complex viscosity change of

PLG-g-CPA solution during the solution to hydrogel transition process; control experiments of hydrogel formation; the stacked 3D image and longitudinal section image of cell loaded PLG-g-CPA hydrogel; rheological property of PLG-g-TA hydrogel (4% (w/v)); In vivo degradation and the tissue compatibility of PLG-g-TA hydrogels (4% (w/v)). This material is available free of charge via the Internet at http://pubs.acs.org.

Author Information Corresponding Author *Email: [email protected]; [email protected] Author Contributions The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. 29



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Acknowledgements This work was supported by the National Natural Science Foundation of China (projects 21574127, 51622307, 51390484 and 51321062) and the authors are grateful for the financial support.

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Injectable, Biomolecule-Responsive Polypeptide Hydrogels for Cell

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