Fabrication and characterization of collagen-based injectable and self-crosslinkable hydrogels for cell encapsulation

Accepted Manuscript Title: Fabrication and Characterization of Collagen-based Injectable and Self-crosslinkable Hydrogels for Cell Encapsulation Authors: Yongli Gao, Weili Kong, Bao Li, Yilu Ni, Tun Yuan, Likun Guo, Hai Lin, Hongsong Fan, Yujiang Fan, Xingdong Zhang PII: DOI: Reference:

S0927-7765(18)30213-3 https://doi.org/10.1016/j.colsurfb.2018.04.009 COLSUB 9262

To appear in:

Colloids and Surfaces B: Biointerfaces

Received date: Revised date: Accepted date:

24-1-2018 2-4-2018 3-4-2018

Please cite this article as: Yongli Gao, Weili Kong, Bao Li, Yilu Ni, Tun Yuan, Likun Guo, Hai Lin, Hongsong Fan, Yujiang Fan, Xingdong Zhang, Fabrication and Characterization of Collagen-based Injectable and Selfcrosslinkable Hydrogels for Cell Encapsulation, Colloids and Surfaces B: Biointerfaces https://doi.org/10.1016/j.colsurfb.2018.04.009 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.

Fabrication and Characterization of Collagen-based Injectable and Self-crosslinkable Hydrogels for Cell Encapsulation Yongli Gao, Weili Kong, Bao Li, Yilu Ni, Tun Yuan, Likun Guo*, Hai Lin, Hongsong Fan, Yujiang Fan, Xingdong Zhang

A

CC E

PT

ED

M

A

N

U

SC R

IP T

National Engineering Research Center for Biomaterials, Sichuan University, 29 Wangjiang Road, Chengdu, Sichuan 610064, China * Corresponding author. National Engineering Research Center for Biomaterials, Sichuan University, 29 Wangjiang Road, Chengdu, Sichuan 610064, China. E-mail address: [email protected]

1

N

U

SC R

IP T

Graphica abstarct

ED

M

A

Design and fabrication of collagen-based injectable and self-crosslinkable hydrogels for cell delivery. （CS-sNHS: activated chondroitin sulfate; -NH2: amino groups; -CO-sNHS: activated carboxyl groups; -CONH: amide groups）

Highlights

ColI/CS-sNHS hydrogels were injectable and self-crosslinkable.

ii

The gelation occurred without the addition of any crosslinking agent.

iii

The physical properties of hydrogels were tuned by regulating the DS of CS-sNHS.

iv

ColI/CS-sNHS hydrogels supported the survival and ECM secretion of chondrocytes.

v

The hydrogels exhibited acceptable inflammatory responses.

A

CC E

PT

i

Abstract Injectable and self-crosslinkable hydrogels have drawn much attention for their potential application as cell delivery carriers to deliver cells to the injury site of 2

arbitrary shape. In this study, injectable and self-crosslinkable hydrogels were designed and fabricated based on collagen type I (Col I) and activated chondroitin sulfate (CS-sNHS) by physical and chemical crosslinking without the addition of any catalysts. The physical properties of hydrogels, including mechanical properties, swelling and degradation properties, were investigated. The results demonstrated that the physical properties of hydrogels, especially the stiffness of hydrogels, were readily tuned by

IP T

varying the degree of substitution (DS) of CS-sNHS without changing the concentration of collagen-based precursor. Chondrocytes were encapsulated into

SC R

hydrogels to investigate the effects of hydrogels on the survival, proliferation and

extracellular matrix (ECM) secretion of cells by FDA/PI staining, CCK-8 test and histological staining. The results suggested that all of these hydrogels supported the

U

survival and ECM secretion of chondrocytes, while there was more ECM secretion

N

around chondrocytes encapsulated in hydrogel Col I/CS-sNHS56% in which the DS of

A

CS-sNHS was 56%. When the neutral precursor solution for hydrogel of Col I or Col I/CS-sNHS56% was subcutaneously injected into SD rats, hydrogels both displayed

M

acceptable biocompatibility in vivo. These results imply that these injectable and

ED

self-crosslinkable hydrogels are suitable candidates for applications in the fields of cells delivery and tissue engineering.

PT

Keywords

Injectable hydrogel; Self-crosslinkable hydrogel; Tunable mechanical property; Cell

A

CC E

encapsulation; Tissue engineering

Introduction Cell therapy is considered to be a promising method of treating various diseases. Cell retention and poor cell survival are limiting the development of traditional cell 3

therapy [1]. Fortunately, some researchers have reported that using cell-compatible biomaterials as cell delivery carriers can solve such problems [2, 3]. Over the past few decades, injectable hydrogels have drawn much attention for their potential application as cell delivery carriers. Injectable and self-crosslinkable hydrogels are superior to the preformed scaffolds in improving patient compliance and overcoming the risk of implant migration [4]. Such hydrogels usually retain large

IP T

amounts of water and exhibit excellent permeability to nutrients and metabolites [5]. Moreover, cells and biologically active molecules can be premixed homogeneously

SC R

into these hydrogels thereby facilitating their uniform incorporation into hydrogels [6]. Minimally invasive surgery makes it possible to delivery hydrogels to the required position with arbitrary shape [7].

U

Typically, the hydrogel precursors are injected and the sol-gel process is achieved

N

in situ by photo-induced reactions, chemical or enzymatic crosslinking reactions, or by

A

altering the physical conditions, such as temperature, pH [4, 7]. However, most of these injectable hydrogels encapsulate cells by drastic changes in environmental conditions

M

or using some toxic organic reagents which are more or less harmful for the

ED

encapsulated cells [8]. An ideal injectable and in situ-forming hydrogel should gel under mild conditions. Furthermore, it should have adjustable physical properties, such as mechanical strength, to meet the needs for its application [9, 10]. Natural polymers

PT

have excellent biocompatibility and biodegradability. Collagen, the major component of ECM, has cell adhesion sites which are conducive to cell proliferation and

CC E

differentiation [11]. Thus far, as an attractive biomaterial for bionic construction of ECM-like microenvironment, collagen type I (Col I) is widely applied in cartilage tissue engineering due to its good biocompatibility and low immunogenicity [12]. Col I

A

self-assemble into hydrogels at physiological conditions (pH 7.4 and 37 ℃ ) by hydrogen bonding, but these spontaneous hydrogels are weak in mechanical strength and will be rapidly degraded by enzymatic hydrolysis [13]. Moreover, as a scaffold for chondrocytes delivery, studies have shown that matrix stiffness can affect the alteration of chondrocyte ECM synthesis [14]. It is obvious that mechanical property is an important factor for the design and fabrication of scaffolds for cartilage tissue 4

engineering [15]. Chondroitin sulfate (CS), a major mucopolysaccharide component in natural cartilage tissue, is in capacity of regulating the swelling pressure for matrix expansion and collagen network tension [10]. In addition, CS can interact with the proteins in ECM and affect the cellular response owing to its biological characteristics, such as hydration of the ECM and binding of effector molecules (e.g. growth factors and

IP T

cytokines) [12, 16]. Some experimental and clinical trials have reported that CS plays a key role in tissue remodeling and regeneration [17].

SC R

Combining the advantages of Col I and CS, in this study, the hydrogels were

designed and fabricated based on the physical crosslinking of Col I and the chemical crosslinking of Col I and activated CS (CS-sNHS). The physical properties of

U

hydrogels were adjustable, and the effects of hydrogels on behaviors of chondrocytes

N

were investigated. Furthermore, the biocompatibility of hydrogels was performed in an

A

animal model.

M

2. Materials and Methods

ED

2.1. Materials

Col I was extracted from newborn calf skin and dissolved in pH3.0 hydrochloric acid solution for use according to the previously published method [18]. Chondroitin

PT

sulfate sodium salt from shark cartilage (CS, 95%), 1-ethyl-3-(3-dimethyl aminopropyl) carbodiimide hydrochloride (EDC•HCl, 99%) and N-hydroxy sulfosuccinimide

CC E

sodium salt (sNHS, 98%) were purchased from Best-reagent Corporation (Chengdu, China). TESCA buffer was purchased from Sigma-Aldrich (USA).

A

2.2. Modification of CS with sNHS CS was activated by sNHS to obtain carboxyl activated CS intermediates.

Specifically, CS powder was dissolved in MilliQ water to form a clarifying solution at a concentration of 5% (m/v). Then the appropriate sNHS solution was added slowly under stirring at pH 4.75. Subsequently, EDC•HCl solution was added slowly and maintained the pH value at 4.75 at room temperature. After 2.5h reaction, pH value of 5

the mixture was adjusted to 7.4 to terminate the reaction. The reaction product was precipitated with ethanol, and the deposit was collected. The deposit was further purified by dialyzing against MilliQ water through a seamless cellulose tube (cutoff Mw, 3500) at 4℃ for 5days, and the MilliQ water was changed once a day. Finally, the purified polymer was freeze-dried by vacuum freeze drier (VirTis) and referred to as CS-sNHS. By adjusting the ratios of CS, sNHS and EDC•HCl, CS-sNHS with different

IP T

DS was obtained.

SC R

2.3. Characterization of CS-sNHS

To calculate the DS of carboxyl groups in CS-sNHS, tyramine (TA) was reacted with CS-sNHS. Specifically, 200mg CS-sNHS was dissolved in 10mL MilliQ water.

U

Then 48.84mg TA was added to the solution at pH 6.0, and the reaction proceeded for 3h at room temperature. The purification and drying process were the same as

N

described above, and the obtained polymer was named as CS-TA. 5mg CS-TA was

A

dissolved in 0.5mL D2O, and 1H-NMR spectrum was recorded by Bruker AV II running

M

at 400MHz.

ED

2.4. Fabrication of Col I/CS-sNHS hydrogels Col I/CS-sNHS hydrogels were prepared at physiological condition (pH 7.4, 37℃)

PT

without addition of any catalyst or crosslinking agent. The weight ratio of Col I and CS-sNHS was 4:1, and their total concentration was 10mg/mL. Concretely, CS-sNHS

CC E

was dissolved in PBS at pH 7.4 to prepare 24mg/mL CS-sNHS solution. In the ice water bath, 1mL acidic collagen solution was neutralized by NaOH solution, and subsequently CS-sNHS solution was added to obtain a Col I/CS-sNHS hydrogel

A

precursor solution. Finally, hydrogel precursor solution was injected into a cylindrical mold (8mm in diameter, 2mm in thickness) and placed in an electric thermostatic water bath at 37℃ for hydrogel formation. Collagen hydrogel was prepared with neutral collagen solution at the same conditions and used as a control. 2.5. Characterization of Col I/CS-sNHS hydrogels 2.5.1. Gelation time test 6

Gelation time of hydrogels was determined by rotational rheometer (TA DHR-2, America). The test geometry was a 40mm diameter plate with 1.0mm gap. The dynamic strain sweep was operated with constant strain of 1% and frequency of 10Hz at time sweep in 400s. Meanwhile, an integrated temperature controller was used to maintain the temperature at 37℃. The gelation time was recorded. 2.5.2. Morphology observation

IP T

Microstructure of hydrogels was observed by scanning electron microscopy (SEM, S-4800, Hitachi). The hydrogels were frozen quickly in liquid nitrogen and lyophilized

SC R

in a vacuum freeze drier for 48h. The lyophilized samples were sectioned and coated with an ultrathin layer of gold/Pt in ion sputter, and then observed by SEM. 2.5.3. Mechanical test

U

The mechanical properties of hydrogels were carried out using dynamic

N

mechanical analyzer (DMA, TA-Q800, USA) at room temperature. The test was

performed to obtain averaged values.

A

measured at an amplitude of 20µm with 5mN prestress. Three parallel hydrogels were

M

2.5.4. Degradation and swelling measurement

ED

For degradation test, the hydrogels were weighed (W0) and then were incubated in 1mL TESCA buffer (50mM TES, 0.36mM calcium chloride, pH 7.4) containing 50µg/mL collagenase I at 37℃. At predetermined time intervals, the hydrogels were

PT

taken out and weighed (Wt). Three parallel hydrogels were performed to obtain averaged values. The degradation behavior of the hydrogel was expressed according to

CC E

the following formula:

Weight loss (%) = (W0-Wt)/ W0 × 100% For the swelling test, the hydrogels were weighed (W0) and then were immersed in

A

PBS at 37℃. The swollen hydrogels were removed at regular intervals and weighed (Wt). The swelling ratio of hydrogels was calculated according to the following formula: Swelling ratio (%) = (Wt-W0)/ W0 × 100% 2.6. In vitro cell culture in hydrogels 7

2.6.1. Chondrocyte isolation and encapsulation Chondrocytes were isolated from articular cartilages of newborn rabbits according to the previously published method [19]. The newborn rabbits were purchased from the Laboratory Animal Center of Sichuan University (Chengdu, China). The experiment was approved by the Animal Care and Use Committee of Sichuan University. The second passage of chondrocytes were harvested, centrifuged and mixed

IP T

gently with the sterile precursor solution of hydrogels at a final cell density of 107

cells/mL. The neutral mixture was injected into cylindrical molds and incubated at 37℃

SC R

for hydrogel formation. Then cell/hydrogel constructs were taken out from the molds and cultured in cell culture dishes. The culture medium was changed once in three days. 2.6.2. Live/dead staining

U

At predetermined intervals, the cell/hydrogel constructs were collected and

N

assessed by live/dead staining and observed by the confocal laser scanning microscope

A

(CLSM, Leica-TCS-SP5). Cell/hydrogel constructs being cultured for 1, 7 and 14days were washed with PBS and stained with the mixed solution containing 1μg/mL

M

fluorescein diacetate (FDA) and 1μg/mL propidium iodide (PI) for 3min. After being

2.6.3. CCK-8 test

ED

rinsed with PBS, the constructs were observed with CLSM.

At predetermined intervals, the original medium in the well of cell culture plate

PT

was removed and 1mL/well of fresh serum-free medium with 10% CCK-8 was added and incubated for 3h at 37℃. The absorbance was measured at 450nm using a

CC E

Microplate Reader (Multiskan FC, America). 2.6.4. Histological assessment At predetermined intervals, the cell/hydrogel constructs were collected, washed

A

with PBS and fixed immediately in 4% paraformaldehyde solution for 2days. Then the constructs were dehydrated in gradient ethanol series, embedded in paraffin wax, sectioned with 5μm thickness, and stained for histological evaluation. 2.7. In vivo degradation and biocompatibility of hydrogels Sprague-Dawley (SD) rats (~200g) were used for in vivo degradation and 8

biocompatibility tests. Then 0.5mL Col I/CS-sNHS56% precursor solution was subcutaneously injected on the left side of the rat's back. Col I precursor solution was injected on the other side of the rat's back and used as a control. At predetermined intervals, rats were sacrificed by neck breaking method and the gel status was observed. Hydrogels together with their surrounding tissues were surgically removed and stored in 4% paraformaldehyde solution for 1week. Histological analysis was performed by

IP T

hematoxylin&eosin staining.

SC R

3. Results and discussion 3.1. Synthesis and characterization of CS-sNHS

The synthetic scheme of CS-sNHS and CS-TA was illustrated in Fig.1a, and their

U

chemical structures were confirmed by 1H-NMR spectrum as shown in Fig.1b.

N

CS-sNHS was synthesized by CS and sNHS at the presence of EDC•HCl, and then

A

purified by dialysis against MilliQ water. CS-sNHS was stable and soluble, especially they were active and could react with some polymers directly without any catalysts. It

M

was difficult to calculate the DS of CS-sNHS from its 1H-NMR spectrum due to the

ED

overlapping of protons in sNHS and CS. Protons in benzene ring could be distinguished from protons in CS main chains easily, so TA was reacted with CS-sNHS to calculate the DS of CS-sNHS indirectly. The DS of CS-sNHS was determined from 1H-NMR

PT

spectrum of CS-TA by comparing the integral area of the methyl protons in CS at 1.9ppm with that of the phenyl protons in TA at 6.7 and 7.1ppm according to the

CC E

following formula: DS=3y/4x. The DS of CS-sNHS was regulated by changing the proportion of CS, sNHS and EDC•HCl. The composition, DS and nomenclature of CS-sNHS were shown in Table 1a. In this study, CS-sNHSs with gradient DS were

A

prepared to regulate physical properties of hydrogels.

Table 1a. The composition, degree of substitution and nomenclature of activated CS-sNHS.

9

Molar ratio of

Degree of substitution

CS:sNHS:EDC•HCl

of CS-sNHS(%)

Nomenclature

1:0.8:0.8

38

CS-sNHS-1

1:1:1

56

CS-sNHS-2

1:1:2

83

CS-sNHS-3

Weight ratio Composition gelation time/(s)

Nomenclature

SC R

of Col I and

IP T

Table 1b. The composition, gelation time and nomenclature of hydrogels.

of hydrogels CS-sNHS 1:0

200

hydrogel S-1

Col I/CS-sNHS-1

4:1

194

hydrogel S-2

Col I/CS-sNHS-2

4:1

163

hydrogel S-3

Col I/CS-sNHS-3

4:1

N

U

Col I

hydrogel S-4

CC E

PT

ED

M

A

151

A

Fig.1. Synthetic scheme of CS-sNHS and CS-TA (a), and 1H-NMR spectra of CS and CS-TA (b).

3.2. Preparation of Col I/CS-sNHS hydrogels In this study, the composition, gelation time and nomenclature of hydrogels were shown in Table 1b. The DS of CS-sNHS increased in the order of hydrogel S-2, S-3 and S-4 and the hydrogel S-1 was as a control. Collagen injectable hydrogel formed by 10

self-assemble, which was a kind of physical-crosslinked hydrogel. Collagen self-assembled into nanofibers when temperature of neutral collagen solution increased to 37℃, meanwhile transparent collagen solution transformed into opaque collagen hydrogel as shown in Fig.2-i. When CS-sNHS was introduced into neutral collagen solution, there were physical and chemical crosslinking occurrence based on self-assemble of collagen and chemical reaction of amino-groups in collagen with

IP T

activated carboxyl groups in CS-sNHS, respectively. Rheological results displayed that

all of these hydrogels formed within 4min. No crosslinking agent was applied during

SC R

fabrication of Col I/CS-sNHS hydrogels. With the increase of DS of CS-sNHS, transparency of Col I/CS-sNHS hydrogels was enhanced due to the decrease of physical crosslinking degree and the increase of chemical crosslinking ratio in

U

hydrogels (Fig.2-ii).

N

3.3. Characterization of Col I/CS-sNHS hydrogels

A

3.3.1. Morphology observation

M

The inner morphology of hydrogels was observed by SEM and their images with high and low magnification were shown in Fig.2-iii, continuous and fibrous network

ED

structure was observed in all hydrogels. Fibers in collagen hydrogel showed a smooth and round morphology (Fig.2-iii a and e), while it seemed that fibers in Col I/CS-sNHS

PT

hydrogels were covered by something unevenly (Fig.2-iii b-d and f-h), and it was more obvious with the increase of DS of CS-sNHS even with the same ratio of Col I and

CC E

CS-sNHS in hydrogels.

Different morphology of these hydrogels was mainly caused by their crosslinking

manners. Collagen hydrogel formed by self-assemble and collagen molecules could

A

freely adjust their structure and position to achieve stable equilibrium state. However, chemical reaction limited the movement of collagen molecules and further limited the self-assemble of collagen in Col I/CS-sNHS hydrogels. The higher the DS of CS-sNHS was, the more obvious the limitation was, which lead to the differences in the morphology of hydrogels. The interconnected fibrous network structure might have an active and comparable effect on cell behaviors and transportation of nutrition and 11

SC R

IP T

metabolites [20, 21].

M

3.3.2. Mechanical property

A

N

U

Fig.2. (i) The gelation phenomenon of neutral Col I solution. The neutral transparent solution became milky opaque hydrogel at 37℃. (ii) Optical images of hydrogels S-1, S-2, S-3 and S-4 in the order of left to right. (iii) SEM images of hydrogels S-1 (a, e), S-2 (b, f), S-3 (c, g) and S-4 (d, h) with low (a-d) and high magnification (e-h). The scale bar is 10µm for images (a-d) and 1µm for images (e-h).

ED

Mechanical properties of hydrogels are essential for their application in tissue engineering. It has been reported that the stiffness of hydrogels has a significant influence on cell behaviors, such as proliferation, migration, differentiation and

PT

cell-dependent processes [22-25]. Mechanical properties of hydrogels are closely related to their component, crosslinking manners, crosslinking degree, water content,

CC E

and so on [26].

In this study, the storage modulus (G′) and loss modulus (G″) of hydrogels were

detected, and the results were shown in Fig.3. It demonstrated that the higher the

A

storage modulus of hydrogels was, the lower the loss modulus of hydrogels was (Fig.3a, 3b). Loss tangent is defined as tanδ = G″/ G′. The magnitude of the loss tangent represents the viscoelastic properties of the hydrogels. The smaller the loss tangent is, the greater the elasticity of the hydrogels is, which means that the hydrogels are more suitable for application as cartilage repair scaffolds. Compared to hydrogel S-1, hydrogel S-3 and S-4 had lower loss tangent (Fig.3c), which indicated 12

that the mechanical properties of collagen hydrogels were effectively improved by the addition of CS-sNHS. Furthermore, storage modulus of hydrogels gradually increased with the increase of DS of CS-sNHS. In these hydrogels, the more activated carboxyl groups meant the higher degree of chemical crosslinking density, while crosslinking density was a factor to regulate mechanical properties of hydrogels [36]. In this research, mechanical properties of hydrogels were regulated by their crosslinking

PT

ED

M

A

N

U

SC R

IP T

manner and crosslinking degree.

CC E

Fig.3. The mechanical properties of storage modulus (a), loss modulus (b) and loss tangent (c) of hydrogels measured by dynamic mechanical analyzer. Data are expressed as means ±SD (n = 3).

A

3.3.3. Water absorption ability Water absorption ability is an important factor for design and fabrication of

cartilage scaffolds. Scaffolds with higher water absorption ability would benefit for transportation of nutrients and metabolites and maintenance of scaffold resiliency [1]. There are some factors affected water absorption ability of scaffolds, such as types of crosslinker, crosslinking degree, and hydrophilic groups in scaffolds materials and so 13

on [27, 28]. In this study, the results of water absorption ability of hydrogels were shown in Fig.4a. The swelling ratio of hydrogels was adjusted by changing the DS of CS-sNHS. All hydrogels reached swelling equilibrium within 27h. Interestingly, higher swelling ratio was showed in hydrogels S-2 and S-3 and lower swelling ratio presented at hydrogel S-4 both relative to hydrogel S-1. The swelling ratio of hydrogel S-2 was a

IP T

little higher than that of hydrogel S-3 before 27h, finally, the swelling ratio of hydrogels

S-2 and S-3 were similar and reached to 37%, while swelling ratio of hydrogels S-4 and

SC R

S-1 were 11% and 17%, respectively.

In the current study, different water absorption capacity mainly contributed to hydrophilic groups in hydrogels, DS of CS-sNHS and chemical crosslinking degree of

U

hydrogels. There are many hydrophilic groups in Col I and CS-sNHS, such as hydroxyl,

N

carboxyl, and amino groups, which contribute to hydration ability and are closely

A

related to water absorption ability of hydrogels. The higher swelling rate of hydrogels S-2 and S-3 was mainly attributed to more hydrophilic groups in CS-sNHS comparing

M

with hydrogel S-1. The lower swelling ratio of hydrogel S-4 was mainly caused by two

ED

reasons: one was the increase of DS of CS-sNHS leading to higher crosslinking degree in hydrogel; and the other one was that there were more hydrophilic groups consumed due to chemical crosslinking in this hydrogel, which led to the reduction of water

PT

absorption ability.

3.3.4. In vitro degradation

CC E

Biodegradability of scaffolds determines their application field. Ideally, hydrogel

scaffold should degrade at a rate in concert with the matrix construction by the seeded cells to form a neo-tissue [29]. In this study, weight loss percentage of hydrogels after

A

being immersed into collagenase I solution at 37℃ was showed in Fig.4b. It was clearly seen that weight loss rate of collagen-based hydrogels depended strongly on their crosslinking manners and crosslinking degree, which was closely associated with the DS of CS-sNHS. Hydrogel S-1 was physically crosslinked by self-assemble, and it degraded with the fastest weight loss rate. Col I/CS-sNHS hydrogels were partially physically crosslinked and partially chemically crosslinked, and the weight loss rate 14

gradually reduced with the increase of DS of CS-sNHS. Their weight loss rate had the same tendency with their water absorption ability. The higher DS of CS-sNHS contributed to higher chemical crosslinking degree in hydrogels, which lead to stable hydrogel structure and low degradation rate. It demonstrated that the degradation rate of scaffolds could be adjusted by crosslinking manners and crosslinking degree.

IP T

3.4. Chondrocytes encapsulation In order to estimate the effect of hydrogels on cell behaviors, chondrocytes were

U

SC R

encapsulated into hydrogels. Live/dead staining, CCK-8 test and histological staining

M

A

N

Fig.4. (a) The swelling ratio curves of hydrogels measured in PBS buffer. (b) The weight loss curves of hydrogels measured in TESCA buffer containing 50µg/mL collagenase I at 37℃. Data are expressed as means ±SD (n = 3).

ED

were performed.

A

CC E

PT

3.4.1. Cell survival and proliferation

15

In order to evaluate the feasibility of hydrogels as biomimetic scaffolds for cell survival, the hydrogels were stained by a live/dead cell staining kit. As shown in Fig.5a, most of cells were alive during culture period, except exactly a few dead cells in each group at 1day. Chondrocytes encapsulated in four kinds of hydrogels proliferated well

A

CC E

PT

ED

M

A

N

U

SC R

IP T

with the extension of culture duration. To quantitatively analyze the proliferation of

16

chondrocytes in these hydrogels, CCK-8 test was performed, and the results demonstrated that chondrocytes in hydrogels proliferated well with the prolongation of culture time as shown in Fig.5b. These results indicated that the hydrogels were suitable for cell survival and proliferation.

A

CC E

PT

ED

M

A

N

U

SC R

IP T

3.4.2. Histological evaluation

Fig.5. (a) Live/dead staining of chondrocytes encapsulated in hydrogel S-1 (a-c), S-2 (d-f), S-3 (g-i) and S-4 (j-l) after being cultured for 1 day (a, d, g, j), 7 days (b, e, h, k) and 14 days (c, f, i, l). The live cells are stained green by FDA, while the dead cells are stained red by PI. The scale bar is 100µm. (b) CCK-8 test for the proliferation of chondrocytes encapsulated in hydrogels. Data are expressed as 17 means ±SD (n = 3).

Hematoxylin&eosin staining, toluidin blue staining and safranin O staining were performed to detect cells morphology and ECM secretion after cell/hydrogel constructs were cultured for 2weeks. The results were shown in Fig.6. Most of cells in hydrogels showed a round morphology and formed small lacunae, as stained by hematoxylin&eosin (Fig.6a, d, g, j). Cell nucleus was stained blue by hematoxylin and cytoplasm was stained light pink by eosin. Secreted ECM such as acid

IP T

mucopolysaccharides around the chondrocytes showed metachromia and appeared

purple with toluidine blue staining (Fig.6b, e, h, k), and basophilic matrix was stained

SC R

red by safranine O staining (Fig.6c, f, i, l), which demonstrated that chondrocytes

encapsulated in these hydrogels maintained their round morphology and secreted ECM. Compared with other hydrogels, there were more ECM secretion around chondrocytes

U

encapsulated into hydrogel S-3 (Fig.6g, h, and i). It seemed that hydrogel S-3 provided

N

a more suitable microenvironment for chondrocytes to maintain morphology and

A

secrete ECM.

Cells adhere to the scaffolds when they are encapsulated in hydrogels [30], and

M

then they proliferate and generate signals that regulate cell behavior [31]. Col I can

ED

provide appropriate microenvironment for cells to adhere and maintain their phenotype owing to their bioactive binding sites [32, 33]. Moreover, it has been proved that Col I can induce rabbit bone marrow derived MSCs to differentiate into chondrocytes both in

PT

vitro and in vivo [34, 35]. CS plays a vital role in maintaining chondrocyte phenotype and promoting ECM secretion [12, 36].

CC E

In this study, the carboxyl groups of CS were partially activated by sNHS. The

hydrogels were fabricated by mixing the neutral solution of Col I and CS-sNHS in moderate conditions without the addition of any catalysts. The storage modulus of

A

hydrogels was

related to the DS of CS-sNHS, which determined the

chemical-crosslinking degree of hydrogels. Chondrocytes maintained round morphology and secreted ECM in these hydrogels. Especially, hydrogel S-3 with moderate storage modulus promoted ECM secretion compared with others. It was noted that the cellular functions of chondrocytes encapsulated in hydrogels were strongly affected by the stiffness of hydrogels [37]. Chondrocytes cultured within the 18

Gtn-HPA hydrogel with medium stiffness produced highest level of sGAG in vitro, and there was markedly more hyaline cartilage formation when the cell/hydrogel constructs with medium stiffness were implanted in the osteochondral defect in a rabbit model [37]. It has been reported that the degree of hydrogel crosslinking also modulated matrix biosynthesis. Lower crosslinking matrix enhanced chondrogenesis with increases in the percentage of cells with chondrocytic morphology, biosynthetic rates of

A

CC E

PT

ED

M

A

N

U

SC R

3.5. In vivo degradation and biocompatibility of hydrogels

IP T

glycosaminoglycan and Col II, and hyaline cartilage tissue formation [26].

Fig.6. Hematoxylin&eosin (a, d, g, j), toluidin blue (b, e, h, k) and safranin O (c, f, i, l) staining of chondrocytes encapsulated in hydrogel S-1 (a-c), S-2 (d-f), S-3 (g-i) and S-4 (j-l) after being cultured for 2weeks. The scale bar is 100µm.

19

In order to assess the in vivo degradation and tissue reactions of injectable hydrogels, 0.5mL of neutral precursor solution for hydrogel S-1 or S-3 were subcutaneously injected into SD rats with a syringe needle. After 1, 2 and 4weeks, the rats were sacrificed, and hydrogels along with their surrounding tissue were harvested. It was obvious that the volume of hydrogels decreased with the prolongation of injection duration, which indicated that hydrogels were subjected to biodegradation

IP T

gradually. There are several enzymes in the body that can degrade Col I and CS. CS can

be degraded by hyaluronidase-PH20, hyaluronidase-1 and hyaluronidase-4 [1], and Col

SC R

I can be degraded by collagenase I and MMPs [38]. Therefore, in vivo degradation of

hydrogels should be mainly attributed to enzymatic cleavage. Samples were stained by hematoxylin&eosin staining to evaluate the inflammatory response at predetermined

U

intervals (Fig.7). Some lymphocytes and macrophages gathered around the hydrogels

N

after injection for one week (Fig.7 a, d, g, j). It indicated that certain inflammatory

A

response took place [39, 40]. It was noteworthy that, with prolongation of injection, inflammatory cells increased at first (Fig.7 b, e, h, k), and then decreased (Fig.7 c, f, i, l).

M

This meant that inflammatory response was gradually maximum when degradation of

ED

hydrogels broke out, and then weakened with the absorption of degradation products. Meanwhile, no obvious tissue necrosis, edema, congestion and hemorrhage were observed during the injection period, and the tissue reactions of hydrogel S-1 and S-3

PT

did not show obvious differences. It demonstrated that these hydrogels exhibited acceptable inflammatory responses, implying that these hydrogels would find some

A

CC E

applications in vivo.

20

IP T SC R U N A M ED

CC E

PT

Fig.7. Hematoxylin&eosin staining of hydrogel S-1 with surrounding tissues (a-f) and hydrogel S-3 with surrounding tissues (g-l) after being subcutaneously injected into the back of SD rats for 1week (a, d, g, j), 2weeks (b, e, h, k) and 4weeks (c, f, i, l), respectively. The scale bar is 500µm for images (a-c, g-i) and 50µm for images (d-f, j-l). Hd, Mc and Lc represent hydrogel, macrophage and lymphocyte, respectively. 4. Conclusions

A

In this study, injectable and self-crosslinkable hydrogels were designed and

fabricated based on Col I and CS-sNHS through physical and chemical crosslinking under mild reaction conditions without the addition of any catalysts. Physical properties of hydrogels, especially mechanical properties, were regulated by adjusting the DS of CS-sNHS. All hydrogels supported chondrocytes survival and ECM secretion. These hydrogels were suitable for cells delivery. Especially, hydrogel S-3 21

provided an appropriate microenvironment for chondrocytes to maintain their morphology and secrete ECM. Moreover, hydrogel S-3 displayed minimal inflammatory responses when injected subcutaneously in SD rats. This study demonstrates that these hydrogels are ideal candidates for application as cells delivery carriers and tissue engineering scaffolds.

IP T

Acknowledgements This work was supported by Science and Technology Support Program of Sichuan

SC R

Province (Grant No. 2016SZ0008), Guangxi Key Research and Development Plan

(Grant No. GuikeAB16450003), the 111 Project (Grant No. B16033), National Natural Science Foundation of China (Grant No. 51403134), and Guangxi Scientific Research

U

and Technological Development Foundation (Grant No. Guikehe 14125008-2-14). We sincerely thank Dr Wen Zou for his careful guidance and valuable suggestions

A

N

in animal experiments.

M

References

[1] S. Lu, C. Gao, X. Xu, X. Bai, H. Duan, N. Gao, C. Feng, Y. Xiong and

M. Liu, Injectable and

Self-Healing Carbohydrate-Based Hydrogel for Cell Encapsulation, ACS Appl Mater Inter, 7 (2015)

ED

13029-13037.

[2] Z. Bao, C. Jiang, Z. Wang, Q. Ji, G. Sun, S. Bi, Y. Liu and

X. Chen, The influence of solvent

formulations on thermosensitive hydroxybutyl chitosan hydrogel as a potential delivery matrix for cell

PT

therapy, Carbohyd Polym, 170 (2017) 80-88. [3] R.R. Smith, E. Marbán and

L. Marbán, Enhancing retention and efficacy of cardiosphere-derived

cells administered after myocardial infarction using a hyaluronan-gelatin hydrogel, Biomatter, 3 (2013)

CC E

e24490.

[4] W.S.S. †, J.S.Y. †, Y. Han Bae and

D.S.L. †, Novel Injectable pH and Temperature Sensitive Block

Copolymer Hydrogel, Biomacromolecules, 6 (2005) 2930-2934. [5] K.N. Kang, D.Y. Kim, S.M. Yoon, J.Y. Lee, B.N. Lee, J.S. Kwon, H.W. Seo, I.W. Lee, H.C. Shin and

Y.M.

Kim, Tissue engineered regeneration of completely transected spinal cord using human mesenchymal

A

stem cells, Biomaterials, 33 (2012) 4828-4835. [6] Y. Ni, Z. Tang, W. Cao, H. Lin, Y. Fan, L. Guo and

X. Zhang, Tough and elastic hydrogel of hyaluronic

acid and chondroitin sulfate as potential cell scaffold materials, Int J Biol Macromol, 74 (2015) 367-375. [7] J. Radhakrishnan, U.M. Krishnan and

S. Sethuraman, Hydrogel based injectable scaffolds for

cardiac tissue regeneration, Biotechnol Adv, 32 (2014) 449-461. [8] B. Yang, Y. Zhang, X. Zhang, L. Tao, S. Li and

Y. Wei, Facilely prepared inexpensive and

biocompatible self-healing hydrogel: a new injectable cell therapy carrier, Polym Chem-Uk, 3 (2012) 22

3235-3238. [9] B. Pei, W. Wei, Y. Fan, X. Wang, F. Watari and

X. Li, Fiber-reinforced scaffolds in soft tissue

engineering, Regen Biomater, 4 (2017) 257-268. [10] J. Radhakrishnan, A. Subramanian, U.M. Krishnan and

S. Sethuraman, Injectable and 3D

Bioprinted Polysaccharide Hydrogels: From Cartilage to Osteochondral Tissue Engineering, Biomacromolecules, 18 (2017) 1-26. [11] A. Sadeghi-Avalshahr, S. Nokhasteh, A.M. Molavi, M. Khorsand-Ghayeni and

M. Mahdavi-Shahri,

Synthesis and characterization of collagen/PLGA biodegradable skin scaffold fibers, Regen Biomater, 4 (2017) 309-314. X. Zhang, Preparation of collagen–

IP T

[12] L. Zhang, K. Li, W. Xiao, L. Zheng, Y. Xiao, H. Fan and

chondroitin sulfate–hyaluronic acid hybrid hydrogel scaffolds and cell compatibility in vitro, Carbohyd Polym, 84 (2011) 118-125.

X. Zhang, Modulation of immunological properties of allogeneic

SC R

[13] T. Yuan, K. Li, L. Guo, H. Fan and

mesenchymal stem cells by collagen scaffolds in cartilage tissue engineering, Journal of biomedical materials research. Part A, 98 (2011) 332-341. [14] C. Chen, J. Xie, L. Deng and

L. Yang, Substrate stiffness together with soluble factors affects

chondrocyte mechanoresponses, ACS Appl Mater Inter, 6 (2014) 16106-16116.

S.J. Bryant, Mechanical loading

U

[15] N.J. Steinmetz, E.A. Aisenbrey, K.K. Westbrook, H.J. Qi and

regulates human MSC differentiation in a multi-layer hydrogel for osteochondral tissue engineering, [16] H. Cao and

N

Acta Biomater, 21 (2015) 142-153.

S.Y. Xu, EDC/NHS-crosslinked type II collagen-chondroitin sulfate scaffold:

A

characterization and in vitro evaluation, J Mater Sci-Mater M 19 (2008) 567-575. and

M

[17] D.A. Wang, S. Varghese, B. Sharma, I. Strehin, S. Fermanian, J. Gorham, D.H. Fairbrother, B. Cascio J.H. Elisseeff, Multifunctional chondroitin sulphate for cartilage tissue-biomaterial integration,

Nat Mater, 6 (2007) 385-392.

X.D. Zhang, Preparation and Characterization of

ED

[18] J. Lu, X.Y. Lin, B. Jiang, X.D. Li, J.Y. Chen and

Collagen by Hydrogel Formation Method, Key Eng Mater, 288-289 (2005) 377-380. [19] L. Yuan, B. Li, J. Yang, Y. Ni, Y. Teng, L. Guo, H. Fan, Y. Fan and

X. Zhang, Effects of composition

PT

and mechanical property of injectable collagen I/ II composite hydrogels on chondrocytes behaviors, Tissue Eng Part A, 22 (2016) 899-906. [20] B. Balakrishnan and

R. Banerjee, Biopolymer-based hydrogels for cartilage tissue engineering,

CC E

Chem Rev, 111 (2011) 4453-4474. [21] M.H. Chang, S. Sant, M. Masaeli, N.N. Kachouie, B. Zamanian, S.H. Lee and

A. Khademhosseini,

Fabrication of three-dimensional porous cell-laden hydrogel for tissue engineering, Biofabrication, 2 (2010) 035003.

[22] E. Hadjipanayi, V. Mudera and

R.A. Brown, Close dependence of fibroblast proliferation on

A

collagen scaffold matrix stiffness, Journal of tissue engineering and regenerative medicine, 3 (2009) 77-84. [23] E. Hadjipanayi, V. Mudera and

R.A. Brown, Guiding cell migration in 3D: a collagen matrix with

graded directional stiffness, Cell motility and the cytoskeleton, 66 (2009) 121-128. [24] P.G. Buxton, M. Bitar, K. Gellynck, M. Parkar, R.A. Brown, A.M. Young, J.C. Knowles and

S.N.

Nazhat, Dense collagen matrix accelerates osteogenic differentiation and rescues the apoptotic response to MMP inhibition ☆, Bone, 43 (2008) 377-385. [25] D. Schop, F.W. Janssen, E. Borgart, J.D.D. Bruijn and

R.V. Dijkhuizen-Radersma, Expansion of 23

mesenchymal stem cells using a microcarrier-based cultivation system: growth and metabolism, J Tissue Eng Regen Med 2(2008) 126–135. [26] S.T. Wei, T.C. Lim, M. Kurisawa and

M. Spector, Modulation of mesenchymal stem cell

chondrogenesis in a tunable hyaluronic acid hydrogel microenvironment, Biomaterials, 33 (2012) 3835-3845. [27] X. Yang, L. Guo, Y. Fan and

X. Zhang, Preparation and characterization of macromolecule

cross-linked collagen hydrogels for chondrocyte delivery, Int J Biol Macromol, 61 (2013) 487-493. [28] H. Park, X. Guo, J.S. Temenoff, Y. Tabata, A.I. Caplan, F.K. Kasper and

A.G. Mikos, Effect of

swelling ratio of injectable hydrogel composites on chondrogenic differentiation of encapsulated [29] Q. Feng, M. Zhu, K. Wei and

IP T

rabbit marrow mesenchymal stem cells in vitro, Biomacromolecules, 10 (2009) 541-546.

L. Bian, Cell-mediated degradation regulates human mesenchymal

stem cell chondrogenesis and hypertrophy in MMP-sensitive hyaluronic acid hydrogels, Plos One, 9 [30] N.R. Paul, G. Jacquemet and

SC R

(2014) e99587.

P.T. Caswell, Endocytic Trafficking of Integrins in Cell Migration,

Current Biology Cb, 25 (2015) 1092-1105. [31] M. Chen, Y. Chen and

G. Diao, Adsorption Kinetics and Thermodynamics of Methylene Blue

onto p-tert-Butyl-calix[4,6,8]arene-Bonded Silica Gel, J Chem Eng Data, 55 (2010) 5109-5116.

K.M. Yamada, Local 3D matrix

microenvironment

spatiotemporal

regulates

cell

migration

U

[32] A.D. Doyle, N. Carvajal, A. Jin, K. Matsumoto and through

dynamics

of

[33] X. Li, Y. Teng, J. Liu, H. Lin, Y. Fan and

N

contractility-dependent adhesions, Nat Commun, 6 (2015) 8720.

X. Zhang, Chondrogenic differentiation of BMSCs

A

encapsulated in chondroinductive polysaccharide/collagen hybrid hydrogels, J Mater Chem B, 5 (2017) [34] L. Zhang, T. Yuan, L. Guo and

M

5109-5119.

X. Zhang, An in vitro study of collagen hydrogel to induce the

chondrogenic differentiation of mesenchymal stem cells, Journal of biomedical materials research.

ED

Part A, 100 (2012) 2717-2725.

[35] L. Zheng, J. Sun, X. Chen, G. Wang, B. Jiang, H. Fan and

X. Zhang, In vivo cartilage engineering

with collagen hydrogel and allogenous chondrocytes after diffusion chamber implantation in

PT

immunocompetent host, Tissue Eng Part A, 15 (2009) 2145-2153. [36] M. Zhu, Q. Feng, Y. Sun, G. Li and

L. Bian, Effect of cartilaginous matrix components on the

chondrogenesis and hypertrophy of mesenchymal stem cells in hyaluronic acid hydrogels, Journal of

CC E

biomedical materials research. Part B, Applied biomaterials, 105 (2017) 2292-2300. [37] L.S. Wang, C. Du, S.T. Wei, A.C.A. Wan, S.J. Gao and

M. Kurisawa, Modulation of chondrocyte

functions and stiffness-dependent cartilage repair using an injectable enzymatically crosslinked hydrogelwith tunable mechanical properties, Biomaterials, 35 (2014) 2207-2217. [38] P. Panwar, G.S. Butler, A. Jamroz, P. Azizi, C.M. Overall and

D. Bromme, Aging-associated

A

modifications of collagen affect its degradation by matrix metalloproteinases, Matrix biology : journal of the International Society for Matrix Biology, 10 (2017) 1016. [39] K. Ren, C. He, C. Xiao, G. Li and

X. Chen, Injectable glycopolypeptide hydrogels as biomimetic

scaffolds for cartilage tissue engineering, Biomaterials, 51 (2015) 238-249. [40] Y. Cheng, C. He, J. Ding, C. Xiao, X. Zhuang and

X. Chen, Thermosensitive hydrogels based on

polypeptides for localized and sustained delivery of anticancer drugs, Biomaterials, 34 (2013) 10338-10347.

24