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Directed osteogenic differentiation of mesenchymal stem cell in three-dimensional biodegradable methylcellulose-based scaffolds
Accepted Manuscript Title: Directed osteogenic differentiation of mesenchymal stem cell in three-dimensional biodegradable methylcellulose-based scaffolds Author: He Shen Yufei Ma Yu Luo Xiaoyun Liu Zhijun Zhang Jianwu Dai PII: DOI: Reference:
S0927-7765(15)30096-5 http://dx.doi.org/doi:10.1016/j.colsurfb.2015.07.062 COLSUB 7266
To appear in:
Colloids and Surfaces B: Biointerfaces
Received date: Revised date: Accepted date:
27-3-2015 18-6-2015 22-7-2015
Please cite this article as: H. Shen, Y. Ma, Y. Luo, X. Liu, Z. Zhang, J. Dai, Directed osteogenic differentiation of mesenchymal stem cell in three-dimensional biodegradable methylcellulose-based scaffolds, Colloids and Surfaces B: Biointerfaces (2015), http://dx.doi.org/10.1016/j.colsurfb.2015.07.062 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.
Graphical Abstract
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Preparation of MC scaffolds with three-dimensional porous structure for regulating the fate of MSCs. Direction osteogenic differentiation of MSCs via MC scaffolds. Stimulatory effect of mechanical properties and 3D structure of the MC scaffolds on inducing the stem cell differentiation without osteogenic inducer supplement.
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Highlights
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Directed osteogenic differentiation of mesenchymal stem cell in three-dimensional
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biodegradable methylcellulose-based scaffolds
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He Shen, a Yufei Ma, a Yu Luo, a, c Xiaoyun Liu, a Zhijun Zhang,*a and Jianwu Dai*a, b
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Nano-tech and Nano-bionics, Chinese Academy of Sciences (CAS), 398 Ruoshui Road, Suzhou,
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215123, China
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b
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College of Preventive Medicine, Third Military Medical University, 30, Gaotanyan Road,
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Chongqing 400038, China
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Institute of Combined Injury, State Key Laboratory of Trauma, Burns and Combined Injury,
College of Chemistry, Chemical Engineering and Biotechnology, Donghua University, Shanghai,
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Key Laboratory of Nano-Bio Interface, Division of Nanobiomedicine, Suzhou Institute of
201620, China
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* Corresponding authors. Email: [email protected], Tel.: +86-10-82614426; Email:
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[email protected], Tel.: +86-512-62872556.
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Abstract:
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Development of three-dimensional (3D) biodegradable scaffolds that can accelerate mesenchymal
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stem cell (MSC) osteogenic differentiation is a decisive prerequisite for treatment of damaged
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skeletal tissue. We report herein the preparation of methylcellulose-based (MC) scaffolds using
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carbonyldiimidazole as cross-linking agent to produce substrates with specific cross-linking
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density and porous structure, as well as their applications for directing hMSC toward osteoblasts.
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The mechanical properties of the scaffolds were controlled by cross-linking density. Human
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MSCs (hMSCs) seeded on the MC scaffolds have penetrated into the pores, and showed high
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viability (> 80%) as revealed by WST assay and Live/Dead assay. Moreover, the results of
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differentiation experiments indicated that hMSCs cultured on MC substrates displayed high level
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of osteogenic differentiation marker expression, alkaline phosphatase activity and osteocalcin
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secretion, suggesting that the MC scaffolds can direct hMSC differentiation towards osteoblasts
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without inducer treatment and cross-linking density of MC scaffolds have stimulatory effect on
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inducing differentiation. The 3D MC scaffolds could be applicable as promising scaffolds for
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bone tissue repair.
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Keywords: three-dimensional scaffold; methylcellulose; cross-linking density; mesenchymal
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stem cell; osteogenic differentiation.
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1.
Introduction
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The regenerative medical applications of biomaterials pursue the reconstruction of damaged
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tissues by controlling the fate of the cells cultured on implanted scaffolds [1]. The biomaterials
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with rationally designed three-dimensional (3D) architectures are ideal tissue engineering
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scaffolds due to their close emulation of the topographies and spatial structures of native
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extracellular matrices (ECMs) that facilitate stem cell proliferation and differentiation [2]. Cells
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cultured in vitro within these artificial 3D structures are subjected to both biochemical and
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biophysical stimuli, and then coaxed to specific cell lineages. Therefore, 3D scaffolds can control
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the stem cell fate more precisely than 2D substrates [3, 4]. To date, numerous 3D scaffolds have
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been developed, such as 3D architectures based on chitosan [5], polylactic acid [6], poly
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(ethylene glycol) (PEG) [7], graphene [8-11], and collagen [12, 13], as well as hydrogels based
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on matrigel [14], hyaluronic acid [15], and alginate [16, 17]. Scaffolds with porous internal
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architecture are considered as ideal biomaterials for tissue engineering as they permit efficient
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diffusion of nutrients and metabolism products, and can be tuned to yield a favorable
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microenvironment that influences cell proliferation and differentiation [18, 19]. Hence, these 3D
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porous scaffolds have been extended to be in combination with mesenchymal stem cells (MSCs)
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for the utilization in cartilage, skeletal tissue, and smooth muscle tissue engineering [20-23].
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With the goal towards tissue regeneration, development of 3D scaffolds that mimic ECM, as
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well as provide an environment and necessary signals for controlling stem cell fate, is highly
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desired [24]. Two of the important scaffold parameters are cross-linking density and pore size,
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which not only affect, to a large extent, the mechanical properties of the scaffolds, but also
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impose impacts on stem cells metabolism [25, 26]. It has been demonstrated that MSCs are able
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to recognize mechanical signals. Complex sensory machinery can involve a group of cell surface
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receptors and intracellular proteins that mediate mechanical signals from substrate to regulate a
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variety of gene expressions [27]. Chatterjee and co-workers have reported that the stiffness of
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PEG hydrogels regulates osteogenic differentiation of the encapsulated stem cells [28]. The
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synthetic scaffolds used for skeletal tissue engineering should ideally act as artificial ECM for
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regenerative applications, and provide suitable condition for directing of MSCs differentiation
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into specific cell lineage. However, these investigations were carried out on 2D tissue-culture
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platforms. Whether MSC cultured on 3D substrates could be induced differentiation by 4 Page 4 of 24
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mechanical property is still unknown. The regulation of MSC differentiation by the mechanical
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properties of 3D scaffolds needs further investigation. Methylcellulose (MC) is a chemically modified polysaccharide with a partial substitution of
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hydroxyl groups with methoxy moieties. Its biocompatibility and capability of gelation make it a
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food additive for thickening [29]. In order to control the gelling process, MC has been combined
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with different components, such as agarose [30] and hyaluronan (HA) [31], leading to different
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characteristics of the MC-based scaffolds. Recently, MC-based biomaterials have been explored
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extensively for tissue regenerative applications, such as supporting substrates and regenerative
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factor carriers [32]. Tate et al. have investigated MC as a drug delivery/tissue engineering
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scaffold for treating traumatic brain injury [30]. Taking the advantages of MC, Hsieh et al. further
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prepared MC-based scaffolds combined with HA for neural stem/progenitor cell culture [24].
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Although some progress has been made on applications of MC-based scaffolds in tissue
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regeneration, the effects of their physical properties on stem cell differentiation were actually
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ignored.
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In the investigation, a series of porous nanostructured MC-based scaffolds with different
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cross-linking density and mechanical properties were developed for bone tissue engineering. We
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synthesized various MC-based scaffolds with low degree of cross-linking (LCL), medium degree
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of cross-linking (MCL), and high enough degree of cross-linking (HCL). In order to address the
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role of pores in the MC-based substrates, we used a scaffold with a low porosity level (LP) as a
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control. The morphology, degradation, wettability and mechanical properties of the MC scaffolds
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were studied. In addition, the toxicity and hemocompatibility of these biomaterials were
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examined. Finally, osteoblast responses towards these 3D scaffolds were investigated. Directing
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osteogenic differentiation of the hMSC cultured on the MC-based substrates with 3D structures
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provides useful insights into designs of new scaffolds for bone tissue engineering applications.
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2.
Materials and methods
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2.1 Preparation of 3D porous MC-based scaffolds
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To prepare MC-based scaffolds, MC (500 mg) was first mixed with gelatin (50 mg) and agarose
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(50 mg) in 25 mL dimethyl sulfoxide (DMSO) to form a homogeneous solution. 5 g ground-NaCl 5 Page 5 of 24
was then added into the above solution, followed by vigorous stirring for 4 h. Cross-linking agent
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(carbonyldiimidazole, CDI) was added into the above homogenized mixed solution at a varying
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concentration, and then the resulted solution was vigorously agitated for 15 min. MC scaffolds
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with low degree of cross-linking (LCL), medium degree of cross-linking (MCL), and high degree
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of cross-linking (HCL) were prepared by adding 250 mg, 500 mg, and 750 mg CDI into the
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solution, respectively. To synthesize the MC scaffold with a low porosity level (LP), MC (500
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mg) was mixed with 50 mg gelatin and 50 mg agarose in 25 mL DMSO, and then 500 mg CDI
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was added. These mixture solutions were incubated in ice bath for scaffold formation, and then
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washed with distilled water 5 times to remove NaCl, DMSO and unreacted reagents. Finally,
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these samples were frozen and lyophilized before use.
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2.2 Morphology of the MC-based scaffolds
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Microstructure of the scaffolds was examined by scanning electron microscope (SEM) with an
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accelerating voltage of 10 kV. All MC-based samples were sputter coated with a 10 nm gold film
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before SEM observation.
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2.3 Swellability of the MC-based scaffolds
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Equilibrium swelling ratio (QS), a measure of swellability of the MC-based scaffolds, was
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estimated by immersing dried the MC samples to swell in PBS at 37 oC for 7 days. QS was
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defined according to the following equation:
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where Ws and Wd are the weight of the swollen sample and the dried sample, respectively [5].
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In addition, swelling measurement was performed to estimate the cross-linking density of
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MC-based scaffolds, as the degree of swelling is known to be dependent upon the crosslinking
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density of the scaffolds. The cross-linking densities of the MC-based scaffolds were calculated
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from the Flory-Rehner equation as follows:
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where Vp, Vo, Dp, Do, and x are the volume fraction of polymers in the swollen mass, molar
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volume of solvent, density of polymers, density of solvent, solvent-polymer interaction term,
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respectively [33]. Mean and standard deviations for the triplicate samples for each scaffold were
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reported.
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2.4 Mechanical testing
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Prior to mechanical testing by a material testing machine (H5K-S, Hounsfield, UK), all frozen
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dried samples were cut into small strips, and then the width, length, and thickness were measured
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with a micrometer, and three strips of each scaffold were chosen for the mechanical test. Stress
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and strain were calculated according to the following equations:
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where σ, ε, P, w, d, l and l0 stand for stress, strain, load, mat width, mat length, extension length,
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and gauge length, respectively. Young’s modulus can be calculated from stress–strain curves [34,
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35].
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2.5 Degradation behavior of the MC-based scaffolds
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The degradation ratio of the MC-based scaffolds was measured after 120 d incubation. LCL,
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MCL, HCL and LP samples, in triplicate, were incubated in physiological saline solution
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containing 0.01% trypsin at 37 oC, respectively. The dried MC samples were weighed after being
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frozen and lyophilized. The degradation ratio (Rd) of MC-based scaffolds was calculated
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according to the following equation:
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where W0 and W120 represent the initial dried weight and the dried weight of the sample at day
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120, respectively. 7 Page 7 of 24
2.6 Cell adhesion, growth, and osteogenesis analysis
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Small pieces of human umbilical cord were washed with PBS and transferred into tissue culture
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plates. The culture medium was changed every 2 days until 80% confluence was reached.
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DMEM/F12 (50/50) supplemented with 10% FBS, 100 U/mL penicillin, and 100 μg/mL
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streptomycin solution were used throughout the routine culture procedures. Human MSCs
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(passage 5-8) were seeded at a density of 2×104, 1×104 and 1×104 cells per well (24-well plates,
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Corning, USA) for the cell adhesion, growth, and osteogenesis assay, respectively.
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Typically, for cell attachment assay, cell suspensions were seeded onto the different scaffolds,
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and cultured for 1, 3, 6 and 12 h, respectively. After incubation, MC-based substrates were
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washed to remove the non-adherent cells. The cell viability was determined via WST assay
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according to the manufacturer’s instruction. 5 days after incubation, Calcein-AM (CAM) and
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propidium iodide (PI) staining were used to detect the live and dead cells cultured on MC-based
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scaffolds and TCP according to the instruction. To evaluate the extent of proliferation, hMSCs
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were cultured onto the scaffolds. After 7, 14, 24 and 33 d treatment, cell density was evaluated by
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total double-stranded DNA content using the PicoGreen DNA quantification kit.
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For cell differentiation analysis, cell suspensions were seeded onto different scaffolds, and
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cultured for 7, 14, 24 and 33 d, respectively. The osteogenic differentiation trends were
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determined by alkaline phosphatase (ALP) activity assay and calcium phosphate secretion test,
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and reverse transcription polymerase chain reaction (RT-PCR). All measurements were carried
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out in three independent experiments.
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2.7 ALP activity
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ALP activity was assessed by the substrate p-nitrophenyl phosphate in an end-point assay. The
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generation of p-nitrophenol production is proportional to the amount of ALP in the solution.
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Standard calibration curve was prepared by p-nitrophenol dilution. The absorbance was read at
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405 nm and then ALP content was calculated.
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2.8 Osteocalcin secretion
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An intact human osteocalcin EIA kit BT-460 (Biomedical Technologies Inc., USA) was used for
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measurement of osteocalcin secretion. On day 13 and 20, the medium was replaced with fresh
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DMEM/F12 (50/50) medium (without FBS). After 24 h of culture, the medium was harvested and
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the osteocalcin content was analyzed using the kit according to our previous work [35].
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2.9 Confocal imaging
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After incubation for 7 d, the hMSCs cultured on MC-based substrates were fixed in 4% formalin
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and the cell nucleus was stained by DAPI. Confocal fluorescence microscopy images were
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captured at different altitude to investigate the cell distribution in the 3D scaffolds.
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2.10 Histological staining
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After 20 d culturing, hMSCs on LCL, MCL, HCL, and LP substrates were stained by different
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agents to study the expression of ALP and mineralization of bone nodules. ALP activity was
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detected histochemically by incubation with a mixture of naphthol AS-MX phosphate and fast
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violet B salt solution [36]. For the Alizarin red staining, scaffolds were incubated with Alizarin
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red [37].
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2.11 RNA isolation and RT-PCR assay
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To investigate the cell phenotype, mRNA was analyzed using RT-PCR. After 5, 10, 15 and 20 d
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co-culture, the specimens were washed by PBS, and then suspended in TRIzol Reagent (Life
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Technologies Co.) to extract total RNA. Then the total RNA was reverse transcribed by using
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transcriptase reaction mix (SuperScript III First-Strand Synthesis System, Life Technologies) for
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cDNA generation.
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Quantitative PCR analysis was performed in triplicate using power SYBR green RT-PCR kit
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(Life Technologies) on a CFX96 system (Bio-Rad, USA). Glyceraldehyde-3-phosphate
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dehydrogenase (GAPDH) was used as an endogenous housekeeping gene. The data were
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calculated using the ΔΔCt method as previously reported [12]. The cells on TCP substrate were
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set as control group, and their relative levels of marker genes expression were artificially set as
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one-fold. The genes and primer sequences are listed in Table S1.
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2.12 Statistical analysis
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The data were expressed as mean ± SD (n = 3). Statistical analysis was performed and p < 0.05
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was considered as statistically significant.
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3.
Results 9 Page 9 of 24
3.1 Characterization of the MC-based scaffolds
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SEM images indicate that the MC scaffolds have three-dimensional, porous, and foamy structures
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(Figure 1A), which may facilitate the formation of bridges across the lesion sites of damaged
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tissue [2]. The freeze-dried scaffolds can be cut into pieces with any desired shape for clinical
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applications. As shown in Figure 1A, numerous pores with diameters ranging from 100 μm to
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600 μm were observed on LCL, MCL and HCL scaffolds, allowing nutrient diffusion through
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these pores. NaCl particles play a vital role in the pore generation. Low porosity level was
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observed on LP scaffold, which was prepared without addition of NaCl particles.
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To confirm the cross-linking densities of LCL, MCL, and HCL scaffolds, the swelling ratios
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of these MC scaffolds were studied. The swelling ratios of the MC-based scaffolds were
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significantly infulenced by CDI concentration (Figure 1B) The swelling rates of LCL, MCL, and
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HCL scaffolds were 108.20, 35.28, and 21.52, respectively. The cross-linking density of
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MC-based scaffolds is shown in Figure 1C. The cross-linking density of the LCL, MCL, and
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HCL scaffolds was 0.003, 0.011, and 0.025 ×10 -5mol cm3, respectively. The cross-linking density
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of the MC scaffolds increased with increasing amount of CDI.
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The mechanical properties of scaffolds can be readily controlled by tuning concentrations of
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polymers or cross-linking molecules. The experiment demonstrated that tensile, stress, and
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Young’s modulus of the MC-based scaffolds could be regulated by changing the CDI
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concentration (Figure 2). With increasing amount of the cross-linker, the Young’s modulus of the
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MC-based scaffolds was significantly increased. In addition, water diffused in the MC-based
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scaffolds very quickly. The water diffusion time of LCL, MCL and HCL scaffolds was
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determined to be 10-20 s (Figure S2A). A favorable and biodegradable scaffold provides
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structural support for initial cell growth and then would gradually degrade after new tissue
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formation [38, 39]. Thus we investigated the degradation behavior of the MC scaffolds. As shown
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in Figure S2B, LCL scaffold displayed less than 30 % of the original mass, while HCL and LP
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scaffolds lost nearly 21 % and 8 % weight, respectively, after 120 days of incubation in vitro. The
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degradation of the MC-based scaffolds was also dependent on their cross-linking ratios. Higher
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cross-linking density and lower porosity improved the stability of the HCL and LP scaffolds.
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3.2 In vitro toxicity of MC-based scaffolds
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As biocompatibility is an essential issue for tissue engineering scaffolds, we studied the in vitro
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toxicity of MC-based scaffolds by WST assay after hMSCs cultured with LCL, MCL, HCL, and
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LP for 3 and 5 d, respectively. The viability of MSCs was used to assess the cytotoxicity of
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scaffolds on which MSCs were cultured. For comparison purpose, traditional TCP substrate was
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employed as a control group (100 % cell viability). As shown in Figure S3, all 3D scaffolds did
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not show any significantly deleterious effects on MSC viability, indicating their good
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biocompatibility.
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The biocompatibility of the scaffolds was also evaluated by hemolysis test. The release of
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hemoglobin was used to quantify the membrane-damaging properties of the LCL, MCL, HCL,
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and LP scaffold. Distilled water and PBS buffer treated erythrocytes were set as 100% and 0%
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values, respectively. As Figure S4 shown, LCL, MCL, HCL and LP scaffold caused 3.0 %, 1.8 %,
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1.8 % and 1.5 % hemolysis, respectively. All these four types of MC scaffolds exhibited neither
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detectable disturbance of the RBC membranes, nor any hemolytic phenomenon, suggesting that
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the MC scaffolds have excellent hemocompatibility, an important premise for their in vivo
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regenerative medical applications.
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3.3 Human MSC growth on 3D MC-based scaffolds
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In this experiment, the cell distribution inside the 3D MC scaffolds was studied by taking
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confocal fluorescence microscopy images at different altitude from top to bottom of these
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scaffolds across the central part (Figure 3). A spatial arrangement of the hMSCs was observed on
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these 3D scaffolds. The MSCs cultured for 7 days penetrated into the pores of the scaffolds and
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formed aggregates, which may reestablish cell-cell contacts (Figure 3A-C) [11]. However, only a
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few hMSCs were detected inside the LP substrate owning to its low porosity, suggesting the
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limitation of cell migration and growth in this case (Figure S5). Additionaly, CAM/PI based
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Live/Dead cell viability assay was used to assess live and dead cells cultured on the scaffolds. As
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shown in Figure S6, most of the cells were live (>95%) on the MCL scaffold and TCP after 5
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days culture.
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In order to understand how hMSCs proliferated on the MC-based scaffolds, DNA was
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extracted from the LCL, MCL, HCL, and LP substrates, as the amount of genes is proportional to
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the cell number. At the initial stage, only a few cells have adhered onto LCL, MCL, HCL, and LP 11 Page 11 of 24
substrate (Figure S7), so the hMSCs cultured on the 3D scaffolds showed slow cell proliferation
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rate, compared to traditional 2D scaffold, on day 7 and 14 (Figure S8). However, with increasing
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incubation time, proliferation of hMSCs on TCP was inhibited due to limited culture space. The
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MSCs cultured on LCL, MCL, and LCL scaffolds showed a lag phase of growth, giving rise to a
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slow rate of proliferation. Inhibition of hMSC growth was also observed on the LP scaffold, due
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to its low porosity level and restricted space. It turns out that the LP scaffold is not a good
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scaffold material for bone tissue reparation, because the density of hMSCs on the LP substrate is
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too low for practical applications.
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3.4 Osteogenic differentiation on 3D porosity scaffolds
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The cell phenotype on different MC-based scaffolds was further investigated to figure out how
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MSCs differentiated on these 3D substrates. Alkaline phosphatase (ALP) activity is one of the
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most commonly used markers for osteogenesis differentiation, as ALP enzyme is involved in the
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early outset of mineralization and is widely found in newly formed bone tissue. The osteogenesis
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differentiation was first examined by ALP staining. The stained ALP positive areas appeared dark
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red color, as shown in Figure 4A. The expression of ALP was observed from the cells cultured on
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LCL, MCL, and HCL scaffolds, whereas it was not present in the control group (TCP).
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The calcium deposition and collagen secretion of cells on the MC-based scaffolds were also
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investigated, as they play important roles in the mineralization during bone reparation [40].
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Alizarin red S staining was utilized for the qualitative assessment of mineralized matrix
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formation. As shown in Figure 4A, the calcium and mineral depositions could be clearly
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observed as red regions on LCL, MCL, and HCL scaffolds incubated with hMSCs for 20 d. The
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results suggest that the LCL, MCL, and HCL scaffolds could induce hMSCs to differentiate into
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osteoblasts after 20 d of incubation, even without the treatment of osteogenic inducer such as
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dexamethasone or bone morphogenetic protein-2.
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The quantitative analyses of ALP activity and extracellular osteocalcin production were
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further carried out to study the cell differentiation behavior. The ALP activity of hMSCs cultured
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on LCL, MCL, and HCL scaffold as well as TCP (control group) was studied after 7, 14, 21 and
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33 days of incubation, respectively (Figure 4B). For the first 7 days, ALP activity of hMSCs on
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both TCP and MC-based substrates presented very low activity levels. During the next 7 days, 12 Page 12 of 24
there was a significant increase in the ALP activity for 3D substrate samples, compared with the
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control group (TCP). The ALP activity of hMSCs on LCL, MCL, and HCL scaffolds was much
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higher than that on TCP substrate at 14, 21 and 33 d, indicating the enhancement of hMSC
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differentiation towards osteoblasts lineage. However, the ALP activity of the cells on LCL, MCL,
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and HCL scaffolds decreased after 21 and 33 d culture, compared to the 14 d culture. We assumed
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that the hMSCs cultured on LCL, MCL and HCL scaffolds were induced to osteogenic
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differentiate at first 14 days. Thus, most of the cells on MC-based scaffolds tended to be
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osteoblasts at 21 and 33 d, which leads to low levels of ALP activity of cells.
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The extracellular osteocalcin production contents of the hMSCs cultured onto TCP, LCL,
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MCL, and HCL scaffolds were also examined after14 and 21 d co-culturing, respectively. In the
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absence of osteogenic inducer (dexamethasone) in the culture medium, hMSCs grown on TCP
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substrate secreted little osteocalcin, because most of the cells on TCP maintained their
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self-renewal property. As expected, the osteocalcin production contents of hMSCs on LCL, MCL,
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and HCL scaffolds were remarkably higher than that on TCP at 14 d and 21 d (Figure 4C). These
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data further indicated the progression of osteogenic differentiation on the 3D substrates. Similar
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to ALP activity, the cells cultured on the MC-based scaffolds at 14 d exhibited higher osteocalcin
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contents than those on the same sample group at 21 d, as most of hMSCs on the MC-based
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scaffolds differentiated into osteoblasts at 14 d.
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To gain a profound understanding of the effects of 3D MC-based scaffolds on directing
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hMSC fate, targeted osteoblastic gene markers were quantitatively assessed by RT-PCR after 5,
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10, 15 and 20 d treatment, respectively. Firstly, relative levels of ALP gene expression were all
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up-regulated in the hMSCs cultured on the LCL, MCL, and HCL scaffolds, compared with those
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seeded on TCP substrate after 10, 15 and 20 d co-culturing (Figure 5A). Osteocalcin (Ocn), a late
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stage osteogenesis and mineralization marker during bone formation, was also examined. With
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the treatment of the MC-based scaffolds, the Ocn expression was enhanced in comparison with
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treatment of TCP, suggesting the induced osteogenic differentiation of hMSCs via MC-based
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platforms. Additionally, the osteoblastic marker genes expression levels of hMSCs on MC
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scaffolds is higher at 15 d than those at 5, 10 and 20 d (Figure 5B). Hence, the experimental
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results indicated that hMSCs cultured on LCL, MCL, and HCL scaffolds were more osteogenic
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and produced more minerals than any other time periods during the first two weeks. More 13 Page 13 of 24
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importantly, the expression levels of ALP and Ocn gene of hMSCs cultured on the scaffolds were
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increased by raising the cross-linking density of the MC-based substrates.
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4.
Discussion
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The presence of pores in scaffolds is necessary for bone tissue reparation because they allow
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MSC proliferation as well as vascularization. Additionally, it has been reported that porosity
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could stimulate osteogenesis in vitro [27]. The pore size of the substrate is recommended to be
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200 - 900 μm for accelerating stem cell growth and differentiation [41]. Although after
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lyophilization the pores created through salt leaching will change their dimensions completely,
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LP scaffold without addition of NaCl particles during prepartions displayed low porous level,
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which might restrict cell migration and growth, as well as nutrition exchange. The mixing of
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gelatin and agarose with MC usually benefits 3D scaffolds for cell growth. Martin et al. reported
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that MC and agarose hydrogel blends and their applications of nerve regeneration which
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combined of agarose’s ability to support nerve growth and the gelling characteristics [31].
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Additionally, gelatin, a denatured collagen, is a biodegradable polymer with high biocompatibility,
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which has been considered as a suitable substrate for cells [42]. It's also demonstrated that gelatin
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scaffolds gelatin-based scaffolds are favorable for cell adhesion and growth [43].
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In addition, surface wettability of biomaterials relates to their surface properties for cell
19
adhesion and proliferation [44]. The good wettability of these 3D substrates is also attributed to
20
their porous structure and abundant hydrophilic groups. Besides, water diffusion time of LP
21
scaffold is much longer than that of LCL, MCL and HCL scaffolds, probably due to that the low
22
porosity level of LP scaffold inhibited the water diffusion and wetting. Furthermore, a trend of
23
decreasing swelling ratio was observed with increasing the concentration of the crosslinking
24
agent. The swelling ratio of the MC scaffold increased, while the cross-linking density of the
25
polymer network decreased [33, 45].
26
Moreover, the cross-linking density and porosity of MC-based scaffolds also impose
27
significant impacts on mechanical property, which is another important parameter of scaffolds for
28
bone tissue engineering [46]. It has been elegantly demonstrated that the substrate elasticity
29
induces MSC differentiation into specific lineages [47]. The 3D space of the MC scaffolds is 14 Page 14 of 24
available for hMSCs proliferation during a long period of time. The porous MC-based substrates
2
allowed cell migration and efficient exchange of nutrients and wastes, as well as provided
3
appropriate mechanical condition for MSCs growth. After long time incubation with these
4
MC-based substrates, the density of the MSCs became higher than that of the control group
5
(TCP). More importantly, the MC-based substrates with different mechanical properties
6
stimulated the related gene expression and osteogenic differentiation without induction.
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The above osteogenic marker analysis revealed that the porous 3D MC-based substrates
8
promote differentiation of MSCs without osteogenic inducer treatment. Interestingly, we found
9
that the ALP activity, osteocalcin secretion, and marker gene expression of the cells incubated
10
with HCL substrate were higher than those of the hMSCs treated with LCL. The origin of
11
directed hMSCs differentiation on the 3D porous platforms could be mainly attributed to two
12
reasons [48]: one is that the porous structure of the 3D MC scaffolds provides ECM biomimetic
13
microenvironment for hMSCs growth and differentiation; the other is that the mechanical
14
properties of LCL, MCL, and HCL scaffolds induced differentiation of the hMSCs.
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In this research, 2D tissue culture platform exhibited a very low level of osteogenic
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differentiation without dexamethasone treatment, because the rigidity of TCP is not relevant to
17
the physiological environment. The 3D scaffold provided conditions that mimic ECM for hMSC
18
growth. Additionally, the MC-based scaffolds could force the cells into an elongated and highly
19
branched osteogenic morphology, which benefits hMSC differentiation [48].
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Moreover, the physical parameters, such as elasticity and morphology of ECM are known to
21
affect the cellular process and regulate stem cell fate [49]. The biomimetic MC scaffolds with a
22
range of Young’s modulus were prepared to investigate whether differentiation of hMSCs was
23
affected by the mechanical properties of 3D substrates. It has been found that MSCs could
24
recognize the mechanical signals from the microenvironment around them via cell membrane
25
reporters and the related intracellular proteins [28]. Discher’s group reported that
26
mechano-transduction played an important role in regulating stem cell fate [47]. They found the
27
elastic property of the substrates alone could induce osteogenic differentiation. Mechanical strain
28
could act as a stimulator to induce differentiation of MSCs into osteoblasts [50], while optimal
29
stiffness of substrates could induce differentiation of MSC into osteoblasts [29]. In this work, we 15 Page 15 of 24
developed the MC-based substrates with controllable Young’s modulus by changing the
2
cross-linking density. We found that the ALP activity, osteocalcin secretion and targeted gene
3
expression were increased with increasing the Young’s modulus of MC scaffolds. Thus, the
4
mechanical property was likely to regulate the osteogenic specification, as previously reported
5
[41, 51]. We also found that the MC-based substrates could mediate hMSCs differentiation
6
without induction, probably due to both the ECM-mimic effect and the mechanical property of
7
the substrates. Therefore, these 3D biomimetic MC scaffolds are expected to be ideal biomaterials
8
in bone regenerative medicine.
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4.
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In summary, we have prepared the 3D MC scaffolds and applied them as scaffolds for directing
12
osteogenic differentiation of hMSCs. These biodegradable 3D scaffolds could afford ECM
13
biomimetic microenvironment for hMSCs growth. Different cross-linking densities lead to
14
different mechanical properties of the scaffolds. The experiments demonstrated that the MC
15
scaffolds directed hMSCs differentiation towards osteoblasts without induction. Moreover, the
16
MC-based scaffolds with higher Young’s modulus can efficiently improve the expression of
17
osteogenic marker genes, activity of ALP, and secretion of osteocalcin, and thereby accelerate
18
osteogenic differentiation. We found that the ECM-mimetic effect and mechanical properties of
19
the 3D MC scaffolds facilitated hMSC differentiation even without inducer supplement. This
20
work provides a paradigm to design new types of 3D substrates for stem-cell fate direction, and
21
thereby may find diverse applications in regenerative medicine.
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Conclusions
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Acknowledgements
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We acknowledge financial support of this work from National Natural Science Foundation of
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China (No. 51361130033), the Ministry of Science and Technology of China (No.
26
2014CB965003), and Strategic Priority Research Program of the Chinese Academy of Science
27
(XDA01030101). H. Shen thanks the Collaborative Academic Training Program for Post-doctoral
28
Fellows at Collaborative Innovation Center of Suzhou Nano Science and Technology.
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16 Page 16 of 24
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Figure 1. (A) SEM images of LCL, MCL, HCL, and LP scaffolds, respectively. The scale bar in
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each image is 1 mm. (B) Equilibrium swelling ratio and (C) crosslinking density of the LCL,
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MCL, and HCL scaffolds, respectively.
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Figure 2. (A) Tensile stress, (B) ultimate strain, and (C) Young’s modulus of LCL, MCL, HCL,
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and LP scaffolds, respectively.
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Figure 3. Confocal fluorescence microscopy images of hMSCs on (A) LCL, (B) MCL, and (C)
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HCL scaffolds, respectively. As shown in (D), images (a), (b) and (c) of each sample were taken
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at different altitude from top to bottom of these scaffolds.
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Figure 4. Osteogenic differentiation visualized by (A) ALP staining, and Alizarin red S staining
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after 20 days of hMSCs incubation onto TCP, LCL, MCL, and HCL scaffolds, respectively. (B)
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ALP activity of hMSCs on different substrates (TCP, LCL, MCL, and HCL scaffolds) after 7 d,
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14 d, 21 d, and 33 d co-culture, without induction. (C) Osteocalcin secretion of the hMSCs
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cultured onto TCP, LCL, MCL, and HCL scaffolds for 14 d and 21 d, respectively. (* p<0.05)
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Figure 5. RT-PCR for variation of osteoblasts marker genes of (A) ALP and (B) Ocn of hMSCs
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cultured on TCP, LCL, MCL, and HCL scaffolds at 5 d, 10 d, 15 d and 20 d, respectively. (*
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p<0.05)
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S0927-7765(15)30096-5 http://dx.doi.org/doi:10.1016/j.colsurfb.2015.07.062 COLSUB 7266
To appear in:
Colloids and Surfaces B: Biointerfaces
Received date: Revised date: Accepted date:
27-3-2015 18-6-2015 22-7-2015
Please cite this article as: H. Shen, Y. Ma, Y. Luo, X. Liu, Z. Zhang, J. Dai, Directed osteogenic differentiation of mesenchymal stem cell in three-dimensional biodegradable methylcellulose-based scaffolds, Colloids and Surfaces B: Biointerfaces (2015), http://dx.doi.org/10.1016/j.colsurfb.2015.07.062 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.
Graphical Abstract
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251658240
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Preparation of MC scaffolds with three-dimensional porous structure for regulating the fate of MSCs. Direction osteogenic differentiation of MSCs via MC scaffolds. Stimulatory effect of mechanical properties and 3D structure of the MC scaffolds on inducing the stem cell differentiation without osteogenic inducer supplement.
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Highlights
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Directed osteogenic differentiation of mesenchymal stem cell in three-dimensional
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biodegradable methylcellulose-based scaffolds
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He Shen, a Yufei Ma, a Yu Luo, a, c Xiaoyun Liu, a Zhijun Zhang,*a and Jianwu Dai*a, b
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1 Page 1 of 24
1 2 3
a
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Nano-tech and Nano-bionics, Chinese Academy of Sciences (CAS), 398 Ruoshui Road, Suzhou,
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215123, China
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b
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College of Preventive Medicine, Third Military Medical University, 30, Gaotanyan Road,
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Chongqing 400038, China
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c
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Institute of Combined Injury, State Key Laboratory of Trauma, Burns and Combined Injury,
College of Chemistry, Chemical Engineering and Biotechnology, Donghua University, Shanghai,
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Key Laboratory of Nano-Bio Interface, Division of Nanobiomedicine, Suzhou Institute of
201620, China
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* Corresponding authors. Email: [email protected], Tel.: +86-10-82614426; Email:
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[email protected], Tel.: +86-512-62872556.
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Abstract:
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Development of three-dimensional (3D) biodegradable scaffolds that can accelerate mesenchymal
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stem cell (MSC) osteogenic differentiation is a decisive prerequisite for treatment of damaged
4
skeletal tissue. We report herein the preparation of methylcellulose-based (MC) scaffolds using
5
carbonyldiimidazole as cross-linking agent to produce substrates with specific cross-linking
6
density and porous structure, as well as their applications for directing hMSC toward osteoblasts.
7
The mechanical properties of the scaffolds were controlled by cross-linking density. Human
8
MSCs (hMSCs) seeded on the MC scaffolds have penetrated into the pores, and showed high
9
viability (> 80%) as revealed by WST assay and Live/Dead assay. Moreover, the results of
10
differentiation experiments indicated that hMSCs cultured on MC substrates displayed high level
11
of osteogenic differentiation marker expression, alkaline phosphatase activity and osteocalcin
12
secretion, suggesting that the MC scaffolds can direct hMSC differentiation towards osteoblasts
13
without inducer treatment and cross-linking density of MC scaffolds have stimulatory effect on
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inducing differentiation. The 3D MC scaffolds could be applicable as promising scaffolds for
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bone tissue repair.
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Keywords: three-dimensional scaffold; methylcellulose; cross-linking density; mesenchymal
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stem cell; osteogenic differentiation.
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1.
Introduction
2
The regenerative medical applications of biomaterials pursue the reconstruction of damaged
3
tissues by controlling the fate of the cells cultured on implanted scaffolds [1]. The biomaterials
4
with rationally designed three-dimensional (3D) architectures are ideal tissue engineering
5
scaffolds due to their close emulation of the topographies and spatial structures of native
6
extracellular matrices (ECMs) that facilitate stem cell proliferation and differentiation [2]. Cells
7
cultured in vitro within these artificial 3D structures are subjected to both biochemical and
8
biophysical stimuli, and then coaxed to specific cell lineages. Therefore, 3D scaffolds can control
9
the stem cell fate more precisely than 2D substrates [3, 4]. To date, numerous 3D scaffolds have
10
been developed, such as 3D architectures based on chitosan [5], polylactic acid [6], poly
11
(ethylene glycol) (PEG) [7], graphene [8-11], and collagen [12, 13], as well as hydrogels based
12
on matrigel [14], hyaluronic acid [15], and alginate [16, 17]. Scaffolds with porous internal
13
architecture are considered as ideal biomaterials for tissue engineering as they permit efficient
14
diffusion of nutrients and metabolism products, and can be tuned to yield a favorable
15
microenvironment that influences cell proliferation and differentiation [18, 19]. Hence, these 3D
16
porous scaffolds have been extended to be in combination with mesenchymal stem cells (MSCs)
17
for the utilization in cartilage, skeletal tissue, and smooth muscle tissue engineering [20-23].
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With the goal towards tissue regeneration, development of 3D scaffolds that mimic ECM, as
19
well as provide an environment and necessary signals for controlling stem cell fate, is highly
20
desired [24]. Two of the important scaffold parameters are cross-linking density and pore size,
21
which not only affect, to a large extent, the mechanical properties of the scaffolds, but also
22
impose impacts on stem cells metabolism [25, 26]. It has been demonstrated that MSCs are able
23
to recognize mechanical signals. Complex sensory machinery can involve a group of cell surface
24
receptors and intracellular proteins that mediate mechanical signals from substrate to regulate a
25
variety of gene expressions [27]. Chatterjee and co-workers have reported that the stiffness of
26
PEG hydrogels regulates osteogenic differentiation of the encapsulated stem cells [28]. The
27
synthetic scaffolds used for skeletal tissue engineering should ideally act as artificial ECM for
28
regenerative applications, and provide suitable condition for directing of MSCs differentiation
29
into specific cell lineage. However, these investigations were carried out on 2D tissue-culture
30
platforms. Whether MSC cultured on 3D substrates could be induced differentiation by 4 Page 4 of 24
1
mechanical property is still unknown. The regulation of MSC differentiation by the mechanical
2
properties of 3D scaffolds needs further investigation. Methylcellulose (MC) is a chemically modified polysaccharide with a partial substitution of
4
hydroxyl groups with methoxy moieties. Its biocompatibility and capability of gelation make it a
5
food additive for thickening [29]. In order to control the gelling process, MC has been combined
6
with different components, such as agarose [30] and hyaluronan (HA) [31], leading to different
7
characteristics of the MC-based scaffolds. Recently, MC-based biomaterials have been explored
8
extensively for tissue regenerative applications, such as supporting substrates and regenerative
9
factor carriers [32]. Tate et al. have investigated MC as a drug delivery/tissue engineering
10
scaffold for treating traumatic brain injury [30]. Taking the advantages of MC, Hsieh et al. further
11
prepared MC-based scaffolds combined with HA for neural stem/progenitor cell culture [24].
12
Although some progress has been made on applications of MC-based scaffolds in tissue
13
regeneration, the effects of their physical properties on stem cell differentiation were actually
14
ignored.
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In the investigation, a series of porous nanostructured MC-based scaffolds with different
16
cross-linking density and mechanical properties were developed for bone tissue engineering. We
17
synthesized various MC-based scaffolds with low degree of cross-linking (LCL), medium degree
18
of cross-linking (MCL), and high enough degree of cross-linking (HCL). In order to address the
19
role of pores in the MC-based substrates, we used a scaffold with a low porosity level (LP) as a
20
control. The morphology, degradation, wettability and mechanical properties of the MC scaffolds
21
were studied. In addition, the toxicity and hemocompatibility of these biomaterials were
22
examined. Finally, osteoblast responses towards these 3D scaffolds were investigated. Directing
23
osteogenic differentiation of the hMSC cultured on the MC-based substrates with 3D structures
24
provides useful insights into designs of new scaffolds for bone tissue engineering applications.
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2.
Materials and methods
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2.1 Preparation of 3D porous MC-based scaffolds
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To prepare MC-based scaffolds, MC (500 mg) was first mixed with gelatin (50 mg) and agarose
29
(50 mg) in 25 mL dimethyl sulfoxide (DMSO) to form a homogeneous solution. 5 g ground-NaCl 5 Page 5 of 24
was then added into the above solution, followed by vigorous stirring for 4 h. Cross-linking agent
2
(carbonyldiimidazole, CDI) was added into the above homogenized mixed solution at a varying
3
concentration, and then the resulted solution was vigorously agitated for 15 min. MC scaffolds
4
with low degree of cross-linking (LCL), medium degree of cross-linking (MCL), and high degree
5
of cross-linking (HCL) were prepared by adding 250 mg, 500 mg, and 750 mg CDI into the
6
solution, respectively. To synthesize the MC scaffold with a low porosity level (LP), MC (500
7
mg) was mixed with 50 mg gelatin and 50 mg agarose in 25 mL DMSO, and then 500 mg CDI
8
was added. These mixture solutions were incubated in ice bath for scaffold formation, and then
9
washed with distilled water 5 times to remove NaCl, DMSO and unreacted reagents. Finally,
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these samples were frozen and lyophilized before use.
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2.2 Morphology of the MC-based scaffolds
12
Microstructure of the scaffolds was examined by scanning electron microscope (SEM) with an
13
accelerating voltage of 10 kV. All MC-based samples were sputter coated with a 10 nm gold film
14
before SEM observation.
15
2.3 Swellability of the MC-based scaffolds
16
Equilibrium swelling ratio (QS), a measure of swellability of the MC-based scaffolds, was
17
estimated by immersing dried the MC samples to swell in PBS at 37 oC for 7 days. QS was
18
defined according to the following equation:
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where Ws and Wd are the weight of the swollen sample and the dried sample, respectively [5].
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In addition, swelling measurement was performed to estimate the cross-linking density of
22
MC-based scaffolds, as the degree of swelling is known to be dependent upon the crosslinking
23
density of the scaffolds. The cross-linking densities of the MC-based scaffolds were calculated
24
from the Flory-Rehner equation as follows:
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where Vp, Vo, Dp, Do, and x are the volume fraction of polymers in the swollen mass, molar
3
volume of solvent, density of polymers, density of solvent, solvent-polymer interaction term,
4
respectively [33]. Mean and standard deviations for the triplicate samples for each scaffold were
5
reported.
6
2.4 Mechanical testing
7
Prior to mechanical testing by a material testing machine (H5K-S, Hounsfield, UK), all frozen
8
dried samples were cut into small strips, and then the width, length, and thickness were measured
9
with a micrometer, and three strips of each scaffold were chosen for the mechanical test. Stress
cr
us
and strain were calculated according to the following equations:
an
10
ip t
2
M
11
12
where σ, ε, P, w, d, l and l0 stand for stress, strain, load, mat width, mat length, extension length,
14
and gauge length, respectively. Young’s modulus can be calculated from stress–strain curves [34,
15
35].
16
2.5 Degradation behavior of the MC-based scaffolds
17
The degradation ratio of the MC-based scaffolds was measured after 120 d incubation. LCL,
18
MCL, HCL and LP samples, in triplicate, were incubated in physiological saline solution
19
containing 0.01% trypsin at 37 oC, respectively. The dried MC samples were weighed after being
20
frozen and lyophilized. The degradation ratio (Rd) of MC-based scaffolds was calculated
21
according to the following equation:
Ac ce pt e
d
13
22 23
where W0 and W120 represent the initial dried weight and the dried weight of the sample at day
24
120, respectively. 7 Page 7 of 24
2.6 Cell adhesion, growth, and osteogenesis analysis
2
Small pieces of human umbilical cord were washed with PBS and transferred into tissue culture
3
plates. The culture medium was changed every 2 days until 80% confluence was reached.
4
DMEM/F12 (50/50) supplemented with 10% FBS, 100 U/mL penicillin, and 100 μg/mL
5
streptomycin solution were used throughout the routine culture procedures. Human MSCs
6
(passage 5-8) were seeded at a density of 2×104, 1×104 and 1×104 cells per well (24-well plates,
7
Corning, USA) for the cell adhesion, growth, and osteogenesis assay, respectively.
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1
Typically, for cell attachment assay, cell suspensions were seeded onto the different scaffolds,
9
and cultured for 1, 3, 6 and 12 h, respectively. After incubation, MC-based substrates were
10
washed to remove the non-adherent cells. The cell viability was determined via WST assay
11
according to the manufacturer’s instruction. 5 days after incubation, Calcein-AM (CAM) and
12
propidium iodide (PI) staining were used to detect the live and dead cells cultured on MC-based
13
scaffolds and TCP according to the instruction. To evaluate the extent of proliferation, hMSCs
14
were cultured onto the scaffolds. After 7, 14, 24 and 33 d treatment, cell density was evaluated by
15
total double-stranded DNA content using the PicoGreen DNA quantification kit.
M
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cr
8
For cell differentiation analysis, cell suspensions were seeded onto different scaffolds, and
17
cultured for 7, 14, 24 and 33 d, respectively. The osteogenic differentiation trends were
18
determined by alkaline phosphatase (ALP) activity assay and calcium phosphate secretion test,
19
and reverse transcription polymerase chain reaction (RT-PCR). All measurements were carried
20
out in three independent experiments.
21
2.7 ALP activity
22
ALP activity was assessed by the substrate p-nitrophenyl phosphate in an end-point assay. The
23
generation of p-nitrophenol production is proportional to the amount of ALP in the solution.
24
Standard calibration curve was prepared by p-nitrophenol dilution. The absorbance was read at
25
405 nm and then ALP content was calculated.
26
2.8 Osteocalcin secretion
27
An intact human osteocalcin EIA kit BT-460 (Biomedical Technologies Inc., USA) was used for
28
measurement of osteocalcin secretion. On day 13 and 20, the medium was replaced with fresh
29
DMEM/F12 (50/50) medium (without FBS). After 24 h of culture, the medium was harvested and
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8 Page 8 of 24
the osteocalcin content was analyzed using the kit according to our previous work [35].
2
2.9 Confocal imaging
3
After incubation for 7 d, the hMSCs cultured on MC-based substrates were fixed in 4% formalin
4
and the cell nucleus was stained by DAPI. Confocal fluorescence microscopy images were
5
captured at different altitude to investigate the cell distribution in the 3D scaffolds.
6
2.10 Histological staining
7
After 20 d culturing, hMSCs on LCL, MCL, HCL, and LP substrates were stained by different
8
agents to study the expression of ALP and mineralization of bone nodules. ALP activity was
9
detected histochemically by incubation with a mixture of naphthol AS-MX phosphate and fast
10
violet B salt solution [36]. For the Alizarin red staining, scaffolds were incubated with Alizarin
11
red [37].
12
2.11 RNA isolation and RT-PCR assay
13
To investigate the cell phenotype, mRNA was analyzed using RT-PCR. After 5, 10, 15 and 20 d
14
co-culture, the specimens were washed by PBS, and then suspended in TRIzol Reagent (Life
15
Technologies Co.) to extract total RNA. Then the total RNA was reverse transcribed by using
16
transcriptase reaction mix (SuperScript III First-Strand Synthesis System, Life Technologies) for
17
cDNA generation.
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1
18
Quantitative PCR analysis was performed in triplicate using power SYBR green RT-PCR kit
19
(Life Technologies) on a CFX96 system (Bio-Rad, USA). Glyceraldehyde-3-phosphate
20
dehydrogenase (GAPDH) was used as an endogenous housekeeping gene. The data were
21
calculated using the ΔΔCt method as previously reported [12]. The cells on TCP substrate were
22
set as control group, and their relative levels of marker genes expression were artificially set as
23
one-fold. The genes and primer sequences are listed in Table S1.
24
2.12 Statistical analysis
25
The data were expressed as mean ± SD (n = 3). Statistical analysis was performed and p < 0.05
26
was considered as statistically significant.
27 28
3.
Results 9 Page 9 of 24
3.1 Characterization of the MC-based scaffolds
2
SEM images indicate that the MC scaffolds have three-dimensional, porous, and foamy structures
3
(Figure 1A), which may facilitate the formation of bridges across the lesion sites of damaged
4
tissue [2]. The freeze-dried scaffolds can be cut into pieces with any desired shape for clinical
5
applications. As shown in Figure 1A, numerous pores with diameters ranging from 100 μm to
6
600 μm were observed on LCL, MCL and HCL scaffolds, allowing nutrient diffusion through
7
these pores. NaCl particles play a vital role in the pore generation. Low porosity level was
8
observed on LP scaffold, which was prepared without addition of NaCl particles.
cr
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1
To confirm the cross-linking densities of LCL, MCL, and HCL scaffolds, the swelling ratios
10
of these MC scaffolds were studied. The swelling ratios of the MC-based scaffolds were
11
significantly infulenced by CDI concentration (Figure 1B) The swelling rates of LCL, MCL, and
12
HCL scaffolds were 108.20, 35.28, and 21.52, respectively. The cross-linking density of
13
MC-based scaffolds is shown in Figure 1C. The cross-linking density of the LCL, MCL, and
14
HCL scaffolds was 0.003, 0.011, and 0.025 ×10 -5mol cm3, respectively. The cross-linking density
15
of the MC scaffolds increased with increasing amount of CDI.
M
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9
The mechanical properties of scaffolds can be readily controlled by tuning concentrations of
17
polymers or cross-linking molecules. The experiment demonstrated that tensile, stress, and
18
Young’s modulus of the MC-based scaffolds could be regulated by changing the CDI
19
concentration (Figure 2). With increasing amount of the cross-linker, the Young’s modulus of the
20
MC-based scaffolds was significantly increased. In addition, water diffused in the MC-based
21
scaffolds very quickly. The water diffusion time of LCL, MCL and HCL scaffolds was
22
determined to be 10-20 s (Figure S2A). A favorable and biodegradable scaffold provides
23
structural support for initial cell growth and then would gradually degrade after new tissue
24
formation [38, 39]. Thus we investigated the degradation behavior of the MC scaffolds. As shown
25
in Figure S2B, LCL scaffold displayed less than 30 % of the original mass, while HCL and LP
26
scaffolds lost nearly 21 % and 8 % weight, respectively, after 120 days of incubation in vitro. The
27
degradation of the MC-based scaffolds was also dependent on their cross-linking ratios. Higher
28
cross-linking density and lower porosity improved the stability of the HCL and LP scaffolds.
29
3.2 In vitro toxicity of MC-based scaffolds
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10 Page 10 of 24
As biocompatibility is an essential issue for tissue engineering scaffolds, we studied the in vitro
2
toxicity of MC-based scaffolds by WST assay after hMSCs cultured with LCL, MCL, HCL, and
3
LP for 3 and 5 d, respectively. The viability of MSCs was used to assess the cytotoxicity of
4
scaffolds on which MSCs were cultured. For comparison purpose, traditional TCP substrate was
5
employed as a control group (100 % cell viability). As shown in Figure S3, all 3D scaffolds did
6
not show any significantly deleterious effects on MSC viability, indicating their good
7
biocompatibility.
ip t
1
The biocompatibility of the scaffolds was also evaluated by hemolysis test. The release of
9
hemoglobin was used to quantify the membrane-damaging properties of the LCL, MCL, HCL,
10
and LP scaffold. Distilled water and PBS buffer treated erythrocytes were set as 100% and 0%
11
values, respectively. As Figure S4 shown, LCL, MCL, HCL and LP scaffold caused 3.0 %, 1.8 %,
12
1.8 % and 1.5 % hemolysis, respectively. All these four types of MC scaffolds exhibited neither
13
detectable disturbance of the RBC membranes, nor any hemolytic phenomenon, suggesting that
14
the MC scaffolds have excellent hemocompatibility, an important premise for their in vivo
15
regenerative medical applications.
16
3.3 Human MSC growth on 3D MC-based scaffolds
17
In this experiment, the cell distribution inside the 3D MC scaffolds was studied by taking
18
confocal fluorescence microscopy images at different altitude from top to bottom of these
19
scaffolds across the central part (Figure 3). A spatial arrangement of the hMSCs was observed on
20
these 3D scaffolds. The MSCs cultured for 7 days penetrated into the pores of the scaffolds and
21
formed aggregates, which may reestablish cell-cell contacts (Figure 3A-C) [11]. However, only a
22
few hMSCs were detected inside the LP substrate owning to its low porosity, suggesting the
23
limitation of cell migration and growth in this case (Figure S5). Additionaly, CAM/PI based
24
Live/Dead cell viability assay was used to assess live and dead cells cultured on the scaffolds. As
25
shown in Figure S6, most of the cells were live (>95%) on the MCL scaffold and TCP after 5
26
days culture.
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27
In order to understand how hMSCs proliferated on the MC-based scaffolds, DNA was
28
extracted from the LCL, MCL, HCL, and LP substrates, as the amount of genes is proportional to
29
the cell number. At the initial stage, only a few cells have adhered onto LCL, MCL, HCL, and LP 11 Page 11 of 24
substrate (Figure S7), so the hMSCs cultured on the 3D scaffolds showed slow cell proliferation
2
rate, compared to traditional 2D scaffold, on day 7 and 14 (Figure S8). However, with increasing
3
incubation time, proliferation of hMSCs on TCP was inhibited due to limited culture space. The
4
MSCs cultured on LCL, MCL, and LCL scaffolds showed a lag phase of growth, giving rise to a
5
slow rate of proliferation. Inhibition of hMSC growth was also observed on the LP scaffold, due
6
to its low porosity level and restricted space. It turns out that the LP scaffold is not a good
7
scaffold material for bone tissue reparation, because the density of hMSCs on the LP substrate is
8
too low for practical applications.
9
3.4 Osteogenic differentiation on 3D porosity scaffolds
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The cell phenotype on different MC-based scaffolds was further investigated to figure out how
11
MSCs differentiated on these 3D substrates. Alkaline phosphatase (ALP) activity is one of the
12
most commonly used markers for osteogenesis differentiation, as ALP enzyme is involved in the
13
early outset of mineralization and is widely found in newly formed bone tissue. The osteogenesis
14
differentiation was first examined by ALP staining. The stained ALP positive areas appeared dark
15
red color, as shown in Figure 4A. The expression of ALP was observed from the cells cultured on
16
LCL, MCL, and HCL scaffolds, whereas it was not present in the control group (TCP).
d
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The calcium deposition and collagen secretion of cells on the MC-based scaffolds were also
18
investigated, as they play important roles in the mineralization during bone reparation [40].
19
Alizarin red S staining was utilized for the qualitative assessment of mineralized matrix
20
formation. As shown in Figure 4A, the calcium and mineral depositions could be clearly
21
observed as red regions on LCL, MCL, and HCL scaffolds incubated with hMSCs for 20 d. The
22
results suggest that the LCL, MCL, and HCL scaffolds could induce hMSCs to differentiate into
23
osteoblasts after 20 d of incubation, even without the treatment of osteogenic inducer such as
24
dexamethasone or bone morphogenetic protein-2.
Ac ce pt e
17
25
The quantitative analyses of ALP activity and extracellular osteocalcin production were
26
further carried out to study the cell differentiation behavior. The ALP activity of hMSCs cultured
27
on LCL, MCL, and HCL scaffold as well as TCP (control group) was studied after 7, 14, 21 and
28
33 days of incubation, respectively (Figure 4B). For the first 7 days, ALP activity of hMSCs on
29
both TCP and MC-based substrates presented very low activity levels. During the next 7 days, 12 Page 12 of 24
there was a significant increase in the ALP activity for 3D substrate samples, compared with the
2
control group (TCP). The ALP activity of hMSCs on LCL, MCL, and HCL scaffolds was much
3
higher than that on TCP substrate at 14, 21 and 33 d, indicating the enhancement of hMSC
4
differentiation towards osteoblasts lineage. However, the ALP activity of the cells on LCL, MCL,
5
and HCL scaffolds decreased after 21 and 33 d culture, compared to the 14 d culture. We assumed
6
that the hMSCs cultured on LCL, MCL and HCL scaffolds were induced to osteogenic
7
differentiate at first 14 days. Thus, most of the cells on MC-based scaffolds tended to be
8
osteoblasts at 21 and 33 d, which leads to low levels of ALP activity of cells.
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The extracellular osteocalcin production contents of the hMSCs cultured onto TCP, LCL,
10
MCL, and HCL scaffolds were also examined after14 and 21 d co-culturing, respectively. In the
11
absence of osteogenic inducer (dexamethasone) in the culture medium, hMSCs grown on TCP
12
substrate secreted little osteocalcin, because most of the cells on TCP maintained their
13
self-renewal property. As expected, the osteocalcin production contents of hMSCs on LCL, MCL,
14
and HCL scaffolds were remarkably higher than that on TCP at 14 d and 21 d (Figure 4C). These
15
data further indicated the progression of osteogenic differentiation on the 3D substrates. Similar
16
to ALP activity, the cells cultured on the MC-based scaffolds at 14 d exhibited higher osteocalcin
17
contents than those on the same sample group at 21 d, as most of hMSCs on the MC-based
18
scaffolds differentiated into osteoblasts at 14 d.
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19
To gain a profound understanding of the effects of 3D MC-based scaffolds on directing
20
hMSC fate, targeted osteoblastic gene markers were quantitatively assessed by RT-PCR after 5,
21
10, 15 and 20 d treatment, respectively. Firstly, relative levels of ALP gene expression were all
22
up-regulated in the hMSCs cultured on the LCL, MCL, and HCL scaffolds, compared with those
23
seeded on TCP substrate after 10, 15 and 20 d co-culturing (Figure 5A). Osteocalcin (Ocn), a late
24
stage osteogenesis and mineralization marker during bone formation, was also examined. With
25
the treatment of the MC-based scaffolds, the Ocn expression was enhanced in comparison with
26
treatment of TCP, suggesting the induced osteogenic differentiation of hMSCs via MC-based
27
platforms. Additionally, the osteoblastic marker genes expression levels of hMSCs on MC
28
scaffolds is higher at 15 d than those at 5, 10 and 20 d (Figure 5B). Hence, the experimental
29
results indicated that hMSCs cultured on LCL, MCL, and HCL scaffolds were more osteogenic
30
and produced more minerals than any other time periods during the first two weeks. More 13 Page 13 of 24
1
importantly, the expression levels of ALP and Ocn gene of hMSCs cultured on the scaffolds were
2
increased by raising the cross-linking density of the MC-based substrates.
3
4.
Discussion
5
The presence of pores in scaffolds is necessary for bone tissue reparation because they allow
6
MSC proliferation as well as vascularization. Additionally, it has been reported that porosity
7
could stimulate osteogenesis in vitro [27]. The pore size of the substrate is recommended to be
8
200 - 900 μm for accelerating stem cell growth and differentiation [41]. Although after
9
lyophilization the pores created through salt leaching will change their dimensions completely,
10
LP scaffold without addition of NaCl particles during prepartions displayed low porous level,
11
which might restrict cell migration and growth, as well as nutrition exchange. The mixing of
12
gelatin and agarose with MC usually benefits 3D scaffolds for cell growth. Martin et al. reported
13
that MC and agarose hydrogel blends and their applications of nerve regeneration which
14
combined of agarose’s ability to support nerve growth and the gelling characteristics [31].
15
Additionally, gelatin, a denatured collagen, is a biodegradable polymer with high biocompatibility,
16
which has been considered as a suitable substrate for cells [42]. It's also demonstrated that gelatin
17
scaffolds gelatin-based scaffolds are favorable for cell adhesion and growth [43].
Ac ce pt e
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4
18
In addition, surface wettability of biomaterials relates to their surface properties for cell
19
adhesion and proliferation [44]. The good wettability of these 3D substrates is also attributed to
20
their porous structure and abundant hydrophilic groups. Besides, water diffusion time of LP
21
scaffold is much longer than that of LCL, MCL and HCL scaffolds, probably due to that the low
22
porosity level of LP scaffold inhibited the water diffusion and wetting. Furthermore, a trend of
23
decreasing swelling ratio was observed with increasing the concentration of the crosslinking
24
agent. The swelling ratio of the MC scaffold increased, while the cross-linking density of the
25
polymer network decreased [33, 45].
26
Moreover, the cross-linking density and porosity of MC-based scaffolds also impose
27
significant impacts on mechanical property, which is another important parameter of scaffolds for
28
bone tissue engineering [46]. It has been elegantly demonstrated that the substrate elasticity
29
induces MSC differentiation into specific lineages [47]. The 3D space of the MC scaffolds is 14 Page 14 of 24
available for hMSCs proliferation during a long period of time. The porous MC-based substrates
2
allowed cell migration and efficient exchange of nutrients and wastes, as well as provided
3
appropriate mechanical condition for MSCs growth. After long time incubation with these
4
MC-based substrates, the density of the MSCs became higher than that of the control group
5
(TCP). More importantly, the MC-based substrates with different mechanical properties
6
stimulated the related gene expression and osteogenic differentiation without induction.
ip t
1
The above osteogenic marker analysis revealed that the porous 3D MC-based substrates
8
promote differentiation of MSCs without osteogenic inducer treatment. Interestingly, we found
9
that the ALP activity, osteocalcin secretion, and marker gene expression of the cells incubated
10
with HCL substrate were higher than those of the hMSCs treated with LCL. The origin of
11
directed hMSCs differentiation on the 3D porous platforms could be mainly attributed to two
12
reasons [48]: one is that the porous structure of the 3D MC scaffolds provides ECM biomimetic
13
microenvironment for hMSCs growth and differentiation; the other is that the mechanical
14
properties of LCL, MCL, and HCL scaffolds induced differentiation of the hMSCs.
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7
In this research, 2D tissue culture platform exhibited a very low level of osteogenic
16
differentiation without dexamethasone treatment, because the rigidity of TCP is not relevant to
17
the physiological environment. The 3D scaffold provided conditions that mimic ECM for hMSC
18
growth. Additionally, the MC-based scaffolds could force the cells into an elongated and highly
19
branched osteogenic morphology, which benefits hMSC differentiation [48].
Ac ce pt e
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15
20
Moreover, the physical parameters, such as elasticity and morphology of ECM are known to
21
affect the cellular process and regulate stem cell fate [49]. The biomimetic MC scaffolds with a
22
range of Young’s modulus were prepared to investigate whether differentiation of hMSCs was
23
affected by the mechanical properties of 3D substrates. It has been found that MSCs could
24
recognize the mechanical signals from the microenvironment around them via cell membrane
25
reporters and the related intracellular proteins [28]. Discher’s group reported that
26
mechano-transduction played an important role in regulating stem cell fate [47]. They found the
27
elastic property of the substrates alone could induce osteogenic differentiation. Mechanical strain
28
could act as a stimulator to induce differentiation of MSCs into osteoblasts [50], while optimal
29
stiffness of substrates could induce differentiation of MSC into osteoblasts [29]. In this work, we 15 Page 15 of 24
developed the MC-based substrates with controllable Young’s modulus by changing the
2
cross-linking density. We found that the ALP activity, osteocalcin secretion and targeted gene
3
expression were increased with increasing the Young’s modulus of MC scaffolds. Thus, the
4
mechanical property was likely to regulate the osteogenic specification, as previously reported
5
[41, 51]. We also found that the MC-based substrates could mediate hMSCs differentiation
6
without induction, probably due to both the ECM-mimic effect and the mechanical property of
7
the substrates. Therefore, these 3D biomimetic MC scaffolds are expected to be ideal biomaterials
8
in bone regenerative medicine.
cr
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1
9
4.
11
In summary, we have prepared the 3D MC scaffolds and applied them as scaffolds for directing
12
osteogenic differentiation of hMSCs. These biodegradable 3D scaffolds could afford ECM
13
biomimetic microenvironment for hMSCs growth. Different cross-linking densities lead to
14
different mechanical properties of the scaffolds. The experiments demonstrated that the MC
15
scaffolds directed hMSCs differentiation towards osteoblasts without induction. Moreover, the
16
MC-based scaffolds with higher Young’s modulus can efficiently improve the expression of
17
osteogenic marker genes, activity of ALP, and secretion of osteocalcin, and thereby accelerate
18
osteogenic differentiation. We found that the ECM-mimetic effect and mechanical properties of
19
the 3D MC scaffolds facilitated hMSC differentiation even without inducer supplement. This
20
work provides a paradigm to design new types of 3D substrates for stem-cell fate direction, and
21
thereby may find diverse applications in regenerative medicine.
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22
Conclusions
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10
23
Acknowledgements
24
We acknowledge financial support of this work from National Natural Science Foundation of
25
China (No. 51361130033), the Ministry of Science and Technology of China (No.
26
2014CB965003), and Strategic Priority Research Program of the Chinese Academy of Science
27
(XDA01030101). H. Shen thanks the Collaborative Academic Training Program for Post-doctoral
28
Fellows at Collaborative Innovation Center of Suzhou Nano Science and Technology.
29
16 Page 16 of 24
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Figure 1. (A) SEM images of LCL, MCL, HCL, and LP scaffolds, respectively. The scale bar in
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each image is 1 mm. (B) Equilibrium swelling ratio and (C) crosslinking density of the LCL,
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MCL, and HCL scaffolds, respectively.
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Figure 2. (A) Tensile stress, (B) ultimate strain, and (C) Young’s modulus of LCL, MCL, HCL,
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and LP scaffolds, respectively.
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Figure 3. Confocal fluorescence microscopy images of hMSCs on (A) LCL, (B) MCL, and (C)
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HCL scaffolds, respectively. As shown in (D), images (a), (b) and (c) of each sample were taken
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at different altitude from top to bottom of these scaffolds.
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Figure 4. Osteogenic differentiation visualized by (A) ALP staining, and Alizarin red S staining
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after 20 days of hMSCs incubation onto TCP, LCL, MCL, and HCL scaffolds, respectively. (B)
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ALP activity of hMSCs on different substrates (TCP, LCL, MCL, and HCL scaffolds) after 7 d,
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14 d, 21 d, and 33 d co-culture, without induction. (C) Osteocalcin secretion of the hMSCs
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cultured onto TCP, LCL, MCL, and HCL scaffolds for 14 d and 21 d, respectively. (* p<0.05)
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Figure 5. RT-PCR for variation of osteoblasts marker genes of (A) ALP and (B) Ocn of hMSCs
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cultured on TCP, LCL, MCL, and HCL scaffolds at 5 d, 10 d, 15 d and 20 d, respectively. (*
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p<0.05)
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