Combinational effect of matrix elasticity and alendronate density on

Transcription

Acta Biomaterialia xxx (2015) xxx–xxx
Contents lists available at ScienceDirect
Acta Biomaterialia
journal homepage: www.elsevier.com/locate/actabiomat
Combinational effect of matrix elasticity and alendronate density
on differentiation of rat mesenchymal stem cells
Pengfei Jiang, Zhengwei Mao ⇑, Changyou Gao
MOE Key Laboratory of Macromolecular Synthesis and Functionalization, Department of Polymer Science and Engineering, Zhejiang University, Hangzhou 310027, China
a r t i c l e
i n f o
Article history:
Received 25 November 2014
Received in revised form 26 February 2015
Accepted 17 March 2015
Available online xxxx
Keywords:
Combinational effect
Stem cell differentiation
Gelatin
Stiffness
Alendronate
a b s t r a c t
Differentiation of mesenchymal stem cells (MSCs) is regulated by multivariate physical and chemical signals in a complicated microenvironment. In this study, polymerizable double bonds (GelMA) and osteoinductive alendronate (Aln) (Aln-GelMA) were sequentially grafted onto gelatin molecules. The biocompatible hydrogels with defined stiffness in the range of 4–40 kPa were prepared by using polyethylene
glycol diacrylate (PEGDA) as additional crosslinker. The Aln density was adjusted from 0 to 4 lM by controlling the ratio between the GelMA and Aln-GelMA. The combinational effects of stiffness and Aln density on osteogenic differentiation of MSCs were then studied in terms of ALP activity, collagen type I and
osteocalcin expression, and calcium deposition. The results indicated that the stiffness and Aln density
could synergistically improve the expression of all these osteogenesis markers. Their osteo-inductive
effects are comparable to some extent, and high Aln density could be more effective than the stiffness.
Ó 2015 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction
The mesenchymal stem cells (MSCs) have the capacity to differentiate into bone, cartilage, muscle, fat and a variety of other
connective tissues, leading to a great deal of interest in the field
of regenerative medicine and tissue engineering [1–6]. Recently,
growing evidence suggests that chemical, physical and mechanical
signals from materials and neighboring cells have a profound
impact on the differentiation of MSCs [7–10]. Therefore, it is of
paramount importance to precisely understand the interaction
between MSCs and the niche consisting of various chemical and
physical signals [11–17].
It is known that physical properties such as elasticity and topography of the extracellular matrix are able to dominate the fate of
stem cells. For instance, Engler et al. [18,19] demonstrated that
MSCs display characteristics of neurogenic, myogenic, and osteogenic phenotypes after being cultured on hydrogel substrates
mimicking the stiffness of neural, muscle, and bone tissues, respectively. McBeath et al. found that different sizes [20] of fibronectin
‘island’ can restrict MSC spreading and then dominate their differentiation. When the MSCs are allowed to adhere, flatten, and
spread they shall undergo osteogenesis, whereas the unspread
and round cells become adipocytes. Recently, it was found that different aspect ratio and subcellular curvature can modulate the
⇑ Corresponding author. Fax: +86 571 87951108.
differentiation of stem cells to adipocytes and osteoblasts [21].
Peng et al. [22–24] further made semi-quantitative investigation
of the effects of cell shape on differentiation of MSCs, and revealed
the optimal aspect ratios for adipogenic and osteogenic differentiation of MSCs. They found that the extents of both adipogenic and
osteogenic differentiations are linearly related to the cell perimeter, which reflects the non-roundness or local anisotropy of cells.
Not only the physical properties, various small functional
groups, peptides and proteins on both stiff substrates, i.e. silicon
wafers, and soft substrates, i.e. poly(ethylene glycol) (PEG) hydrogels, can modulate MSC differentiation [25–30]. Moreover, Kilian
and Mrksich demonstrated that the affinity and density of ligands
at the cell-biomaterial interface also can be engineered to influence
stem cell fate [31]. Among these molecules, alendronate sodium
(Aln) is a kind of bisphosphonate drug, which is able to promote
osteogenic differentiation of BMSCs via several mitogen-activated
protein kinase (MAPK) pathways, such as extracellular signalrelated kinases (ERKs) 1/2 and Jun amino-terminal kinases (JNK1/
2/3) pathways, in a dose-dependent manner [32–34]. Besides,
physically or chemically immobilized Aln also can induce
osteogenic differentiation of MSCs. Zhu et al. [34] created a density
gradient surface of Aln onto the polycaprolactone (PCL) membrane,
and found that MSCs over-express osteogenic marker proteins on
the surface, dependent on the local Aln density. Kim et al. [35]
demonstrated that physical immobilization of Aln and bone
morphogenic protein-2 on a titanium surface showed synergistic
effect on improving osteoblast activity.
E-mail address: zwmao@zju.edu.cn (Z. Mao).
http://dx.doi.org/10.1016/j.actbio.2015.03.018
1742-7061/Ó 2015 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
Please cite this article in press as: Jiang P et al. Combinational effect of matrix elasticity and alendronate density on differentiation of rat mesenchymal
stem cells. Acta Biomater (2015), http://dx.doi.org/10.1016/j.actbio.2015.03.018
2
P. Jiang et al. / Acta Biomaterialia xxx (2015) xxx–xxx
However, differentiation of stem cells usually happens in a
complicated microenvironment which contains multivariate signals [7,13]. Therefore, it is of paramount importance to understand
the impact of multivariate signals on stem cell fate, especially the
interplay between different types of signals. Sometimes these
signals have a synergistic effect on the differentiation of MSCs.
For example, Jiang et al. [36] demonstrated the synergistic effect
of nanofiber topography and released neuronal induction factor,
retinoic acid, on enhancing MSC neural commitment. Sometimes
the combinational effects became more complicated. Zouani et al.
[37] studied the effect of mechanical properties and special growth
factor in the same microenvironment on stem cell fate. Their
results demonstrate that chemical grafting on relative stiff
matrices (13–70 kPa) with an osteogenic factor (BMP-2 mimetic
peptide) results only in osteogenic differentiation. When grafted
on even softer hydrogel matrices (0.5–3.5 kPa), the BMP-2 mimetic
peptide has no effect on the stem cell differentiation. Therefore, in
order to predict the fate of MSCs in a complicated artificial environment, a more careful and case-sensitive study is required to understand the combinational effects of different types of signals.
In this work, the combinational effect of substrate stiffness and
alendronate density is studied in terms of MSCs differentiation
(Fig. 1b). Gelatin is chosen as the backbone of hydrogels due to
its good biocompatibility and potential of modification [38–40].
In order to fabricate the hydrogels with controllable mechanical
property, polymerizable double bonds are introduced onto the
gelatin molecules via the reaction between methacrylic anhydride
(MA) and amino groups of gelatin. Aln molecules are further
grafted onto the gelatin backbone through aldehyde-activated
reaction. Furthermore, polyethylene glycol diacrylate (PEGDA) is
used as crosslinker to modulate the crosslinking density of the
hydrogel and thus the stiffness (Fig. 1a). The combinational impact
of the hydrogel stiffness and Aln density on MSCs’ neuronal, myogenic, and osteogenic differentiation is first evaluated in terms of
expression of b-tubulin, MyoD, and calcium, respectively. Then
the osteogenic differentiation of MSCs, which is significantly influenced by these two factors in the current study, is studied in terms
of alkaline phosphatase (ALP) activity, expressions of collagen type
I and osteocalcin, and calcium deposition.
2. Experiment section
2.1. Materials
Gelatin (type B), polyethylene glycol diacrylate (PEGDA), 2,4,
6-trinitrobenzenesulfonic acid (TNBS), bovine serum albumin
(BSA), ascorbic acid, ammonium persulfate (APS), and tetramethylethylenediamine (TEMED) were purchased from Sigma–
Aldrich, USA. Methacrylic anhydride (MA) was bought from Alfa
Aesar, USA. Alendronate sodium (Aln) was obtained from
Spectrum, USA. PicoGreen dsDNA kit was purchased from life
technologies, USA. o-Cresolphthalein complexone (CPC), 8-hydroxyquinoline, and 2-amino-2-methyl-1-propanol (AMP) were
obtained from TCI chemical, Japan. Other chemicals were of analytical grade and used as received. The water used in the experiments was purified by a Milli-Q water system (Millipore, USA).
2.2. Synthesis and characterization of methacrylated gelatin
Methacrylated gelatin (GelMA) was synthesized according to
the method reported previously [41,42]. Briefly, 4 g gelatin was
dissolved in 40 mL phosphate buffer (pH = 8.0) at 70 °C. After being
cooled to 45 °C, 40 lL MA was added at a rate of 10 lL/min under
stirring, and the mixture was allowed to react for 1 h. Into the
solution 500 mL cold ethanol ( 20 °C) was added to precipitate
the methacrylated gelatin. After centrifugation, the precipitates
were dissolved in water, sealed in a dialysis bag with a cut-off
molecular weight of 3.5 kDa, and dialyzed against water for 3 d.
The solution was lyophilized to obtain the white porous product,
which was stored at 20 °C until use.
The substitution degree of MA was also quantified by measuring the contents of amino groups in gelatin before and after reaction by the Habeeb method using TNBS [43]. Briefly, 0.25 mL
0.01% (w/v) TNBS water solution, 0.25 mL 0.01% gelatin or GelMA
solution, and 0.25 mL 4% NaHCO3 solution were mixed in a centrifuge tube. After being incubated at 37 °C for 2 h, the absorbance
at 420 nm was determined by UV–vis spectroscopy (UV-2550,
Shimadzu, Japan). The concentration of amino groups in gelatin
or GelMA was calculated by referring to a standard curve
generated with a series of glycine solutions with different
concentrations.
2.3. Synthesis and characterization of Aln-grafted GelMA (Aln-GelMA)
Fig. 1. (a) Schematic illustration of the preparation of 6 hydrogels with different Aln
contents and crosslinking degrees based on methacrylated gelatin (GelMA) and
alendronate-grafted GelMA (Aln-GelMA). (b) Schematic illustration of the combinational effect of hydrogel stiffness and Aln density on the differentiation of MSCs.
Alendronate sodium (Aln-NH2) was firstly reacted with excess
glutaraldehyde in water overnight at 45 °C to obtain the aldehyde-modified Aln (Aln-CHO), which was then precipitated and
washed with a large amount of cold acetone. After drying, 60 mg
Aln-CHO was added into 10 mL GelMA PBS solution (100 mg/mL),
and reacted overnight at room temperature. The product was dialyzed against water for 3 d. The solution was lyophilized to obtain
yellow porous Aln-GelMA, which was stored at 20 °C until use.
The chemical structure of the product was characterized by 31P
nuclear magnetic resonance (31P NMR) (500 MHz, Cambridge).
The Aln ratio in the Aln-GelMA was determined similarly by the
aforementioned Habeeb method. Besides, the molybdate blue
method was also used to determine the phosphorus content in
Aln-GelMA [44]. Briefly, the Aln-GelMA was burned in a muffle furnace at 700 °C for 1 h. Residues were dissolved in 0.5 mL 16%
H2SO4, and then mixed with 0.5 mL 2.5% (w/v) ammonium molybdate solution and 0.5 mL 10% (w/v) ascorbic acid solution. After
being incubated at 37 °C for 2 h, the absorbance at 800 nm was
determined by UV–vis spectroscopy. The phosphorus content in
Aln-GelMA was obtained by referring to a standard curve created
with K2HPO4 at the same conditions.
2.4. Hydrogel preparation and characterization
The GelMA macromonomers, APS, and TEMED were mixed to
form a reaction solution. PEGDA was used as the crosslinker to
modulate the cross-linking degree and thereby the stiffness of
hydrogels. Aln-GelMA was added to adjust the Aln concentration
in the hydrogels. The final concentration of each component is
summarized in Table 1. 20 mM APS and TEMED were used as redox
initiators. The reaction lasted for 6 h at 37 °C. The obtained
hydrogels were washed with water to remove unreacted macromonomers and initiators.
The hydrogels were freeze-dried and characterized by FTIR
spectroscopy. The incorporated Aln concentration in the hydrogels
was determined by measuring the phosphorus content as
described above.
The equilibrium swelling ratio of the hydrogels, which is correlated to the cross-linking density of the hydrogel network, was
characterized. The mass of swollen hydrogel (Wwet) was measured
after it was incubated in distilled water for 48 h at room temperature. The swelling ratio (SR) of the hydrogel was determined
according to SR = (Wwet Wdry)/Wdry, where Wdry is the original
weight of the hydrogel.
The mechanical properties of the hydrogels (cylindrical shape,
15 mm in diameter and 5 mm in height) were measured by a
mechanical tester (Instron 5543, USA) in a water tank containing
PBS at 37 °C with a compression rate of 2 mm per min until failure
occurred. The compressive modulus of the hydrogels was obtained
from the linear region of the stress (2–3% strain). Each value was
averaged from 4 parallel measurements.
2.5. Cell isolation and culture
Bone mesenchymal stem cells (BMSCs) were isolated from bone
marrow of Sprague–Dawley rats (6–8 weeks old) according to the
methods reported previously [45]. The procedures were performed
in accordance with the ‘‘Guidelines for Animal Experimentation’’
by the Institutional Animal Care and Use Committee, Zhejiang
University. Briefly, the BMSCs were obtained from the femoral
shafts of rats by flushing out with 10 mL of culture medium (low
glucose Dulbecco’s modified Eagle’s medium, LDMEM) supplemented with 10% fetal bovine serum (FBS, Life Technologies, New
York, USA), 100 lg/mL penicillin and 100 U/mL streptomycin).
The released cells were collected in a 9 cm cell culture dish
(Corning, USA) containing 10 mL culture medium and incubated
in a humidified atmosphere of 95% air and 5% CO2 at 37 °C. After
the cells reached about 80% confluence, they were detached, and
serially subcultured. The BMSCs at passage 2 were used in this
study. The stemness of the BMSCs was verified by differentiation
tests (see Supporting information).
The hydrogel samples were cut into a cylindrical shape with a
diameter of 15 mm, and then were placed into a 24-well plate.
The samples were sterilized in 75% ethanol for 30 min at room
temperature, and followed by four washes in phosphate buffered
saline (PBS) and two washes in culture medium. 15,000 cells in
500 lL culture medium (LDMEM/10% FBS) were seeded onto the
Table 1
Recipe of different hydrogels.
Sample
GelMA
(mg/mL)
Aln-GelMA
(mg/mL)
PEGDA
(mg/mL)
GelMA
Aln-GelMA-1
Aln-GelMA-2
GelMA/PEGDA
Aln-GelMA-1/PEGDA
Aln-GelMA-2/PEGDA
200
200
200
200
200
200
0
0.064
1.28
0
0.008
0.16
0
0
0
20
20
20
3
top of the hydrogels in the 24-well plate, and were incubated for
7 or 21 d. The culture medium (LDMEM/10% FBS) without additional osteogenic supplementation was changed every 3 d.
2.6. Cell morphology and cell number
After culturing on the hydrogels for 21 d, the MSCs were
washed with PBS 3 times, and fixed with 4% formaldehyde solution
for 30 min at room temperature. They were further treated with
0.5% Triton/PBS solution at 4 °C for 10 min. After being washed
with PBS 3 times, they were treated with 1% BSA/PBS solution to
block nonspecific adsorptions for 2 h. The cells were finally stained
with DAPI (100 ng/mL) for nucleus and rhodamine phalloidin solution (0.2 lM, Life Technologies) for cytoskeleton (F-actin) at 37 °C
for 1 h. After washing with PBS 3 times, the cells were observed
under a fluorescence microscope (IX81, Olympus).
After being cultured on the hydrogels for 21 d, the cell number
was quantified by the PicoGreenÒ assay. In brief, the cells were
lysed by repeated freeze–thaw cycles in the presence of 0.2%
Triton X-100. Total double-stranded DNA was quantified after the
samples were incubated with PicoGreen fluorescence dye in the
assigned buffer, and the fluorescence emission intensity at
520 nm was measured according to the manufacturer’s instruction
(PicoGreenÒ dsDNA Quantitation Kit, Invitrogen, USA). The cell
number of each sample was calculated by referring to a standard
curve recorded at the same conditions with known cell number
of MSCs.
2.7. Alkaline phosphatase (ALP) quantification
Phenylphosphate can be hydrolyzed by ALP and form free phenol, which can react with 4-amino-antipyrine in the presence of
alkaline potassium ferricyanide to form a red-colored complex,
whose absorbance at 490 nm is directly proportional to the ALP
activity in the specimen. After being cultured on the hydrogels
for 7 d, the MSCs were washed with PBS 3 times, and treated with
0.5% Triton/PBS at 4 °C for 24 h. After ALP was totally released, the
solution was mixed with reagents from the colorimetric Kit
(KeyGEN Biotech, China) according to the user’s manual. The
absorbance at 490 nm was recorded by a microplate reader and
the activity of ALP is calculated by referring to a standard curve.
The ALP activity per 104 cells was reported.
2.8. Immunofluorescence staining
After being cultured on the hydrogels for 21 d, the MSCs were
washed 3 times with PBS, and fixed in 4% paraformaldehyde for
30 min at 37 °C. After being washed 3 times in PBS, they were further treated in 0.5% (v/v) Triton X-100/PBS at 4 °C for 10 min to
increase the permeability of the cell membrane. After 3 washes
in PBS, they were incubated in 1% BSA/PBS at 37 °C for 30 min to
block the non-specific interactions. Then the cells were incubated
with a mouse monoclonal antibody against collagen type I and
osteocalcin (Abcam, USA) for 1 h, respectively. After being washed
twice in 1% BSA/PBS, they were further stained with fluorescent
labeled IgG (Beyotime, China) and DAPI at room temperature for
1 h, and followed by 3 washes in PBS. The cells were observed
under confocal laser scanning microscopy (CLSM, SP5, Leica,
Germany). The images obtained were further analyzed by Image J
software (National Institutes of Health, USA).
2.9. Western blotting
After culturing on 7 or 21 d on different hydrogels, the MSCs
were washed with PBS three times and completely homogenized
in radio immunoprecipitation assay buffer (RIPA) with protease
4
inhibitors. The lysates were centrifuged at 12,000 rpm at 4 °C for
15 min, and separated on a SDS–PAGE. All the gels have been run
under the same experimental conditions. After being transferred
to a polyvinylidene fluoride (PVDF) membrane (Millipore, MA,
USA), the proteins were incubated overnight with antibodies
(Abcam, USA) and detected using an enhanced chemiluminescence
(ECL Western Blotting Substrate, Pierce, USA) system. The
integral optical density (IOD) was determined using the software
Bandscan 5.0.
2.10. Calcium deposition
The amount of deposited calcium by MSCs was measured by
using the o-cresolphthalein complexone method [46,47]. Firstly,
AMP buffer and staining solution were prepared, respectively.
The AMP buffer was prepared by dissolving 7.06 g AMP in 35 mL
water, whose pH values was adjusted to 10.7 using 6 M HCl, and
the final volume was adjusted to 50 mL by water. The staining
solution was prepared by dissolving 5 mg CPC and 50 mg
8-hydroxyquinoline into 3 mL 12 M HCl solution, whose volume
was adjusted to 50 mL water. After being cultured for 21 d, the
cells were incubated in 0.5 M HCl for 8 h on an ice/water bath
under shaking. After centrifugation at 1000g for 5 min, the
supernatant (5 lL) was mixed with 100 lL staining solution and
100 lL AMP buffer. 5 min later, the absorbance at 570 nm was
determined by a microplate reader (BioRad 680, USA). The calcium
content of each sample was determined by referring to a standard
curve generated by calcium chloride at the same conditions. The
calcium content per 104 cells was reported.
2.11. Statistical analysis
At least three independent experiments were carried out if
not otherwise stated. Results are reported as mean ± standard
deviation, and are analyzed using a paired student’s t-test. The
significant difference level was set at p < 0.05.
3. Results and discussion
3.1. Characterization of gelatin-based macromonomers (GelMA and
Aln-GelMA)
The polymerizable carbon double bonds were grafted to gelatin
molecules to obtain GelMA by the reaction with methacrylic
anhydride (MA) molecules under alkaline environment.
Compared to the 1H NMR spectrum of gelatin, in the spectrum of
GelMA new resonance peaks appeared at 5.60 and 5.36 ppm which
are assigned to the protons of H2C@C(CH3)–, confirming the success of MA grafting (Fig. S1a). The degree of MA substitution
(SDMA) was determined to be 12.0 ± 0.3% and 14.6% by Habeeb
assay and 1H NMR spectroscopy [41], respectively.
Aldehyde molecules were further grafted onto GelMA via coupling between the amine groups of Aln and GelMA [34,48].
Compared with the FTIR spectrum of Aln (Fig. S1b), the absorbance
at 1714, and 2934, 2865 cm 1 appeared, which are assigned to the
stretching vibration of C@O, and the Fermi resonance between
stretching vibration and bending vibration of C–H, respectively.
This result confirms the successful introduction of the aldehyde
group onto Aln molecule. The Aln-CHO was grafted onto the backbone of GelMA by a simple incubation. Although no apparent difference can be found between the 1H NMR spectra of Aln-GelMA
and GelMA (Fig. S1a), the 31P NMR spectra (Fig. 2) reveal that only
the Aln-GelMA had an obvious resonance peak at 17.6 ppm. The
total substitution degree (SDMA+Aln) of Aln-GelMA was determined
to be 22.0 ± 0.4%, implying that the SDAln was about 10%, and the
Fig. 2. 31P NMR spectra of gelatin, methacrylated gelatin (GelMA) and Aln grafted
GelMA (Aln-GelMA).
Aln concentration was 31 lmol in 1 g Aln-GelMA. The Aln content
in the Aln-GelMA was also determined by the molybdate blue
method [44]. The phosphorus element content in 1 g Aln-GelMA
was found to be 58.4 lmol, suggesting that the Aln content in 1 g
Aln-GelMA was 29.2 lmol since one Aln molecule has two phosphonate groups. These results are consistent with each other, and
reveal the successful synthesis of Aln-GelMA.
3.2. Preparation and characterization of hydrogels
In order to maximize the adjustable range of mechanical properties of the hydrogels [49], herein the highest concentration of
gelatin-based macromonomer, i.e. 20%, was used in this study. By
using 20 mM APS and TEMED as initiators, the hydrogels were
formed within 5 min at 37 °C. After gelation, the hydrogels were
kept under 37 °C for another 6 h for complete polymerization.
Demonstrated by Engler et al., hydrogels with low modulus
(0.1–10 kPa) are feasible for neurogenesis, and hydrogels with relatively high modulus (25–40 kPa) are able to promote osteogenesis
[19]. In this study, the stiffness of hydrogels was further modulated
by the addition of 20 wt.% PEGDA, which improved the compressive modulus of the resulting hydrogels from 4 to 40 kPa.
Moreover, the Aln density in the hydrogels was varied by adjusting
the ratio between GelMA and Aln-GelMA. Totally 6 types of hydrogels with variable stiffness and Aln density were prepared as
shown in Table 1.
Compared to the FTIR spectra of GelMA and Aln-GelMA hydrogels (Fig. 3a), in the spectra of the GelMA/PEGDA and Aln-GelMA/
PEGDA hydrogels, new peaks appeared at 1101 and 1730 cm 1,
which are assigned to the C–O asymmetric stretching vibration
of –C–O–C and the C@O stretching of the acetate group respectively. This result suggests the successful incorporation of PEGDA
into the hydrogel. However, due to the rather low concentration
of Aln in the hydrogels, no obvious difference was found in
the FTIR spectra of GelMA/PEGDA, Aln-GelMA-1/PEGDA, and
Aln-GelMA-2/PEGDA hydrogels.
Hydrogel swelling is related with the hydrogel network and
mechanical properties [50,51]. The swelling ratio of the hydrogels
decreased significantly after PEGDA incorporation (Fig. 3b) as a
result of improvement of crosslinking density. By contrast, the
compressive moduli of the hydrogels (Fig. 3c) increased from
4 to 40 kPa. The minor difference in Aln concentration did not
have significant influence on the swelling and mechanical properties of the hydrogels. By variation of the feeding ratio of AlnGelMA, the Aln concentration in the hydrogels was adjusted from
0 to 0.2 ± 0.03 lM (0.2 lM), and further to 3.9 ± 0.1 lM
(4 lM), respectively (Fig. 3d). Therefore, the gelatin-based
5
Fig. 3. (a) FTIR spectra of gelatin-based hydrogels. (b) Swelling ratio, and (c) compressive modulus of different hydrogels. (d) Aln concentration in different hydrogels.
hydrogels with two different stiffness (denoted as 4 and 40 kPa)
and three different densities of Aln (denoted as 0, 0.2 Aln, and
4 Aln) have been successfully prepared.
3.3. Morphology and proliferation of MSCs
After culturing for 21 d on all hydrogels, the MSCs showed similar spindle morphology because of the high cell density, suggesting
all hydrogels had good cytocompatibility (Fig. 4a–f). As shown in
Fig. 4g, in general a similar number of MSCs was found on the
hydrogels without or with a lower Aln concentration, likely due
to the good biocompatibility of the hydrogels. However, about
20% decrease (p < 0.05) of the MSC number was found on the
hydrogels containing the highest density of Aln. This phenomenon
is likely attributed to the differentiation of MSCs which suppresses
cell proliferation, as revealed in the following results.
3.4. Osteogenic differentiation of MSCs
The stemness of obtained MSCs was first verified by differentiation tests. The MSCs were incubated in adipogenic and osteogenic
medium for 14 d, and then stained with Oil red and Alizarin Red S,
respectively. Results (Fig. S2) indicate that the MSCs could undergo
adipogenic differentiation and osteogenic differentiation in the
corresponding differentiation mediums, as evidenced by the lipid
droplet formation and calcium deposition, respectively. In order
to reveal the potential combinational impact of hydrogel stiffness
and Aln density on the MSCs’ multilineage differentiation, the
expressions of myogenic marker MyoD1 and neurogenic marker
b-tubulin were firstly evaluated (Figs. S3 and S4). The MSCs did
not show an obvious expression of MyoD1 and b-tubulin on all
types of hydrogels, indicating they did not undergo myogenic or
neurogenic differentiation under current stimuli. Since the well
spread MSCs are inclined to undergo osteogenesis [34], several
kinds of osteogenesis-related markers were then investigated.
Alkaline phosphatase (ALP) is an important osteogenic marker
in earlier stage of osteogenesis [34]. The expressions of several
differentiation hallmarks at relative later stage were also studied
at protein level. Collagen type I (COL) and osteocalcin (OCN) are
both important for osteogenic differentiation. It has been proved
that COL is the most abundant protein in the organic/inorganic
composite matrix of bone tissue [52]. OCN is the most abundant
noncollagenous protein in the bone matrix, which plays an
essential role in bone formation and remodeling [53]. An
osteogenic tissue is capable of forming an extracellular matrix that
can regulate mineralization, which represents its ultimate
phenotypic expression [54]. Therefore, after being cultured for
21 d, the calcium deposition was measured and used as a hallmark
for osteogenic differentiation [55,56].
After 7 d culture, the expressions of ALP (Figs. 5 and 7a) by the
MSCs were significantly enhanced along with the increase of Aln
concentration in the hydrogels regardless of their stiffness, revealing that the osteo-inductive effect of Aln is dose-dependent
[32,34]. At the same Aln density, the expressions of ALP by the
MSCs were significantly enhanced on the stiffer hydrogels. The
highest expressions of ALP were observed on the stiffest hydrogel
with the highest Aln density (40 kPa/4 Aln), suggesting the synergistical effect of stiffness and Aln on the osteogenesis of MSCs. In
contrast, the expressions of COL and OCN (Fig. 7a) by the MSCs
were only slightly enhanced in the hydrogels with higher Aln concentrations and stiffness. This might be attributed to the fact that
COL and OCN were only actively expressed at the later differentiation stage.
After 21 d culture, the expressions of ALP (Fig. 7b) by the MSCs
on the softer hydrogels regardless of the Aln concentration and the
stiffer hydrogel without Aln were significantly reduced, compared
6
Fig. 4. Cell morphology of MSCs being cultured on (a) 4 kPa, (b) 4 kPa/0.2 Aln, (c) 4 kPa/4 Aln, (d) 40 kPa, (e) 40 kPa/0.2 Aln, (f) 40 kPa/4 Aln hydrogels for 21 d, respectively.
Scale bar = 100 lm. The actin fibers (red) and cell nucleus (blue) were stained with rhodamine phalloidin and DAPI, respectively. (g) Cell numbers quantified by PicoGreen
dsDNA assay. ⁄indicates significant difference at p < 0.05 level.
Fig. 5. ALP activity per 104 cells being cultured on hydrogels for 7 d, respectively.
⁄
and ⁄⁄ indicate significant difference at p < 0.05 and p < 0.01 levels, respectively.
to the counterparts after 7 d culture, respectively. This might be
attributed to the fact that ALP is mainly expressed at the earlier differentiation stage. However, the expressions of ALP by the MSCs on
the stiffer hydrogels with Aln (40 kPa/0.2 Aln and 40 kPa/4 Aln)
were still quite high, suggesting that the synergistical effect of stiffness and Aln can prolong the expression of ALP, and thereby might
be feasible for osteogenesis of MSCs. The expressions of COL and
OCN at 21 d were significantly enhanced than their counterparts
after 7 d culture (Fig. 7), and also were enhanced along with the
increase of Aln concentration in the hydrogels regardless of their
stiffness (Figs. 6 and 7b). Similar to the ALP expressions at 7 d, at
the same Aln density, the expressions of COL and OCN by the
MSCs at 21 d were significantly enhanced on the stiffer hydrogels.
The highest expressions of COL and OCN were observed on the
Fig. 6. (a–f) Immunofluorescence staining of COL in MSCs being cultured on the
hydrogels for 21 d, respectively: (a) 4 kPa, (b) 4 kPa/0.2 Aln, (c) 4 kPa/4 Aln, (d)
40 kPa, (e) 40 kPa/0.2 Aln, and (f) 40 kPa/4 Aln. (g–l) Immunofluorescence staining
of OCN in MSCs being cultured on the hydrogels for 21 d, respectively: (g) 4 kPa, (h)
4 kPa/0.2 Aln, (i) 4 kPa/4 Aln, (j) 40 kPa, (k) 40 kPa/0.2 Aln, and (l) 40 kPa/4 Aln.
Scale bar = 20 lm.
stiffest hydrogel with the highest Aln density (40 kPa/4 Aln), proving again the synergistical effect of stiffness and Aln on the
osteogenesis of MSCs. As shown in Fig. 8, the calcium deposition
7
Fig. 7. Western blotting analysis of ALP, OCN and COL expressed by MSCs on various hydrogels after (a) 7 d and (b) 21 d culture, respectively. The left panel shows the photo
of gels and the right panel shows the relative integral optical density of ALP, COL and OCN calculated from the images by Bandscan software. The protein expression level was
normalized to that of the respective expression of b-actin, which was used as a reference standard. ⁄ and ⁄⁄ indicate significant difference at p < 0.05 and p < 0.01 levels,
respectively. All gels have been run under the same experimental conditions.
by the MSCs was also enhanced along with the increase of Aln
density and hydrogel stiffness.
Besides the synergistical effect of hydrogel stiffness and Aln
density, it might be more important to find out which one is more
effective for osteogenesis between the stiffness and Aln density.
Since the process of mineralization is most significant for dictating
osteogenesis [54], it is used to figure out the importance of these
two factors. Fig. 8 shows that the calcium deposition increased to
1.3–2.6 folds on the stiffer hydrogels compared to the softer ones
with the same Aln density. Meanwhile, the calcium deposition
increased to 1.4 or 2.2 fold along with the increase of Aln density
from 0 to 0.2 lM on the softer and stiffer hydrogels, respectively.
This means the effect of stiffness and low density of Aln (0.2 lM)
is more or less comparable. When the Aln density increased from
0 to 4 lM, the calcium deposition on the softer and stiffer hydrogels increased to 2.0 or 5.0 folds, respectively. This result implies
the osteogenic effect of higher density of Aln (4 lM) is much more
effective than the stiffness (40 kPa). Therefore, one can conclude
that (i) higher stiffness and Aln incorporation can synergistically
Fig. 8. Calcium contents determined by the o-cresolphthalein complexone method
in MSCs being cultured on hydrogels for 21 d, respectively. ⁄ and ⁄⁄ indicate
significant difference at p < 0.05 and p < 0.01 levels, respectively.
8
promote the osteogenesis, and (ii) the osteo-inductive effect of Aln
is more effective than the stiffness if its density is high enough.
Several previous studies have demonstrated the synergetic
effect of different signals on stem cell differentiation. For example,
Kim et al. [17] found that nanotopography and co-culture with
endothelial cells synergistically promote osteogenesis, and the
nanotopography seemed to take a more crucial role. Kaur et al.
[57] has also demonstrated the synergistic effects of nanotopography provided by tobacco mosaic virus and phosphate on
the osteogenic differentiation of MSCs. Zhao et al. [58] built titanium nanotubes of different diameters loaded with strontium.
They found that 10 nm diameter of nanotubes with long-lasting
strontium release showed the best osteogenic properties of
MSCs. However, there are relative less reports to compare the
impact of two competing factors on stem cell differentiation. For
example, Zouani et al. demonstrated that chemical grafting of
BMP-2 mimetic peptide on relative stiff matrices (13–70 kPa)
results in osteogenic differentiation, while the similar grafting on
very soft hydrogel matrices has no effect on the stem cell differentiation [37]. Besides, Banks et al. [59] demonstrated that the
stiffest substrates direct osteogenic lineage commitment of adipose-derived mesenchymal stem cells (ASCs) regardless of the
presence or absence of growth factors (BMP-2 or PDGF), while
softer substrates require biochemical cues to direct cell fate.
Furthermore, Engler et al. [60] demonstrated that differentiation
of MSCs is unaffected by the collagen density tethering on hydrogels with different stiffness. All these results suggest that the
mechanical property of the substrate usually has a stronger impact
on the differentiation of stem cells over chemical cues, such as
extracellular protein and growth factors. However, in this study
we found that the osteo-inductive effect of higher Aln density
(4 lM) is more effective than the high stiffness (40 kPa). This might
be attributed to the different impact of chemical cues, suggesting
more careful case by case study is required to reveal the impact
of stem cell differentiation in a complicated microenvironment.
4. Conclusion
The gelatin-based hydrogels with defined stiffness (4 and
40 kPa) and Aln density (0, 0.2 and 4 lM) were successfully prepared. Enhancing the stiffness and Aln density were found to
improve osteogenesis of mesenchymal stem cells synergistically
in terms of ALP, COL, OCN and calcium expressions. Furthermore,
the osteo-inductive effect of Aln molecules and higher modulus
of the substrate is comparable to some extent. The insight understanding of the interplay between the chemical cues and substrate
stiffness is useful to unveil the differentiation behaviors of MSCs
within a complicated microenvironment containing multivariate
signals, and subsequently guide the design of biomaterials for
controlling stem cell fate.
Acknowledgments
This study is financially supported by the Natural Science
Foundation of China (21434006, 21374097), the National Basic
Research Program of China (2011CB606203), and Open Project of
State Key Laboratory of Supramolecular Structure and Materials
(sklssm201412).
Appendix A. Figures with essential color discrimination
Certain figures in this article, particularly Figs. 1, 4, 6 and 7 are
difficult to interpret in black and white. The full color images can
be found in the on-line version, at http://dx.doi.org/10.1016/
j.actbio.2015.03.018.
Appendix B. Supplementary data
Supplementary data associated with this article can be found, in
the online version, at http://dx.doi.org/10.1016/j.actbio.2015.03.
018.
References
[1] Barry FP, Murphy JM. Mesenchymal stem cells: clinical applications and
biological characterization. Int J Biochem Cell Biol 2004;36:568–84.
[2] Caplan AI, Bruder SP. Mesenchymal stem cells: building blocks for molecular
medicine in the 21st century. Trends Mol Med 2001;7:259–64.
[3] Baksh D, Song L, Tuan RS. Adult mesenchymal stem cells: characterization,
differentiation, and application in cell and gene therapy. J Cell Mol Med
2004;8:301–16.
[4] Deans RJ, Moseley AB. Mesenchymal stem cells. Exp Hematol 2000;28:875–84.
[5] Wang W, Li B, Yang J, Xin L, Li Y, Yin H, et al. The restoration of full-thickness
cartilage defects with BMSCs and TGF-beta 1 loaded PLGA/fibrin gel constructs.
Biomaterials 2010;31:8964–73.
[6] Fang L, Fu X, Sun T, Li J, Cheng B, Yang Y, et al. An experimental study on the
differentiation of bone marrow mesenchymal stem cells into vascular
endothelial cells. Chin J Burn 2003;9:22–8.
[7] Discher DE, Mooney DJ, Zandstra PW. Growth factors, matrices, and forces
combine and control stem cells. Science 2009;324:1673–7.
[8] Reilly GC, Engler AJ. Intrinsic extracellular matrix properties regulate stem cell
differentiation. J Biomech 2010;43:55–62.
[9] Wang W, Li B, Gao C. Modulating the differentiation of BMSCs by surface
properties of biomaterials. Prog Chem 2011;23:2160–8.
[10] Lao L, Zhu Y, Zhang Y, Gao Z, Zhou F, Chen L, et al. Mineralization of collagencoated electrospun poly(lactide-co-glycolide) nanofibrous mesh to enhance
growth and differentiation of osteoblasts and bone marrow mesenchymal
stem cells. Adv Eng Mater 2012;14:B123–37.
[11] Cha C, Liechty WB, Khademhosseini A, Peppas NA. Designing biomaterials to
direct stem cell fate. ACS Nano 2012;6:9353–8.
[12] Marklein RA, Burdick JA. Controlling stem cell fate with material design. Adv
Mater 2010;22:175–89.
[13] Lutolf MP, Gilbert PM, Blau HM. Designing materials to direct stem-cell fate.
Nature 2009;462:433–41.
[14] Fisher OZ, Khademhosseini A, Langer R, Peppas NA. Bioinspired materials for
controlling stem cell fate. Acc Chem Res 2009;43:419–28.
[15] McNamara LE, McMurray RJ, Biggs MJ, Kantawong F, Oreffo RO, Dalby MJ.
Nanotopographical control of stem cell differentiation. J Tissue Eng 2010;2010:
120623.
[16] Saha K, Pollock JF, Schaffer DV, Healy KE. Designing synthetic materials to
control stem cell phenotype. Curr Opin Chem Biol 2007;11:381–7.
[17] Kim J, Kim HN, Lim K-T, Kim Y, Pandey S, Garg P, et al. Synergistic effects of
nanotopography and co-culture with endothelial cells on osteogenesis of
mesenchymal stem cells. Biomaterials 2013;34:7257–68.
[18] Even-Ram S, Artym V, Yamada KM. Matrix control of stem cell fate. Cell
2006;126:645–7.
[19] Engler AJ, Sen S, Sweeney HL, Discher DE. Matrix elasticity directs stem cell
lineage specification. Cell 2006;126:677–89.
[20] McBeath R, Pirone DM, Nelson CM, Bhadriraju K, Chen CS. Cell shape,
cytoskeletal tension, and RhoA regulate stem cell lineage commitment. Dev
Cell 2004;6:483–95.
[21] Kilian KA, Bugarija B, Lahn BT, Mrksich M. Geometric cues for directing the
differentiation of mesenchymal stem cells. Proc Natl Acad Sci USA 2010;107:
4872–7.
[22] Peng R, Yao X, Ding J. Effect of cell anisotropy on differentiation of stem cells
on micropatterned surfaces through the controlled single cell adhesion.
[23] Yao X, Peng R, Ding J. Cell-material interactions revealed via material
techniques of surface patterning. Adv Mater 2013;25:5257–86.
[24] Peng R, Yao X, Cao B, Tang J, Ding J. The effect of culture conditions on the
adipogenic and osteogenic inductions of mesenchymal stem cells on
micropatterned surfaces. Biomaterials 2012;33:6008–19.
[25] Benoit DS, Schwartz MP, Durney AR, Anseth KS. Small functional groups for
controlled differentiation of hydrogel-encapsulated human mesenchymal
stem cells. Nat Mater 2008;7:816–23.
[26] Curran JM, Chen R, Hunt JA. The guidance of human mesenchymal stem cell
differentiation in vitro by controlled modifications to the cell substrate.
[27] Lee JS, Lee JS, Murphy WL. Modular peptides promote human mesenchymal
stem cell differentiation on biomaterial surfaces. Acta Biomater 2010;6:21–8.
[28] Hu Y, Cai K, Luo Z, Zhang Y, Li L, Lai M, et al. Regulation of the differentiation of
mesenchymal stem cells in vitro and osteogenesis in vivo by microenvironmental
modification of titanium alloy surfaces. Biomaterials 2012;33:3515–28.
[29] Zhang Y, Deng X, Scheller EL, Kwon T-G, Lahann J, Franceschi RT, et al. The
effects of Runx2 immobilization on poly (epsilon-caprolactone) on osteoblast
differentiation of bone marrow stromal cells in vitro. Biomaterials 2010;31:
3231–6.
[30] Zhenqing LI, Yanyi XU, Haichang LI, Jianjun G. Immobilization of insulin-like
growth factor-1 onto thermosensitive hydrogels to enhance cardiac progenitor
cell survival and differentiation under ischemic conditions. Sci Chin Chem
2014;57:568–78.
[31] Kilian KA, Mrksich M. Directing stem cell fate by controlling the affinity and
density of ligand–receptor interactions at the biomaterials interface. Angew
Chem 2012;124:4975–9.
[32] Von Knoch F, Jaquiery C, Kowalsky M, Schaeren S, Alabre C, Martin I, et al.
Effects of bisphosphonates on proliferation and osteoblast differentiation of
human bone marrow stromal cells. Biomaterials 2005;26:6941–9.
[33] Fu L, Tang T, Miao Y, Zhang S, Qu Z, Dai K. Stimulation of osteogenic
differentiation and inhibition of adipogenic differentiation in bone marrow
stromal cells by alendronate via ERK and JNK activation. Bone 2008;43:40–7.
[34] Zhu Y, Mao Z, Gao C. Control over the gradient differentiation of rat BMSCs on a
PCL membrane with surface-immobilized alendronate gradient. Biomacromolecules
2013;14:342–9.
[35] Kim S, Yun Y-P, Park K, Kim H-J, Lee D-W, Kim J, et al. The effects of
functionalized titanium with alendronate and bone morphogenic protein-2 for
improving osteoblast activity. Tissue Eng Regen Med 2013;10:353–61.
[36] Jiang X, Cao HQ, Shi LY, Ng SY, Stanton LW, Chew SY. Nanofiber topography
and sustained biochemical signaling enhance human mesenchymal stem cell
neural commitment. Acta Biomater 2012;8:1290–302.
[37] Zouani OF, Kalisky J, Ibarboure E, Durrieu M-C. Effect of BMP-2 from matrices
of different stiffnesses for the modulation of stem cell fate. Biomaterials
2013;34:2157–66.
[38] Hu X, Ma L, Wang C, Gao C. Gelatin hydrogel prepared by photo-initiated
polymerization and loaded with TGF-beta 1 for cartilage tissue engineering.
Macromol Biosci 2009;9:1194–201.
[39] Wu J, Tan H, Li L, Gao C. Covalently immobilized gelatin gradients within
three-dimensional porous scaffolds. Chin Sci Bull 2009;54:3174–80.
[40] Li Y, Lin C, Wang L, Liu Y, Mu X, Ma Y, et al. Maintenance of human embryonic
stem cells on gelatin. Chin Sci Bull 2009;54:4214–20.
[41] Ovsianikov A, Deiwick A, Van Vlierberghe S, Dubruel P, Moller L, Drager G,
et al. Laser fabrication of three-dimensional CAD scaffolds from photosensitive
gelatin for applications in tissue engineering. Biomacromolecules 2011;12:
851–8.
[42] Shin SR, Aghaei-Ghareh-Bolagh B, Dang TT, Topkaya SN, Gao X, Yang SY, et al.
Cell-laden microengineered and mechanically tunable hybrid hydrogels of
gelatin and graphene oxide. Adv Mater 2013;25:6385–91.
[43] Nichol JW, Koshy ST, Bae H, Hwang CM, Yamanlar S, Khademhosseini A. Cellladen microengineered gelatin methacrylate hydrogels. Biomaterials
2010;31:5536–44.
[44] Chen PS, Toribara TY, Warner H. Microdetermination of phosphorus. Anal
Chem 1956;28:1756–8.
[45] Soleimani M, Nadri S. A protocol for isolation and culture of mesenchymal
stem cells from mouse bone marrow. Nat Protoc 2009;4:102–6.
9
[46] Anderson SB, Lin CC, Kuntzler DV, Anseth KS. The performance of human
mesenchymal stem cells encapsulated in cell-degradable polymer-peptide
hydrogels. Biomaterials 2011;32:3564–74.
[47] Smith LA, Liu X, Hu J, Ma PX. The influence of three-dimensional nanofibrous
scaffolds on the osteogenic differentiation of embryonic stem cells.
[48] Wang B, Zhang Y, Mao Z, Gao C. Cellular uptake of covalent poly(allylamine
hydrochloride) microcapsules and its influences on cell functions. Macromol
Biosci 2012;12:1534–45.
[49] Van den Bosch E, Gielens C. Gelatin degradation at elevated temperature. Int J
Biol Macromol 2003;32:129–38.
[50] Hong Y, Song H, Gong Y, Mao Z, Gao C, Shen J. Covalently crosslinked chitosan
hydrogel: properties of in vitro degradation and chondrocyte encapsulation.
Acta Biomater 2007;3:23–31.
[51] Slaughter BV, Khurshid SS, Fisher OZ, Khademhosseini A, Peppas NA.
Hydrogels in regenerative medicine. Adv Mater 2009;21:3307–29.
[52] Blomqvist C, Risteli L, Risteli J, Virkkunen P, Sarna S, Elomaa I. Markers of type I
collagen degradation and synthesis in the monitoring of treatment response in
bone metastases from breast carcinoma. Br J Cancer 1996;73:1074–9.
[53] Nakamura A, Dohi Y, Akahane M, Ohgushi H, Nakajima H, Funaoka H, et al.
Osteocalcin secretion as an early marker of in vitro osteogenic differentiation
of rat mesenchymal stem cells. Tissue Eng C Methods 2009;15:169–80.
[54] Moeinzadeh S, Barati D, He X, Jabbari E. Gelation characteristics and
osteogenic differentiation of stromal cells in inert hydrolytically degradable
micellar polyethylene glycol hydrogels. Biomacromolecules 2012;13:2073–86.
[55] Stern J, Lewis WHP. The colorimetric estimation of calcium in serum with ocresolphthalein complexone. Clin Chim Acta 1957;2:576–80.
[56] Hayden RS, Fortin J-P, Harwood B, Subramanian B, Quinn KP, Georgakoudi I,
et al. Cell-tethered ligands modulate bone remodeling by osteoblasts and
osteoclasts. Adv Funct Mater 2013;24:472–9.
[57] Kaur G, Wang C, Sun J, Wang Q. The synergistic effects of multivalent ligand
display and nanotopography on osteogenic differentiation of rat bone marrow
stem cells. Biomaterials 2010;31:5813–24.
[58] Zhao L, Wang H, Huo K, Zhang X, Wang W, Zhang Y, et al. The osteogenic
activity of strontium loaded titania nanotube arrays on titanium substrates.
[59] Banks JM, Mozdzen LC, Harley BaC, Bailey RC. The combined effects of matrix
stiffness and growth factor immobilization on the bioactivity and differentiation
capabilities of adipose-derived stem cells. Biomaterials 2014;35:8951–9.
[60] Wen JH, Vincent LG, Fuhrmann A, Choi YS, Hribar KC, Taylor-Weiner H, et al.
Interplay of matrix stiffness and protein tethering in stem cell differentiation.
Nat Mater 2014;13:979–87.

Combinational effect of matrix elasticity and alendronate density on

Transcription

Similar documents

“The 79th of Stem Cell Biology and Regenerative Medicine Forum”

Engineering Process in the Classroom Event

ALN2015 Brochure - African Leadership Network

Globaloria sparks students` imaginations as they learn to create and

Hydrogels

4. Biomaterials and Tissue Engineering 4.2.1. The use of

Multi-Responsive and Tough Hydrogels Based on Triblock

Study Guide - Swift Classroom

EQT 102 ENGINEERING MATHEMATICS II (MATEMATIK

Attention Educators: