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Dynamic Article LinksC<Soft Matter
Cite this: Soft Matter, 2011, 7, 6501
www.rsc.org/softmatter PAPER
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View Article Online / Journal Homepage / Table of Contents for this issue
Substrate induced differentiation of human mesenchymal stem cells onhydrogels with modified surface chemistry and controlled modulus†
Mathieu Lanniel,a Ejaz Huq,b Stephanie Allen,a Lee Buttery,c Philip M. Williamsa and Morgan R. Alexander*a
Received 31st January 2011, Accepted 21st April 2011
DOI: 10.1039/c1sm05167a
Polyacrylamide hydrogels were prepared with variable stiffness within a range of effective surface
Young’s modulus values from 5.5 kPa to 152 kPa measured in the hydrated state using atomic force
microscopy (AFM). The gel surface was modified with either collagen or plasma polymer coatings
containing amino, carboxyl or phosphate moieties. Analysis of the surface chemistry using X-ray
photoelectron spectroscopy and AFM indentation showed that the coated gels present very different
surface chemistries while maintaining the range of stiffness. The density of human mesenchymal stem
cells (hMSC) adhered to the materials was found to depend on the surface chemistry, with the highest
cell densities achieved for collagen coated gels. The spread of each cell was shown to be greater for the
stiffer surfaces independent of surface chemistry. To assess the differentiation of the hMSCs, antibody
staining was carried out using markers for osteogenic (Runx2), myogenic (MyoD1) and neurogenic (b-
III tubulin) cell types which revealed a dependence of marker protein expression upon both surface
stiffness and chemistry. The expression of the osteogenic Runx2 marker was maximal for cells cultured
on gels of 41 kPa stiffness when modified with the phosphate plasma polymer. Myogenic MyoD1
expression was maximal on the carboxyl coated gels of intermediate stiffness (10 kPa to 17 kPa).
Neurogenic differentiation indicated by b-III tubulin expression was seen to be greatest on the carboxyl
surfaces and for the lowest surface stiffness substrates. Using soluble factors in the medium to induce
osteogenic behaviour resulted in the formation of bone nodules and matrix calcification for gel stiffness
values higher than 10 kPa, especially on amino-functionalized coatings but not for collagen coated gels.
The results indicate that control over differentiation fate of hMSCs can be exerted using not only
surface stiffness, a result previously widely reported, but also surface chemistry working in tandem with
the influence of compliance. This has great significance in developing stem cell therapies when synthetic
surfaces are used as scaffolds, delivery vehicles or culture ware.
1. Introduction
Stem cell therapy shows promise in the treatment of many
human diseases including myocardial infarction, neurological
disorders such as stroke and the replacement of damaged and
diseased bone.1–5
Surfaces are a key area in controlling cells as the interactions of
cells with the natural extracellular matrix (ECM) in the human
body and with artificial surfaces in biotechnology are essential to
aLaboratory of Biophysics and Surface Analysis, School of Pharmacy,University of Nottingham, University Park, Nottingham, NG7 2RD, UK.E-mail: [email protected]; Fax: +44 (0)115 9515102; Tel: +44 (0)115 951 5119bRutherford Appleton Laboratory, Harwell Science and InnovationCampus, Didcot, OX11 0QX, UKcCentre for Biomolecular Sciences, School of Pharmacy, University ofNottingham, University Park, Nottingham, NG7 2RD, UK
† Electronic supplementary information (ESI) avaiable. See DOI:10.1039/c1sm05167a
This journal is ª The Royal Society of Chemistry 2011
many cellular functions including survival, proliferation and
differentiation. Cell adhesion to surfaces occurs as a result of
signalling via transmembrane proteins, such as integrins, that
bind to specific motifs within ECM proteins in the body or those
adsorbed to synthetic materials from serum. The binding of
integrins to the ECM causes them to cluster and leads to the
recruitment of cytoplasmic factors, which initiates intracellular
signalling cascades in response to ECM binding.6 These signals
from the ECM, combined with soluble factors and cell–cell sig-
nalling, determine cell fate.
Mesenchymal stem cells (MSC) are attractive for regenerative
medicine because they can be easily expanded in vitro while
maintaining pluripotency and have the capacity to differentiate
along various lineages of the skeletal connective tissues: osteo-
blasts, chondrocytes, bone marrow adipocytes and hematopoi-
esis-supportive stromal cells.7–11 Certain observations suggest
that they can also give rise to skeletal muscle cells and may even
be capable of forming non-mesodermic cells such as neurons or
astrocytes.12,13
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A number of studies using different types of cells have iden-
tified matrix properties with major effects on cell behaviour.
Nanoscale roughness of the culture surface has been found to
improve both cell adhesion and proliferation14,15 and to direct
hMSCs towards the osteogenic lineage with a similar efficiency to
osteogenic medium supplements.16 Control of stem cell shape by
confinement on chemically patterned surfaces has also been
shown to influence stem cell differentiation.17 Surface chemistry
has also been shown to be an important factor in hMSC differ-
entiation. Functionalisation of the bulk chemistry of gels with
low concentrations of monomethacrylated monomers containing
either an amino, a tert-butyl, a phosphate, a fluoro or a carboxy-
lic group in polyethylene glycol has been shown to have signifi-
cant effect upon hMSC fate.18 Markers of hMSC differentiation
and staining of extracellular matrix showed that phosphate
containing groups promoted osteogenesis, tert-butyl groups
promoted adipogenesis and carboxylic groups promoted chon-
drogenesis. It was proposed that encapsulation of the cells in
three dimensional PEG environment allowed deconvolution of
the effects of surface chemistry and cell spreading by keeping the
cell area constant. This allowed the workers to conclude that the
effect on cell differentiation was caused only by the chemical
groups present on the gel surface. Interestingly, the chemical
functionalities able to direct differentiation of the stem cells are
representative of the chemical environment encountered by the
cells of these lineages. Phosphate groups have been shown to
have a role in bone formation19 whereas tert-butyl groups are
similar to the hydrophobic groups present in the lipid based
extracellular matrix of adipose cells.20
A number of studies have highlighted the importance of
substrate stiffness on the response of mature cells to synthetic
surfaces.6,21 Cells respond to matrix rigidity by sampling their
environment through integrin focal adhesion complexes linking the
ECM to the cytoskeleton, a process known as mechano-
transduction. Engler et al.22 investigated the effect of matrix stiff-
ness on hMSC differentiation where these cells were cultured on
collagen I coated polyacrylamide gels for which a varying rigidity
had been achieved by controlling the extent of chemical cross-
linking. Their striking discovery was that, using the same culture
medium in each case, on substrates with low (0.1 kPa to 1 kPa),
medium (8 kPa to 17 kPa) and high (25 kPa to 40 kPa) surface
stiffness, hMSCs started to show a specification towards the
neurogenic, myogenic and osteogenic lineages, respectively. This
observation was based on cell morphology and expression of early
differentiation markers. Expression of terminal differentiation
markers was limited, showing the importance of complementary
factors, such as soluble growth factors, to achieve a complete
differentiation. In a later study, Rowlands et al.23 used the same gel
systemasEngler et al.but tested four differentECMproteins cross-
linked on the gels. They observed an interplay between surface
stiffness and the identity of the pre-attached proteins on hMSC
differentiation. This shows that hMSC lineage specification
requires the right matrix composition as well as the appropriate
mechanical properties from the matrix. These two studies used
proteins crosslinked to the gel surfacebutBenoit et al.18andCurran
et al.24 have shown that even simple chemical groups deposited on
the culture surface could influence stem cell fate. However, these
studies didnot test the combined effect of different surface chemical
functionalities and variable surface stiffness.
6502 | Soft Matter, 2011, 7, 6501–6514
Surface modification to improve cell attachment can be ach-
ieved by deposition of a thin organic coating resulting from
plasma polymerization.25 The introduction of amine groups
using plasma treatment has been shown to direct hMSCs towards
the osteogenic lineage.26 Plasma polymerised allylamine
(ppAAm) deposition is a way of creating surfaces with amino
surfaces, which is largely independent of the substrate chemistry
and has previously been shown to improve cell attachment.27–31
Surfaces containing carboxylic groups have also been found to
influence hMSC lineage specification and can be obtained by
surface coating with plasma polymerised acrylic acid (ppAAc).32
Finally, surfaces containing phosphate groups have been found
to increase expression of osteogenic markers in hMSCs.18 Lin
et al. observed that plasma polymerized trimethyl phosphite
(ppTMP) deposition was a means to introduce phosphate groups
into surfaces.33
In this work, we combine the control of surface stiffness using
variably compliant polyacrylamide gels with a control of surface
chemistry using plasma polymer deposition. The effect of these
coatings on hMSC behaviour is also compared to the effect of
collagen coatings. The aim of this study is to investigate the
interplay between these stiffness and surface chemistry factors on
hMSC adhesion and differentiation. Correlations between
chemistry and cell area, cell number and phenotype were iden-
tified. Furthermore, synthetic chemistry–stiffness combinations
superior to collagen coating in guiding the formation of bone
nodules were also identified.
2. Experimental section
2.1. Substrates
2.1.1. Polyacrylamide gel preparation. All chemicals were
purchased from Sigma-Aldrich unless otherwise stated. Variably
compliant polyacrylamide gels were prepared following
a method of Pelham and Wang.34 Coverglasses were soaked in
0.1 M NaOH and air dried. A small aliquot of 3-amino-
propyltrimethoxysilane was spread on the glass surface. After 5
minutes, the coverglasses were washed and soaked in distilled
H2O. The coverglasses were then immersed for 30 minutes in
a 0.5 vol% solution of glutaraldehyde in PBS. After this, they
were washed extensively in distilled water and air dried. N,N0
Methylene bis acrylamide was added to a 10 wt% acrylamide
solution in distilled water to achieve final bis acrylamide
concentrations ranging from 0.01 wt% to 0.6 wt%. The cross-
linking was induced by addition of 0.5 vol% of 10 wt% ammo-
nium persulfate solution (Fisher Scientific) and 0.05 vol% of N,
N,N0,N0-tetramethylethylenediamine. An accurately dispensed
volume of 50 mL of the polymerizing solution was added on each
coverglass and coverslips were placed on top. After polymeri-
zation, the gels were rinsed with 200 mM HEPES. Based on the
volume of gel deposited, the coverslip area (22 mm � 22 mm)
and the fact that no gel was lost and no air was added in the gel,
the gel thickness was estimated to be 100 mm.
2.1.2. Plasma polymerization. Plasma polymerisation was
carried out in a T-shaped borosilicate chamber. Two external
inductively coupled copper band electrodes that were connected
to a 13.56 MHz radio frequency power source (Coaxial Power
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System Ltd.) were used to initiate the plasma. The power was
matched manually such that the reflected power was lower than
1 W. The gas pressure was monitored via a Pirani gauge and
controlled by needle valves (BOC Edwards). All monomers
(allylamine, acrylic acid, trimethyl phosphite) were obtained
from Sigma-Aldrich and were degassed by at least two freeze–
thaw cycles prior to use. After 30 seconds of surface activation
using oxygen plasma (20 W, 40 Pa), the plasma polymers were
deposited on the dry polyacrylamide gels. The peak deposition
power was 20 W with a 10 ms cycle within which the plasma was
on during 2 ms and off for 8 ms. The duty cycle of the pulsed
deposition was therefore 0.2 (the duty cycle is defined as g ¼ ton/
(ton + toff) where ton is the time where the plasma is switched on
and toff when it is switched off). The equivalent power of the
deposition (Peq) is defined as Peq ¼ Ppeak � g, in this experiment
Peq was 4 W. The deposition was carried out until a fixed amount
of polymer was deposited, according to a quartz crystal sensor.
2.1.3. Collagen type I coating. Rat tail collagen type I was
covalently bound to the gel surfaces using sulfo-SANPAH
(Pierce, Rockford, IL). A thin layer of 1 mg mL�1 sulfo-SAN-
PAH solution was placed on top of the gels, which were then
exposed to ultraviolet light for ten minutes to bind the sulfo-
SANPAH to the surface of the gels. Collagen binding to the gels
was then carried out for eighteen hours at 4 �C. Following
collagen cross-linking, the gels were rinsed with PBS prior to cell
seeding.
2.2. Surface characterization
2.2.1. AFM imaging and nanoindentation. Images of the
surface topography were obtained on a Dimension 3000 AFM
(Digital Instruments, Veeco) controlled with NanoScope (V5.30,
2005). Dry gels were used for imaging as hydrated gels were too
soft to be imaged. The samples were imaged using tapping mode
and the acquired micrographs had a dimension of 10 mm �10 mm. The images were processed with the Scanning Probe
Image Processor (SPIP, Version 3.3.6.5, 2005, Image
Metrology). The route mean square (RMS) surface roughness
was calculated on 1 � 1 mm areas after line wise plane correc-
tions. The data presented are an average of five measurements.
All errors shown represent standard deviations.
AFM nanoindentation was used to measure the elasticity of
hydrated polyacrylamide gels before and after surface coating. A
Molecular Force Probe-1D (Asylum Research, Santa Barbara,
CA, USA) and DNP-S probes (Veeco) were used. The spring
constant of the cantilever was 0.13 N m�1, as calculated using the
thermal method.35 The deflection sensitivity was measured using
a reference silicon wafer. The loading force applied on the
surfaces was kept constant at 30 nN and was within the linearity
domain of the photodetector. In order to determine the model of
indenter shape that had to be used for the Young’s modulus
calculation, the tip geometry was estimated using a silicon TGT1
grating sample (NT-MDT, Moscow, Russia), consisting of
a matrix of sharp spikes. The tip height image obtained was then
analyzed using a tip apex geometry calculation macro. This
macro calculates the projected radius, the spherical radius, the
cone half angle corresponding to the three models of tip geom-
etry, as well as the projected area of the tip. The force curves
This journal is ª The Royal Society of Chemistry 2011
obtained were then analysed using a force curve analysis macro,
in which the effective Young’s modulus values of the poly-
acrylamide gels were obtained using the linear portion of the
retraction region of the force curves. Polyacrylamide has been
previously found to have a Poisson ratio of 0.48,36 therefore this
value was used to convert reduced modulus values into effective
surface Young’s modulus values.
2.2.2. Water contact angle (WCA) measurements. The static
WCA of the samples were measured with a CAM 200 Optical
Contact Angle Meter (KSV Instruments LTD). Immediately
after bringing the surface in contact with the water droplet,
fifteen images were taken at one second intervals. The average
contact angle of each image was calculated using a Young
Laplace fit and extrapolated back to the point at which the drop
contacted the surface.
2.2.3 X-Ray photoelectron spectroscopy. XPS analysis was
carried out on a Kratos AXIS ULTRA instrument with a mon-
ochromated Al Ka X-ray source (1486.6 eV) operated at 15 mA
emission current and 10 kV anode potential. Electron flood was
used for charge neutralization. The pass energy was 80 eV for
survey scans with a step size of 1 eV and 20 eV for high resolution
scans, which had a step size of 0.1 eV. Photoelectrons emitted
normal to the sample surface were analysed. Empirically derived
sensitivity factors provided by the manufacturer were used to
achieve quantification of the elemental composition in atomic
percent from survey scan spectra. The C 1s envelope from the
different coatings was fitted with component peaks of equal full
width half maximum (FWHM). The component corresponding
to the C–C environment was placed at a binding energy of
285.0 eV to account for charging at the surface. In order to
estimate the thickness of the coatings, the Beer–Lambert
approach was used. This approach gives the coating thickness
based on the attenuation of the signal from an element present in
the substrate and not in the coating using the following formula:
t¼�l* ln (I/I0), where t is the coating thickness, l is the inelastic
mean free path of the core level studied, I is the peak intensity of
this element obtained on the coated sample and I0 is the peak
intensity of this element for the uncoated sample. For estimation
of ppAAm coating thickness, the Si 2p signal was used
(l ¼ 3.1 nm),37 whereas for ppAAc and ppTMP coatings, the
attenuation of the N 1s signal was used (l ¼ 2.9 nm).
2.3. Cell culture
hMSCs (TCS cellworks Ltd) at passage number five derived from
the bone marrow were expanded on gelatine in Mesenchymal
Stem Cell Medium (MSCM, TCS cellworks Ltd). All cells were
plated on the coated polyacrylamide gels, which were not
exposed to a preliminary incubation in the medium, at 1000 cells
cm�3 and were cultured for fourteen days in low-glucose Dul-
becco’s modified Eagle’s medium (DMEM) supplemented with
20 vol% fetal bovine serum (FBS), 50 mg mL�1 streptomycin, and
50 units mL�1 penicillin. Medium was changed every two days.
Uncoated and oxygen plasma treated polyacrylamide gels did
not support cell adhesion. For osteogenic induction from the
medium, cells were cultured for fourteen days in low-glucose
DMEM supplemented with 20 vol% fetal bovine serum (FBS),
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50 mg mL�1 streptomycin, 50 units mL�1 penicillin, 50 mM
ascorbate 2 phosphate, 10 mM b-glycerol phosphate and 100 nM
dexamethasone. Changes in cell shape and cell area were
observed under an optical microscope after plating on the
polyacrylamide gels.
2.4 Immunocytochemistry study
After fourteen days of culture, cells were fixed using 4 wt%
paraformaldehyde in PBS pH 7.4 for ten minutes at room
temperature. The samples were then washed twice using ice cold
PBS. In order to permeabilize the cells, the samples were incu-
bated for ten minutes with PBS containing 0.25 vol% Triton
X-100 and washed three times in PBS for five minutes. Samples
were then incubated in blocking buffer (PBS + 1 wt% BSA +
0.3 M glycine) for thirty minutes. The cells were then incubated
with the diluted antibody (10 mg mL�1) against either Runx2,
MyoD1 or b-III tubulin (Abcam) in 1 wt% BSA in PBS in
a humidified chamber for one hour at room temperature. The
cells were then washed three times for five minutes with PBS and
incubated with the diluted secondary antibody (Sigma-Aldrich)
in 1 wt% BSA for one hour at room temperature in the dark.
After incubation, the samples were rinsed three times for five
minutes with PBS in the dark and stained with DAPI nuclear
staining (Invitrogen). Negative controls were included, where the
samples were incubated only with the secondary antibodies to
test for non-specific staining. These controls did not show
significant staining (ESI, Fig. S1†). To study the effect of surface
stiffness on the cytoskeleton and formation of stress fibers, cells
from ppAAm coated gels were also stained with FITC-Phalloidin
(Sigma-Aldrich), following the same procedure as for antibody
staining.
2.5. Confocal microscope analysis
Fluorescently labelled cells were examined under a Leica TCS SP
confocal microscope (Leica Microsystems GmbH) and the rela-
tive fluorescence intensity was measured using the same gain
settings and pinhole size for all the samples of each marker
studied. Stacks of images were used in order to include the whole
cell volume in the analysis.
Fig. 1 Effective Young’s modulus values obtained using AFM nano-
indentation on polyacrylamide gels in water, without coating (�) or
2.6. ECM staining
Alizarin red S was prepared by dissolving 1 g of the dye (Fisher
Scientific) in 100 mL deionised water. The solution was filtered
with 0.45mm filter and pH adjusted to 4.2 using ammonium
hydroxide. The polyacrylamide gels with the fixed hMSCs were
washed three times with deionised water before 0.5 mL Alizarin
red S solution was added to each gel for five minutes. The stain
was then removed and the gels were washed with deionised water
for five times until the water was clear. ECM staining was then
observed under a light microscope.
coated with ppAAm (>), ppAAc (,), ppTMP (O) and collagen type I(B). Error bars represent standard deviations. One way ANOVA testing
indicated that, for each bis acrylamide concentration studied, the
measured Young’s modulus values were significantly affected by the type
of coating (‘‘**’’ indicate that at least one of the means is different from
the others at the 0.01 level).
2.7. Statistical analysis
All statistical analysis (ANOVA, Bonferroni post hoc tests) was
performed using the Origin software (Originlab).
6504 | Soft Matter, 2011, 7, 6501–6514
3. Results and discussion
Neither the surface properties of polyacrylamide gels formed
using the present method nor the properties of these materials
coated with collagen, ppAAm, ppAAc or ppTMP have been
published previously. Therefore, surface chemical, topographical
and mechanical characterisations were carried out in order to
determine the physicochemical nature of the surfaces that may
influence the cell response. The effect of coating the gel on
surface topography and Young’s modulus were assessed using
AFM, while surface chemistry was studied using XPS analysis
and sessile drop WCA measurements.
3.1 Surface modulus characterisation
AFM nanoindentation was carried out on uncoated and coated
gels in PBS solution in order to determine the stiffness of the
outermost surface. Examples of force versus piezo displacement
curves obtained on uncoated polyacrylamide gels are shown in
the ESI (Fig. S2†). The effective Young’s moduli measured are
shown in Fig. 1 for all coatings and cross-linker concentrations.
An increase in modulus with increasing bis acrylamide concen-
tration was observed for all coatings. The modulus was found to
be dependent upon the type of coating used as indicated by one
way ANOVA testing performed for each bis acrylamide
concentration (ESI, Table S1†). The stiffness values achieved on
uncoated gels ranged from 7 kPa to 152 kPa. For the four surface
coatings tested in the present study, the Young’s modulus of the
stiffest gels dropped after surface coating from 152 kPa to values
comprised between 75 kPa and 90 kPa, depending on the coating.
This suggests that the deposited layer of collagen or plasma
polymer contributed to the measured modulus and was softer
than the stiffest gels.
This journal is ª The Royal Society of Chemistry 2011
Fig. 2 (A) AFM images (10 mm � 10 mm) of the coated and uncoated
polyacrylamide gels (Z-scale is 50 nm from black to white). Scale bar is
2 mm. (B) AFM image (2 mm� 2 mm) of a collagen coated polyacrylamide
gel containing 0.01% bis acrylamide (Z-scale is 20 nm from black to
white). Scale bar is 0.5 mm.
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After coating of the polyacrylamide gels, the stiffness values
achieved ranged from about 6.5 kPa to 90 kPa for all coating
chemistries tested. This wide range covers the stiffness values
previously shown to induce osteogenesis and myogenesis in
hMSCs.22 Gels with lower cross-linking contents aiming to
achieve lower stiffness values resulted in tacky surfaces and were
therefore not included in the study. For all nanoindentation
measurements, a load of 30 nN was applied on the gel surface
resulting in an indentation depth from 4 mm for the softest gels to
0.2 mm for the hardest gels. This technique therefore measures
the mechanical properties of the outermost surface and not of the
bulk material. As this indentation depth is larger than the
average coating thickness on the gel (between 0.5 nm and 30 nm
based on XPS and ellipsometry measurements), the measured
stiffness is clearly a combination of both the surface coating and
the gel. Previously, we have shown that the deposition of a soft
plasma polymer coating on a hard silicon surface greatly reduces
the measured surface Young’s modulus values (70 GPa to 1 GPa)
until the substrate is no longer sensed at a coating thickness of
approximately 150 nm.38 The indentation depth in the present
study is much smaller than 15% of the total gel thickness (about
100 mm), so the measured effective Young’s modulus is unlikely
to be influenced by the underlying glass coverslip.39 Buxboim
et al. studied the behaviour of hMSCs cultured on poly-
acrylamide gels of various thicknesses on hard substrates. They
showed that the cells started to be affected by the hard substrates
for gels thinner than 20 mm.40 As the gels used in the present
study are much thicker, the cells are unlikely to be influenced by
the stiffness of the supporting glass surface. Studies investigating
the range of forces applied by cells on their substrates found
varying values depending on cell types. Human fibroblasts
typically exerts forces of about 20 nN per focal adhesion,41
whereas myocytes exerts higher forces up to 70 nN.42 The load
applied to the surface in this study during AFM nanoindention
(30 nN) is in a typical range of cell induced stress and strain so the
measured surface stiffness is relevant to the one felt by cells
applying forces to the surface.
3.2. Topographical characterisation
AFM images were obtained for all gel and coating conditions in
order to study how the cross-linking degree of the poly-
acrylamide gels as well as the plasma polymer and collagen
coatings influence the surface topography. Surface roughness
values were then calculated for each condition using 10 mm �10 mm images. Examples of AFM images are shown in Fig. 2 for
the five cross-linker contents of polyacrylamide gels before and
after coating. These are all shown with the same height range for
ease of comparison. Roughness values calculated from these
images are shown in Table 1.
Roughness values measured from the AFM images were low
(sub-nm), but showed significant differences between coated gel
categories, for both surface stiffness and surface chemistry
(p-value < 0.01, ESI, Tables S2 and S3†). However, no systematic
variation of surface roughness depending on bis acrylamide
concentration was identified (Fig. S3, ESI†). Uncoated poly-
acrylamide gels were found to have the lowest roughness values
and showed a smooth surface at the nanometre scale. Coating
with ppAAm, ppAAc and ppTMP did not introduce any
This journal is ª The Royal Society of Chemistry 2011
discernible features, although an increase in roughness was
observed following coating. Collagen coated gels showed
a characteristic nanofibrillar structure with the highest RMS
roughness values due to the presence of collagen fibers on the
surfaces, visible on the AFM images for all gel compositions. The
AFM images confirm that the collagen layer coated on the gels in
continuous over the surface. In summary, all the plasma polymer
coated samples showed very low roughness values, comprised
between 0.15 nm and 0.32 nm and surface roughness was not
greatly influenced by the gel cross-linking degree. The absence of
trend in surface roughness depending on bis acrylamide
concentration suggests that the increase in effective Young’s
modulus depending on gel bis acrylamide concentration shown
in Fig. 1 is not caused by variation in surface roughness.
Engineered nanostructures (pits 100 nm deep and 120 nm
diameter in PMMA) have been shown to direct hMSCs towards
the osteogenic lineage without the use of supplements in the
culture medium.16 An increased expression of osteopontin and
osteocalcin after twenty one days of culture was observed as well
as the presence of early bone nodule formation and matrix
mineralization. In the present study, the range of roughness and
the size of the features (<10 nm) observed is much lower than the
range previously shown to influence hMSC differentiation.16,43
Soft Matter, 2011, 7, 6501–6514 | 6505
Table 1 Roughness values calculated from the AFM images obtained for each surface coating and gel composition
Uncoated Young’s modulus /kPa 5.5 14 33.5 63 152 Overall roughness value/nmRoughness/nm 0.19 � 0.02 0.15 � 0.02 0.22 � 0.06 0.14 � 0.03 0.15 � 0.02 0.17 � 0.04
ppAAm Young’s modulus/kPa 8.5 10 27 53 90 Overall roughness value/nmRoughness/nm 0.22 � 0.02 0.23 � 0.02 0.16 � 0.01 0.26 � 0.04 0.26 � 0.05 0.22 � 0.045
ppAAc Young’s modulus/kPa 6.5 10 17 35 75 Overall roughness value/nmRoughness/nm 0.32 � 0.04 0.20 � 0.02 0.25 � 0.02 0.25 � 0.01 0.30 � 0.06 0.26 � 0.05
ppTMP Young’s modulus/kPa 7 14 23 41 79 Overall roughness value/nmRoughness/nm 0.25 � 0.02 0.27 � 0.01 0.28 � 0.01 0.28 � 0.01 0.27 � 0.02 0.27 � 0.02
Collagen type I Young’s modulus/kPa 7 24 38 63 85 Overall roughness value/nmRoughness/nm 0.45 � 0.02 0.34 � 0.02 0.42 � 0.04 0.33 � 0.03 0.35 � 0.02 0.38 � 0.05
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3.3. WCA measurements
The chemistry of the uppermost nanometre of the surface
controls both the WCA and cellular response. Consequently, this
measurement method is widely used to investigate biomaterials.
Uncoated polyacrylamide gels were found to have an initial
water contact angle value of 31�. Oxygen etching of the gels
caused a decrease in contact angle value to 17�. Collagen coated
gels had the highest initial contact angle value at 85�. For
ppAAm, ppAAc and ppTMP, plasma deposition lead to
a decrease in contact angle compared to uncoated gels, with
initial water contact angle values of 25� for ppAAm and ppTMP
and 19� for ppAAc coatings.
3.4. Chemical characterisation of plasma polymer modified gel
surfaces
To determine the chemical composition of the coatings, XPS
analysis was carried out. The survey scans obtained for all the
coating conditions are shown in Fig. 3 and the corresponding
atomic concentrations of the elements detected in the wide scans
are shown in Table 2.
On the wide scan for the uncoated polyacrylamide gels, apart
from the three main peaks corresponding to O 1s, N 1s and C 1s
consistent with the polyacrylamide composition, two smaller
peaks are observed at 102 eV and 153 eV binding energy, cor-
responding to Si 2p and Si 2s, respectively. This indicates the
presence of a small amount of silicon contamination of the
polyacrylamide gels, probably in the form of silicone. As silicon
is not present in the monomers used for plasma deposition, the
Fig. 3 XPS wide scans for an uncoated polyacrylamide gel and gels
coated with collagen type I, ppAAm, ppAAc or ppTMP.
6506 | Soft Matter, 2011, 7, 6501–6514
attenuation of the Si 2p signal following ppAAm coating was
used to estimate the thickness of the ppAAm coating. For
ppAAc and ppTMP coatings, attenuation of the N 1s signal from
the underlying polyacrylamide gel was used to obtain a thickness
estimate. The observation of a Si 2p peak in the ppAAm coating
(0.2% in atomic concentration, shown in Table 1) and a N 1s
peak in the ppAAc (6.2% in atomic concentration) and ppTMP
coatings (10.5% in atomic concentration) shows that these
coatings are thinner than the analysis depth of XPS. For the
collagen coating, no Si 2p signal was observed suggesting
the coating is thicker than the analysis depth of XPS and the
obtained chemistry represents the collagen coating only.
In order to quantify the functional groups at the surface, the C
1s peak was modelled using the minimum number of synthetic
components required to fit the raw data. The number and shape
of the components are shown in Fig. 4 and their positions and
functionalities assigned on the basis of knowledge of the sample
and comparison with standard compounds reported in the
literature44 are shown in Table 3.
All data have been charge corrected to place the most intense
component of the C 1s core level at 285.0 eV, which was attrib-
uted to the hydrocarbon bond (C–C or C–H). For the uncoated
polyacrylamide gel, one additional component was observed at
286.4 eV, representing 15.3% of the total C 1s peak and was
attributed to N–C–N bonds present in the polyacrylamide
structure. Another component was fitted at 288.2 eV, repre-
senting 10.3% of the total C 1s signal and was attributed to the
carbon involved in the amide bond (C(]O)N).
For the collagen coating, the C 1s spectrum also showed high
binding energy components. Component III at 288.1 eV was
attributed to the carbon involved in the amide environment (C
(]O)N–C) and component II at 286.3 eV was assigned to the
other carbon in the peptide bond (C–C(]O)N).
For the ppAAm coating, thickness estimation using the
attenuation of the Si 2p signal from the underlying gel gave
a value of 13 nm. Analysis of the survey scans showed an increase
in nitrogen content following plasma coating. For the C 1s peak,
a component representing 35.4% of the total C 1s signal was
fitted at 286.2 eV binding energy (component II) and was
attributed to C–N bonds. Component III, representing 9% of the
peak, was fitted at 288.0 eV, which is likely to correspond to
C]O groups formed by oxidation of the coating. These results
are in general agreement with previous detailed characterisation
studies of ppAAm coatings showing a mixture of amine, imine,
nitrile and carbonyl groups.45
For the ppTMP coating, the attenuation of the N 1s signal
compared to the uncoated gel provided a coating thickness
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Table 2 Atomic concentrations of elements detected using the XPS wide scans for the different coating conditions studied, with thickness estimatemeasured using substrate attenuation
Surface codingMonomermolecular formula
Estimatingcoating thickness/nm
Atomic concentrations of the surface coatings (%)
C N O Si P Na Cu CI F
ppAAc C3H4O2 4.5 67.1 6.2 26.1 0.3 0.0 0.1 0.2 0.0 0.0ppTMP C3H9O3P 0.5 44.1 10.5 34.1 8.0 1.1 0.3 0.1 0.2 1.6ppAAm C3H7N 13 69.8 22.5 6.8 0.2 0.0 0.0 0.2 0.1 0.4Collagen type I >15 63.3 16.7 18.2 0.0 0.0 0.9 0.1 0.0 0.8Uncoated polyacrylamide gel C3H5NO 0 43.3 12.5 33.9 8.1 0.0 0.4 0.2 0.1 1.5
Fig. 4 XPS C 1s core levels for the different coating conditions fit with
synthetic components of equal FWHM.
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estimate of 0.5 nm. The surface phosphorus concentration was
1.1%. The C1s core level component at a binding energy of
286.2 eV (component II) was assigned to C–N bonds present in
the underlying polyacrylamide gel after oxygen etching. A small
proportion is likely to correspond to C–P–O bonds, consistent
with the phosphorus elemental composition. Previous work by
Lin et al. was unable to differentiate between the assignment of
the phosphorus environment to phosphate or phosphite groups
and our data are consistent with this.33 The thin ppTMP coating
was accompanied by an increase in surface wettability compared
to uncoated polyacrylamide, which is likely to be influenced, at
Table 3 C 1s peak components observed by XPS analysis for each coating con
Coating Components Bin
Uncoated polyacrylamide gel C1s I 285C1s II 286C1s III 288
Collagen C1s I 285C1s II 286C1s III 288
ppAAM C1s I 285C1s II 286C1s III 288
ppTMP C1s I 285C1s II 286
ppAAc C1s I 285C1s II 286C1s III 288
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least in part, by the preliminary oxygen plasma activation of the
polyacrylamide gel surface, which lead to a very low initial
contact value (17�).The thickness of the ppAAc coating was estimated as 4.5 nm
from the attenuation of the N 1s signal from the underlying gel.
For the C 1s peak, two additional components were required in
addition to the hydrocarbon component. The component at
286.5 eV was attributed to C–OR bonds (alcohol or ether func-
tionalities) and II at 288.7 eV corresponding to either ester or
carboxylic groups. Since component II is smaller than compo-
nent III, it is likely that there are carboxylic acid functionality at
the surface, although a full quantitative assignment between
carboxylic and ester groups would require the use of
derivatization.46–48
3.5 Cell attachment
Human MSCs were seeded at 1000 cells cm�3 on the coated gels
and manual cell counts were performed under the light micro-
scope after three and fourteen days of culture. Examples of light
microscope images obtained for each culture condition are
shown in the ESI (Fig. S4 and S5†). Cell densities quantified
using multiple cell counts from light microscope images are
presented in Fig. 5. The hMSCs were found to attach to all
coated gels whereas uncoated gels and gels treated with oxygen
plasma did not support cell adhesion. From this observation, we
infer that the coatings were sufficiently stable in aqueous media
to allow cell attachment. Although the XPS data suggested that
the ppTMP coating is very thin, cell attachment on this sample
dition attributed to functional groups according to Beamson and Briggs44
ding energy/ev% oftotal C1s
Correspondingfunctional groups
.0 74.4 Aliphatic carbon
.4 15.3 N–C–N
.2 10.3 (C]O)N
.0 66.6 Aliphatic carbon
.3 12 C–(C]O)N
.1 21.4 (C]O)N
.0 55.6 Aliphatic carbon
.2 35.4 C–N
.0 9 C]O
.0 71.2 Aliphatic carbon
.2 28.8 C–O, C–P–O
.0 69.2 Aliphatic carbon
.5 14.5 C–O
.7 16.3 COOR
Soft Matter, 2011, 7, 6501–6514 | 6507
Fig. 5 Cell attachment as a function of surface stiffness and chemistry after three days (A) and fourteen days (B) of culture on the coated poly-
acrylamide gels. The values represent the mean density calculated from light microscope images of 1 mm2 area. Errors bars represent standard devi-
ations. For both time points, one way ANOVA was used to compare the cell densities depending on the surface stiffness for each coating. ‘‘*’’ and ‘‘**’’
indicate that at least one of the means is significantly different from the others at the 0.05 level and 0.01 level, respectively. ‘‘-’’ indicates that the means are
not significantly different at the 0.05 level. The overall mean cell densities for each coating were also compared by one way ANOVA, which showed that
after fourteen days of culture at least one mean was significantly different from the others. Pairwise comparisons showing significant differences are
shown below the graph. ‘‘#’’ and ‘‘##’’ indicate that the means compared are significantly different from each other at the 0.05 level and 0.01 level,
respectively. ‘‘-’’ below the graph of Fig. 5A indicates that the means are not significantly different at the 0.05 level.
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category was comparable to the other coatings showing that
plasma coating was successful. After three days of culture,
a significant difference in cell density was observed between gels
of different stiffness for each surface chemistry (Table S4, ESI†).
A significant difference in cell density depending on the surface
coating was also observed by one way ANOVA. However, post
hoc Bonferroni means comparisons found no significant differ-
ences when the coatings studied were compared pairwise (Tables
S5 and S6, ESI†). There was no trend of cell density versus
stiffness common to the different chemistries.
After fourteen days of culture, the cell densities dropped for all
gels except for the collagen coated gels, which maintained similar
cell densities after three and fourteen days of culture. The cell
densities on collagen coated gels were found to be significantly
higher than the ones observed on the plasma coatings, which had
comparable cell densities. Although a trend towards an increase
in cell density with surface modulus of the polyacrylamide gels
could be identified, the differences were not statistically signifi-
cant at the 0.05 level (Tables S7, S8, S9 and S10, ESI†). Plasma
polymer- and collagen coated glass showed increased cell
6508 | Soft Matter, 2011, 7, 6501–6514
numbers after fourteen days of culture compared to three days
indicating good cell viability and proliferation.
The higher cell densities observed on the coated glass surfaces
compared to the coated polyacrylamide gels might suggest that
the high stiffness of glass (70 GPa49) promotes stronger adhesion,
as previously observed on mouse NIH-3T3 fibroblasts.50 The
lower adhesion to the flexible coated gels could be contributed to
by swelling of the gel in cell culture medium compared to the
plasma polymer coating of the dry material. This might result in
incomplete plasma polymer coating coverage in the wet state
although it is not possible to discern whether this occurs.
Although the exact structure of the water swollen plasma poly-
mer coated gels is unclear, they achieve good cell adhesion after
hydration indicating that the plasma coating is sufficiently
resilient to achieve this desired functional outcome.
Collagen coating showed the highest cell attachment after
fourteen days of culture, indicating that the collagen cross-linked
to the gel surface promotes cell adhesion more effectively than
the proteins adsorbed from the culture media. The reduction in
the number of the attached cells on the plasma coated gels
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between three and fourteen days of culture suggests that the
plasma coatings might be soluble in aqueous conditions and
degrade over time, a hypothesis supported by previous
studies.51,52 However, another study using plasma polymerized
allylamine and hexane for cell culture found that proteins present
in the cell culture medium could act as a protective coating for
the plasma polymers, preventing its degradation.53
3.6. Cell area and cytoskeletal organization
Cell areas were measured after fixation and the results are shown
in Fig. 6. A high variation in cell areas was observed on each
sample with a range of very different cell shapes. The mean cell
area ranged from 1000 mm2 to 5000 mm2 and generally increased
with increasing substrate stiffness but were not affected by the
surface chemistry or the culture medium used. ANOVA analysis
showed that the surface stiffness significantly affected cell area
for all surface coatings cultured in DMEM (p-value < 0.01,
Tables S11 and S12, ESI†) and for cells on ppAAc- and collagen
coated gels cultured in osteogenic medium (p-value < 0.01,
Tables S13 and S14†). Post hoc Bonferroni tests showed that the
significant effect of surface stiffness on cell area observed by
ANOVAwas mostly due to the mean area values obtained on the
softest gel category for all the coatings studied. Cell shape on the
softest gels was different compared to cells on the other gel
categories, with very low cell area and a rounded appearance
Fig. 6 Cell area as a function of surface stiffness. hMSCs cultured for
fourteen days on polyacrylamide gels coated with plasma polymers and
collagen type I and cultured in DMEM (A) or in osteogenic medium (B).
Error bars represent standard deviations. One way ANOVA analysis
showed significant differences in cell area depending on surface stiffness
(p < 0.01).
This journal is ª The Royal Society of Chemistry 2011
seen on gels with stiffness values from 6.5 kPa to 8.5 kPa. Cell
spreading is related to cell adhesion, which involves the cyto-
skeleton. Actin filaments are linked to focal adhesions, which
provide the pathway of force transmission inside the cells. Actin
filament staining using FITC-Phalloidin was used to study
cytoskeletal changes depending on stiffness for ppAAm coated
gels and the obtained confocal microscope images are shown in
Fig. 7.
Actin staining on the softest ppAAm coated gels (8.5 kPa)
showed very diffuse structures. On stiffer gels (10 kPa and
27.5 kPa), the actin structure became more organized and for the
higher stiffness values (53 kPa to glass stiffness), the actin fila-
ments were clearly visible and organized, with the presence of
stress fibers, which appear as long, thick actin bundles that span
across the cell body. Staining of cells on coated glass was similar
to the one obtained on the stiffer gels (53 kPa and 90 kPa).
These results are in agreement with the cell area measurements
where cells on soft gels have a very limited spreading for all
surface coatings tested and cell area increases with surface stiff-
ness but reaches a plateau for stiffness higher than 40 kPa. In
a previous study on the effect of surface stiffness on hMSC,
Buxboim et al. observed that cell morphology did not change
between cells on gels having an osteoid like stiffness (34 kPa) and
cells on glass.40 They also observed cell areas comprised between
500 mm2 and 2500 mm2, which are in the same range as the present
study. Rowlands et al. observed a rounded cell morphology for
0.7 kPa surface stiffness with no visible cytoskeletal organization,
for all protein coatings tested.23 This cell morphology is very
similar to the one observed in the present study on the softest gel
category for each coating (between 6.5 kPa and 8.5 kPa).
Fig. 7 Confocal microscope images of FITC-phalloidin stained hMSCs
cultured on ppAAm coated polyacrylamide gels of varying surface
stiffness. Scale bar is 200 mm.
Soft Matter, 2011, 7, 6501–6514 | 6509
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3.7. Assessment of cell differentiation
Expression of differentiation markers was studied after fourteen
days of culture on the coated gels and glass surfaces. Runx2,
MyoD1 and b-III tubulin protein expression was studied by
immunocytochemistry. The effect of osteogenic induction through
soluble chemical factors on Runx2 expression was also studied.
Fluorescence intensities for the three markers are shown in Fig. 8.
Examples of confocal images obtained on the surfaces with the
maximal expressionof differentiationmarkers (ppAAc forMyoD1
and b-III tubulin, ppTMP for Runx2) are shown for selected
surface stiffness values. Examples of confocal images for each
surface chemistry and each marker are available in ESI (Fig. S6†).
For the plot of b-III tubulin in Fig. 8A, it was apparent that
hMSCs on ppAAc coated gels had a higher expression than the
cells on the other surface coatings for the whole stiffness range
studied. For all surface chemistries, one way ANOVA found
a significant effect of surface stiffness on the fluorescence inten-
sity measured (p-value < 0.01, Tables S15 and S16, ESI†). There
was a clear trend towards a decrease of b-III tubulin expression
with increasing surface stiffness on ppAAm and ppAAc coated
samples. Bertani et al. observed that expression of b-III tubulin is
strongly associated with a neuronal phenotype in mice but in
human can show a broader expression encompassing also non-
neural tissues.54 This might explain why in the present study, b-
III tubulin expression was observed for all surface stiffness values
studied. They also observed a weak constitutive expression of b-
III tubulin in hMSCs.
Fig. 8 Relative fluorescence intensity per cell as a function of surface stiffness
on glass coverslips coated with various pulsed plasma polymer coatings (ppAA
tubulin expression (examples of images obtained for hMSCs on ppAAc coated
for hMSCs on ppAAc coated samples are shown); (C) Runx2 expression (e
shown). For each marker, the fluorescence values are divided by the maximum
bars represent standard deviations. ANOVA testing showed that the fluoresce
marker, the significant differences between the data point with the maximal fluo
MyoD1 and 41 kPa on ppTMP for Runx2) and the remaining data points for
is significantly different from the maximal data point, at the 0.01 level).
6510 | Soft Matter, 2011, 7, 6501–6514
For MyoD1, ANOVA analysis showed a significant effect of
surface stiffness on MyoD1 expression for all surface chemistries
studied (p-value < 0.05, Tables S17 and S18, ESI†). A clear
increase in MyoD1 expression was observed for cells on ppAAc
coated gels of Young’s modulus values comprised between
6.5 kPa and 35 kPa compared to the other surface coatings. At
the stiffness with the maximum level of MyoD1 expression
(10 kPa), fluorescence levels on the ppAAc coating were 2.5 times
higher than on the ppTMP coating, which showed the second
most elevated MyoD1 expression. Fluorescence levels dropped
for the 75 kPa surface, which could be caused by unfavourable
conditions for cell growth on this sample supported by the
observation that this sample had a very low cell density. Fluo-
rescence levels on glass were comparable to those observed for
the gel of 35 kPa surface stiffness, while being significantly lower
than the maximal levels of MyoD1 expression observed for the
lowest stiffness values. A similar pattern of expression was
observed at lower levels on ppTMP coated gels. For collagen and
ppAAm coated gels, a maximum of expression was observed for
stiffness values of 65 kPa and 90 kPa, respectively, the fluores-
cence intensity being only half the maximum one observed for
cells cultured on ppAAc coated gels. The surface stiffness values
achieving maximal MyoD1 expression (10 kPa to 17 kPa) are
within the stiffness range encountered by hMSCs in striated
muscle.55 Myogenic potential of ppAAc coatings has not been
tested previously, past studies suggesting these coatings can
support osteogenesis32 and, for polyacrylic acid coated poly-
styrene, chondrogenesis.56
in hMSCs cultured in DMEM for fourteen days on polyacrylamide gels or
m (:), ppTMP (-), ppAAc (;)) or collagen type I (C) coating. (A) b-III
samples are shown); (B)MyoD1 expression (examples of images obtained
xamples of images obtained for hMSCs on ppTMP coated samples are
fluorescence value observed for the marker. Scale bar is 200 mm. Error
nce of each marker was significantly affected by surface stiffness. For each
rescence value (6.5 kPa on ppAAc for b-III tubulin, 10 kPa on ppAAc for
the same coating are shown (‘‘**’’ indicates that the data point considered
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For Runx2, the ppTMP surfaces induced a higher expression
of Runx2 compared to the other surfaces. The levels of RunX2
expression were also found to be significantly different between
the surface stiffness categories for the plasma polymer coated
gels (p-value < 0.01, Tables S19 and S20, ESI†). For ppTMP
surfaces, the fluorescence intensity was maximal for 41 kPa
surface stiffness. The staining for RunX2 was more intense in the
nucleus than in the cytosolic region (shown in Fig. S7, ESI†). For
the other surface coatings, low levels of Runx2 expression were
observed, apart from an increase at 53 kPa on ppAAm surfaces,
which is thought to be in the relevant stiffness range encountered
by hMSCs in developing bone.22 Levels of Runx2 expression on
glass were consistently low on all tested surface chemistries as
expected, as the stiffness of glass is much higher than the phys-
iological stiffness range.22 Studies have suggested that collagen I
can promote osteogenic differentiation as it is an important part
of developing bone. Benoit et al. observed an important increase
in osteogenesis on surfaces containing phosphate groups.18 This
is consistent with the findings on the ppTMP surfaces of the
present study suggesting that the presence of phosphate groups
as well as the appropriate stiffness act together to promote
Runx2 expression.
The effect of osteogenic medium induction on the expression
of Runx2 was also investigated on all samples and the results are
shown in Fig. 9. For clarity, only ppTMP and ppAAm are
included in the analysis (a figure with all the surface chemistries
can be found in ESI, Fig. S8†). Fluorescence data obtained on
Fig. 9 Relative fluorescence intensity as a function of surface stiffness in
hMSCs cultured for fourteen days on ppAAm coated gels in osteogenic
medium (O) and on ppTMP coated gels in DMEM (-) or in osteogenic
medium (,). The fluorescence values are divided by the maximum value
observed for cells on ppTMP coated gels of 14 kPa surface stiffness in
osteogenic medium. Scale bar is 200 mm. Error bars represent standard
deviations. The significant differences between the data point with the
maximal Runx2 fluorescence value (14 kPa on ppTMP) and the
remaining data points for the same coating are shown (‘‘**’’ indicates that
the data point considered is significantly different from the maximal data
point, at the 0.01 level).
This journal is ª The Royal Society of Chemistry 2011
cells cultured on ppTMP gels in DMEM were also included. For
all surface chemistries except ppAAm coated surfaces, a signifi-
cant effect of the surface stiffness on RunX2 expression was
found by ANOVA (Tables S21 and S22, ESI†). The data also
showed that cells cultured in osteogenic medium did not have an
elevated expression of Runx2 compared to cells cultured in
DMEM at the time point used in the present study (fourteen days
of culture). The maximum of Runx2 expression was shifted for
lower gel stiffness values (14 kPa for cells in osteogenic medium
compared to 41 kPa for cells in DMEM). Runx2 expression for
cells on ppAAm coated gels was low across the range of stiffness
studied.
Studies of Runx2 expression in hMSCs have found contra-
dictory results, some studies suggesting an upregulation of
Runx2 expression during osteogenic specification at the mRNA
and protein level.57–60 Others have observed that Runx2 expres-
sion levels stay constant and an increase in osteogenesis is
reflected by an increase in DNA-binding potential of Runx2 to
its target genes.61–63 This second hypothesis might explain why no
overall increase in Runx2 expression was observed in the present
study for cell cultured in osteogenic medium compared to cells
cultured in DMEM. The lack of increase of RunX2 expression
for cells in osteogenic medium compared to cells cultured in
DMEM might also be explained by the fact that RunX2 is an
early marker of osteogenesis and is downregulated at later stages
in favour of later osteogenic marker such as osteopontin or
osteocalcin.64 Dalby et al.16 observed that expression of an early
marker of osteogenesis such as alkaline phosphatase increased
from day 7 and started to decrease from day 14 where later
markers started to be expressed (osteopontin, osteocalcin). Also,
after fourteen days, they did not find differences in early markers
expression between normal and osteogenic medium, which is also
observed in this study for RunX2.
3.8. Extracellular matrix staining
After fourteen days of culture on the coated polyacrylamide gels,
cells cultured in osteogenic medium showed the presence of cell
aggregates, similar to mineralized nodules observed during
osteogenesis. The number of these structures formed relative to
the surface stiffness and chemistry is shown in Fig. 10A, with
light microscope images corresponding to each chemistry
(Fig. 10B). To test for the presence of calcification within these
aggregates, Alizarin red S staining of calcium was carried out and
examples of staining for each chemistry are shown in Fig. 10C.
One way ANOVA found a significant effect of both surface
stiffness and surface chemistry on the nodule density (Tables S23,
S24 and S25, ESI†). Pairwise post hoc Bonferroni tests found that
ppAAm gels had a significantly higher nodule density than the
other coatings tested. For ppAAm, cell nodules were observed
for all the stiffness values except 8.5 kPa and glass. Alizarin red S
showed a clear staining of the nodules compared to their
surroundings, suggesting mineralization of these structures. This
suggests a specification of the hMSCs towards the osteogenic
lineage. A comparatively high number of nodules was also
observed for ppTMP coated gels of 23 kPa stiffness. ppAAc and
collagen coated gels showed a lower number of nodules and for
just one stiffness category. This suggests that surface chemistry
Soft Matter, 2011, 7, 6501–6514 | 6511
Fig. 10 (A) Quantification of cell nodules on hMSCs cultured with osteogenic medium for fourteen days. The values represent the mean density
calculated from six light microscope images of 1 mm2 area. Errors bars represent standard deviations. The mean nodule densities observed for each
surface stiffness category within each coating were compared by ANOVA. ‘‘**’’ indicates that at least one of the means is significantly different from the
others at the 0.05 level and 0.01 level, respectively. ‘‘-’’ indicates that the means are not significantly different at the 0.05 level. The overall mean nodule
densities for each coating were also compared by one way ANOVA, which showed that at least one mean was significantly different from the others.
Pairwise comparisons showing significant differences are shown below the graph. ‘‘#’’ and ‘‘##’’ indicate that the means is significantly different from
each other at the 0.05 level and 0.01 level, respectively. (B) Light microscope images of nodules. (C) Staining of calcific deposition on cell nodules using
Alizarin red S.
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interacts with signals from the culture medium to promote or
prevent osteogenesis.
Cells on collagen coated gels showed very little evidence of
calcification but it might be that calcification will occur at a later
time point than the one studied here (fourteen days), as past
studies have observed calcification in hMSCs cultures occurring
between 14 and 28 days of osteogenic induction.16,23 The same
might be valid for the cells on ppAAc and ppTMP coated gels.
The conclusion from the data is that ppAAm coatings seem to
favour bone nodule formation and calcification more rapidly or
more effectively than the other surface coatings. The effect of the
surface stiffness is less clear, although the fact that no nodules
were observed on the softest gel category for any chemistry
suggests that these gels were too soft to provide sufficient
mechanical signal to allow osteogenic differentiation. No
nodules were observed for cells cultured in DMEM without
osteogenic supplements, suggesting that the osteogenic induction
was efficient. This shows that the culture medium used has an
6512 | Soft Matter, 2011, 7, 6501–6514
important effect for the osteogenic induction, although antibody
staining of RunX2 protein showed no general increase in RunX2
expression between hMSCs cultured in DMEM and osteogenic
medium.
The comparatively high number of bone nodules observed for
hMSCs on ppAAm coated gels in osteogenic conditions is in
agreement with previous observations showing an increase in
hMSC osteogenic differentiation on glass surfaces functionalised
with NH2 groups compared to other chemistries (–COOH, –CH3
and–OH).65
3.9. Comparison with previous studies on substrate stiffness
induced differentiation of hMSCs
Two previous studies have investigated the effect of surface
stiffness on hMSC lineage specification using the same poly-
acrylamide gel system as the present study. While Engler et al.22
used gels coated with cross-linked collagen type I to allow cell
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adhesion, Rowlands et al.23 tested the effect of different ECM
proteins cross-linked on the gels (collagen I/IV, fibronectin and
lamnin). The present study focussed on the effects of plasma
polymer coatings on hMSC adhesion and differentiation while
including collagen I coated gels to allow comparison with the two
previous studies mentioned.
Engler et al. showed that surface Young’s modulus values
between 0.1 kPa and 1 kPa induced neurogenesis in hMSCs,
8 kPa and 17 kPa induced myogenesis and 25 kPa and 40 kPa
induced osteogenesis. They used AFM nanoindentation to
determine the effective Young’s modulus values. Rowlands et al.
obtained a similar range of stiffness values, between 0.7 kPa and
80 kPa, using the bulk technique of compressive testing for
Young’s modulus determination and showed that a surface
stiffness of 25 kPa promoted myogenesis on fibronectin coated
gels and a surface stiffness of 80 kPa promoted osteogenesis on
collagen type I coated gels. A similar range was measured by
AFM nanoindentation in the present study, although the stiff-
ness range was slightly shifted towards higher Young’s modulus
values.
Cell areas measured by Rowlands et al. increased for 9 kPa gels
and stayed in the same range for stiffer gels whereas increasing
actin skeleton organization was observed, which is consistent
with the results of the present study. The observation of a round
cell morphology contrasts with the results from Engler et al. who
observed formation of branched morphology in hMSCs cultured
on the softest collagen coated polyacrylamide gels (0.1 kPa to
1 kPa), which was interpreted as an initial specification towards
neurogenic lineage. When studying neurogenesis in hMSCs, Qian
and Saltzman found that cells responsive to neuronal induction
developed small, rounded cell bodies but with long cellular
processes forming secondary and tertiary branches.66 This would
suggest that cells in the present study were not specified towards
the neurogenic lineage. For cytoskeletal organization of the cells
depending on surface stiffness, similar observations to our study
and Rowlands et al. study were made by Engler et al.
For MyoD1, the surface stiffness range achieving maximal
expression in the present study (10 kPa to 17 kPa on ppAAc
coated gels) is the same as the one shown to induce myogenesis in
hMSCs by Engler et al. and Rowlands et al. In addition, Row-
lands et al. observed that MyoD1 expression is not very depen-
dent on matrix composition for collagen coated gels. In contrast,
the present study showed an important effect of surface chem-
istry on MyoD1 expression, with ppAAc coatings allowing the
highest expression.
For b-III tubulin, Engler et al. observed protein expression on
collagen coated substrates only in cells cultured on the softest
surfaces from 0.1 kPa to 1 kPa. The cells on these surfaces also
showed development of branched phenotype. The protein was
mainly localised in the axon-like extensions of the cells. This
phenotype was neither observed by Rowlands et al. or by the
present study.
For Runx2, the stiffness range achieving the highest protein
expression in the present study (24 kPa to 79 kPa on ppTMP
coated gels) is within to the stiffness range described by Engler
et al.22 as corresponding to the one of developing bone, between
25 kPa and 100 kPa. Engler et al. observed an increase in Runx2
expression after one week of culture on collagen coated poly-
acrylamide gels in DMEM and in osteogenic culture medium.
This journal is ª The Royal Society of Chemistry 2011
Whereas for DMEM, the upregulation was limited to gels with
stiffness values of 30 kPa to 40 kPa, osteogenic medium lead to
a general increase in Runx2 expression with a higher peak at
30 kPa to 40 kPa. Rowlands et al.23 obtained an increased Runx2
expression for collagen type I coated gels at a stiffness of 80 kPa
compared to the lower stiffness category (25 kPa) and other
proteins coatings, but did not include an intermediate stiffness
value of 40 kPa. The osteogenic effect of collagen coating on
hMSCs suggested by both Engler et al. and Rowlands et al. was
not observed in this study at the time point studied, as the
phosphate containing surfaces showed the highest Runx2
expression.
4. Conclusions
The present study highlights the combined effects of surface
chemistry and variable substrate stiffness on hMSC adhesion,
differentiation and proliferation using variably compliant poly-
acrylamide gels coated with three different plasma polymers
(ppAAm, ppAAc and ppTMP) or type I collagen. It was
demonstrated that plasma polymerisation is able to create a cell
supporting interface while conserving the mechanical properties
of the underlying substrate. It was shown that cell attachment to
the gels after three days of culture, cell spreading and cytoskeletal
organizations were mainly affected by surface stiffness for all the
surface chemistries studied, which supported cell adhesion. Long
term cell survival was related to both surface chemistry and
stiffness. For Runx2 and MyoD1 expression, a combined effect
of chemistry and stiffness was observed, whereas b-III tubulin
expression was mainly affected by surface chemistry. The use of
osteogenic medium for cell culture resulted in the formation of
bone nodules and matrix calcification for gel stiffness values
higher than 10 kPa, especially on the ppAAm coated gels.
Previous studies have used different ECM proteins coated on
the surfaces with varying stiffness and ppAAm and ppAAc
coatings have been used for hMSC culture but this is the first
study to test the effect of different plasma polymer coated
surfaces while at the same time achieving a range of physiolog-
ically relevant Young’s modulus values. This has the advantage
that all the materials used to form the substrate are synthetic.
Acknowledgements
ML is supported through studentship funding from the BBSRC
and Rutherford Appleton Laboratories. The authors also
gratefully thank Professor Xinyong Chen for useful discussions
and assistance with regards the AFM stiffness measurements.
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