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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: morgan.alexander@nottingham.ac.uk; 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

This journal is ª The Royal Society of Chemistry 2011

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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