Effect of structure, topography and chemistry on fibroblast adhesion and morphology

Miguel A. Mateos-Timoneda • Oscar Castano •

Josep A. Planell • Elisabeth Engel

Abstract Surface biofunctionalisation of many biode-

gradable polymers is one of the used strategies to improve

the biological activity of such materials. In this work, the

introduction of collagen type I over the surface of a bio-

degradable polymer (poly lactic acid) processed in the forms

of films and fibers leads to an enhancing of the cellular

adhesion of human dermal fibroblast when com- pared to

unmodified polymer and biomolecule-physisorbed polymer

surface. The change of topography of the material does not

affect the cellular adhesion but results in a higher

proliferation of the fibroblast cultured over the fibers.

Moreover, the difference of topography governs the cel-

lular morphology, i.e. cells adopt a more stretched con-

formation where cultured over the films while a more

elongated with lower area morphology are obtained for the

cells grown over the fibers. This study is relevant for

designing and modifying different biodegradable polymers for

their use as scaffolds for different applications in the field

of Tissue Engineering and Regenerative Medicine.

M. A. Mateos-Timoneda (&) · O. Castano ·

J. A. Planell · E. Engel (&)

CIBER de Bioingenierıa, Biomateriales y Nanomedicina

(CIBER-BBN), Barcelona, Spain

e-mail: mamateos@ibecbarcelona.eu

E. Engel

e-mail: eengel@ibecbarcelona.eu

M. A. Mateos-Timoneda · O. Castano · J. A. Planell · E. Engel

Institute for Bioengineering of Catalonia (IBEC), Barcelona,

Spain

O. Castano · J. A. Planell · E. Engel

Department of Material Science and Metallurgical Engineering,

Technical University of Catalonia (UPC), Barcelona, Spain

1 Introduction

Among the myriad of synthetic polymers that are being

used in the biomedical field, poly(lactic acid) (PLA) is one of

the most studied due to its interesting properties, such as

controlled degradation rate, good mechanical properties,

and biocompatibility [1]. However, due to its hydropho-

bicity, the cellular interaction with this material is far from

optimal. Moreover, the lack of chemical groups on its structure

does not allow the grafting of (bio)molecules which enhance

the cellular response. Therefore, many different physical and

chemical methods have been developed for the modification

of the surface of PLA and other biodegradable polymers to

improve its biological activity, directly from the groups

originated or to the covalently attachment of biologically

relevant molecules [2, 3]. Wet chemical methods (i.e.

hydrolysis and aminol- ysis) are amid the most used methods

for surface modifi- cation of biodegradable polymers, leading

to carboxyl and amino groups on the surface of the material

[4–8]. These approaches are simple and cheap methods for

chemical modification of the materials surface, however

they also lead to a modification of the surface roughness

which can have an effect in the posterior cell activity [9,

10]. The choice of the (bio)molecule used to enhance the

biological performance of the material will depend on the

specific application envisioned for the material/device.

Neverthe- less, the use of proteins derived from the

extracellular matrix (ECM) or short peptides derived from

them are the most common approaches [11–13]. Topography of

the surface of the material is another factor known to affect

the cell behavior, specially of stem cells [14–16]. Micro- and

nanostructures produced on the surface of materials induce

changes in cell alignment, polarization, elongation, migration,

proliferation and gene expression [17–19]. This

effect might be attributed to the size of the extracellular

matrix structures which are in the submicron range [20].

One simple and extensively used technique to create

topography in the form of fibers is electrospinning [21, 22].

Electrospinning produces fibers in the micro/nanometer scale

which amplify certain cell responses as contact guidance and

differentiation [23].

In this article, we study the biofunctionalisation with

collagen type I of PLA both in the forms of films and fibers and

investigated the cellular response of human dermal

fibroblast over these substrates. The effect of the different

topographies was evaluated studying the behavior and

morphology of the cells seeded over the diverse

topographies.

2 Materials and methods

2.1 Materials

Poly(95DL/5L) lactic acid copolymer with an intrinsic viscosity

of 6.15 dL cm3

was purchased from PURAC, The Netherlands.

Human recombinant collagen type I was supplied by Fibrogen,

USA. All the others chemicals were purchased by Sigma-

Aldrich and used without further purification.

2.2 Preparation of PLA films

Polymeric films were obtained by solvent-casting of a

2.5 % polymer solution in CHCl3 under a solvent saturated

atmosphere. In short, 2.5 g of PLA were dissolved in 100 ml

of CHCl3, and let the solution stir till everything is completely

dissolved. Afterwards, the solution is cast into a petri dish (12

cm diameter) and the solvent is allowed to evaporate in a

solvent saturated atmosphere for 3 days. The film is cut into

discs with 1.5 cm diameter and they are stored in a

desiccator. The thickness of the obtained PLA

discs is of 150 lm.

2.3 Preparation of PLA fibers

Polymeric fibers were obtained by electrospinning of a 3 % PLA

solution in 2,2,2-trifluoroethanol (TFE). Briefly, PLA solution

was loaded into a syringe and delivered for 20 h at a rate of

0.5 mL/h. An 8 kV potential was used to electrospun the

PLA solution onto a stationary target with a distance

between tip and collector of 12 cm, to permit solvent

evaporation before arriving onto the collector. The ambient

conditions were controlled and set at 13 ºC with a relative

humidity of 52 %.

2.4 Biofunctionalisation of the PLA films and fibers

The biofunctionalisation of the polymeric surfaces was

performed using a three-step protocol: In the first step,

introduction of carboxylic groups by basic hydrolysis of the

surface was performed using NaOH (0.5 M, 15 min for the

fibers and between 2 and 6 h for the films). In the second

step, the activation of the resulting carboxylic groups was

achieved by incubation of the substrates with an EDC/NHS

solution (0.2/0.1 M respectively, MES buffer, pH 6.3) for 4 h.

After several washings of the resulting surfaces, col-

lagen type I solution (10–100 lg/ml, MES buffer, pH 6.3)

was brought into contact with the previously activated

polymer films and fibres and were incubated for 24 h. The

obtained substrates were washed several times with water,

air-dried and store in a desiccator.

2.5 Materials characterisation

The morphology and roughness of the films and fibers were

visualized using scanning electron microscopy (SEM) (ESEM

Quanta 200 FEI, XTE 325/D8395) and optical interferometry

(WYCO NT1100, Veeco), respectively. Contact angle (CA)

measurements were performed by the sessile drop method

using an optical contact angle device (OCA15, Dataphysics,

Germany).

Chemical analysis of the surface of the polymeric films was

obtained by means of Time of Flight-Secondary Ions Mass

Spectroscopy (TOF–SIMS) (TOF–SIMS IV, Ion Tof).

Mechanical testing was performed using an Adamel Lhomargy

DY-34 compression and tension tester (Adamel

Lhomargy, France) with samples of 15 9 10 mm and thickness of 161 ± 27 lm for films and 54 ± 9 lm for

fibrous scaffolds. The tensile stress test was monitored

using a speed of 10 mm/min.

2.6 Immunofluorescence of the collagen-functionalised

PLA fibres and films

Immunofluorescence of the collagen-functionalised mate-

rials was performed using mouse monoclonal anti collagen

IgG1 (Santa Cruz Biotechnology, USA) (dilution 1:100) and

Alexa fluor 488 goat anti mouse IgG (H + L) (Invit- rogen)

(dilution 1:250) as primary and secondary anti-

bodies, respectively. In brief, the samples were block for 20

min at room temperature with PBS-Glycine-BSA. After washing

with PBS-Glycine (2 x 5 min), the samples were incubated

with the primary antibody for 45 min at 37 ºC

and washed with PBS-Glycine (2 x 5 min). The incuba-

tion of the secondary antibody was performed in dark during

45 min at 37 ºC. The samples were visualized with a Nikon

E600 upright fluorescence microscope. Different

controls were performed by incubation of either only pri-

mary or secondary antibodies.

2.7 Quantification of the attachment of collagen

to the films

For the quantification of the covalently attachment of

collagen, DQ collagen type I, from bovine skin, fluorescein

conjugated (Invitrogen) (FL-collagen) was used. PLA discs

previously hydrolysed and activated were incubated in

darkness for 24 h in a FL-collagen solution (10–100 lg/ml, MES

buffer, pH 6.3). The supernatant solutions were col-

lected and their fluorescence was measured (excitation: 495

nm; emission: 515 nm). The measurements were car- ried out

in 5 different samples. For visualization purposes, the

resulting surfaces were visualized with a Nikon E600 upright

fluorescence microscope.

2.8 In vitro studies

Human Dermal Fibroblasts (PromoCell, Germany) were

cultured in Dulbecco’s Modified Eagle’s Medium (DMEM)

supplemented with 10 % fetal bovine serum (FBS), 1 mM

Sodium Pyruvate, 2 mM L-Glutamine, and Penicillin–

Streptomycin.

2.8.1 Cell adhesion studies

104

HDF cells were seeded into disc-shape films with a

diameter of 1.5 cm and squares of 1 cm2

of fibers for 4 h in

serum free medium. WST-1 assay was performed to cal-

culate the number of cells attached to the biomaterials

surface. The cells were fixed with 4 % paraformaldehyde

and were stained with rhodaminephalloidin and DAPI.

Phalloidin stains the actin filaments of the cells in red and

DAPI stains the nuclei in blue. Samples were viewed with a Leica

TCS40 confocal microscopy.

2.8.2 Cell proliferation studies

104

HDF cells were seeded into disc-shape films with a

diameter of 1.5 cm and squares of 1cm2 of fibers for 4 h in

serum free medium. After this time, the medium was replaced

with complete medium and the cells were allowed to grow over

the different surfaces. The medium was changed every 2

days. At the selected time points, WST-1 assay was

performed to measure the cellular proliferation and other set

of samples were fixed and stained to measure the different

morphological descriptors using Image J software [24].

3 Results and discussion

With the aim of study the effect of the topography in the

biological response of the biofunctionalised PLA, films and

fibers were prepared. The films were prepared by solvent-

casting, while the fibers were prepared by electrospinning

[25]. Afterwards, both films and fibers were functionalised

with collagen type I following a 3-step protocol. The first

step is the basic hydrolysis of the surface of the material,

creating a carboxylic-rich surface [26]. Using standard

carbodiimide chemistry, the resulting carboxylic groups were

activated obtaining a NHS-activated surface. The last step is

the incubation of the samples with a collagen type I solution

(Fig. 1).

It is well-known that basic hydrolysis of the surface of

PLA leads to the formation of carboxylic groups on the

surface of the materials due to the breakage of the ester

groups of the polymeric chain. This hydrolysis of the ester

groups can alter the roughness of the material surface, and in

the case of the polymeric fibers, it leads to a decrease of

their thickness. Therefore, contact angle measurements (CA)

were performed to check the change on hydropho- bicity of

the surface upon basic hydrolysis. Scanning

Fig. 1 Schematic

representation of the

biofunctionalisation procedure

for the PLA films and fibers

Fig. 2 CA and SEM images of untreated PLA (a, d), and PLA films hydrolysed for 2 h (b, e) and 6 h (c, f). The scale bar of the SEM images is

600 nm for d and 100 mm for e and f, respectively

Fig. 3 Tof-SIMS spectra of untreated PLA film (top) and PLA film

after the hydrolysis and activation steps (bottom)

Fig. 4 Immunofluorescence images of collagen I-functionalised films

(left) and fibers (right). Both pictures were taken at 409 magnification

and the scale bar represents 100 lm

electron microscopy (SEM) and white light interferometry

studies were carried out to examine the surface topography

upon hydrolysis. The water contact angle of PLA surfaces

decreases from the initial value (80.5 ± 5.18) to more hydrophilic values (34.0 ± 8.1º) increasing the time of

hydrolysis up to 6 h (Fig. 2a). This increase in hydrophi-

licity is accompanied by a change in the roughness of the

surface as can be seen in Fig. 2b–d. The roughness

increases from 79 nm of the as-obtained PLA films to

420 nm and [2 lm after 2 and 6 h of hydrolysis, respec-

tively. Similarly, the CA of the fibers also decreases from

150 ± 378 to 55 ± 128 due to the hydrolysis of the poly-

mer surface. In this case, the degradation of the polymer

leads to a decrease of the thickness of the fibers in 9 %

(1.09 and 25 0.99 lm before and after hydrolysis, respec-

tively). However, mechanical testing of the films before

and after hydrolysis did not show a decrease in the

mechanical properties (young modulus and amax) of the

films. In the case of the fibers, a decrease of the young

modulus was observed after the hydrolysis (data not shown).

For the activation step, the previously hydrolysed

materials were incubated with an EDC/NHS solution (0.1/

0.2 M, MES buffer pH 6.2) for 4 h. The presence of the

more hydrophobic NHS-activated ester groups on the sur-

face of the materials was investigated using CA measure-

ments and Tof–SIMS analysis [27]. The spectra of the

activated surface shows the appearance of a new peak at

m/z 182 that can be attributed to the fragmentation of PLA

containing the NHS ester group on its structure (Fig. 3). It is

important to notice that the intensity of the peaks due to

different fragments of PLA (m/z 141, 142 and 159, 160)

decreases when the surface of the materials have been

activated, indicating the presence of the NHS ester groups

on the surface of the films surfaces.

The last step in the biofunctionalisation procedure is the

actual immobilization of the biomolecule of interest. In our

case, we immobilized human recombinant collagen type I, as it

is the main component of the extracellular matrix (ECM)

of many tissues, such as bone tissue [28]. Conse- quently, the

previously activated surfaces (both in the form of films and

fibers) were incubated for 10 24 h with dif-

ferent concentrations of collagen I (ranging from 10 to

100 lg/ml, MES buffer, pH 6.2). The presence of immo-

bilized collagen I in the surface of the materials was

investigated by immunofluorescence (Fig. 4). Both

Table 1 Quantification of the Collagen I content

Sample Collagen attached (lg) Density (lg cm2)

Covalent 10 lg/ml 2.7 ± 0.6 0.2 ± 0.1

Covalent 25 lg/ml 4.5 ± 0.1 0.3 ± 0.1

Covalent 50 lg/ml 9.8 ± 0.4 0.7 ± 0.1

Covalent 100 lg/ml 18.2 ± 2.5 1.3 ± 0.2

Physisorbed 50 lg/ml 0.3 ± 0.1 0.02 ± 0.01

Fig. 5 Fluorescence microscopy images of a pristine PLA, b hydro-

lysed and activated PLA, incubated with 50 lg/ml of FL-collagen.

The white line in the left picture indicates the border of the surface.

The scale bar represents 100 lm

controls using only the primary or secondary antibodies, and

in samples of pristine PLA showed no fluorescence,

indicating the selectivity of the proposed method for the

detection of immobilized collagen I on the surface of the

different topographies made of PLA.

In order to quantify the amount of collagen that is

covalently grafted to the polymeric surface, the immobili-

zation of fluorescently labeled collagen (FL-collagen) was

studied. PLA discs of 1.5 cm of diameter were previously

activated following the previously described protocol and

were incubated with different concentrations of FL-colla- gen (from 10 lg/ml to 100 lg/ml) during 24 h. After- wards, the fluorescence of the supernatant solution was measured (calibration curve, y = 30x 0.00129 + 0.229,

R2

= 0.9511). The results are summarized in Table 1. The

grafting of collagen shows a linear behavior and only

around 20 % of the initial amount is covalently attached to the

surface when it has been previously activated and only 1 %

when the collagen is just physisorbed on the surface.

Moreover, visualization of the resulting surfaces shows the

presence of the fluorescent collagen only when the surfaces

have been hydrolysed and activated previously to the col-

lagen attachment (Fig. 5).

The initial adhesion of human dermal fibroblast (HDF) on

the surface of both films and fibers was studied after

incubation for 4 h in a serum-free medium, to rule out the

cell adhesion due to the adsorption of serum proteins. The

results are depicted in Fig. 6. The seeding efficiency on

both fibers and films is very low, between 4 and 10 % of

Fig. 6 HDF cell seeding efficiency for the different materials and

topographies after 4 h of incubation in serum free medium

the seeded cells are attached to the surfaces after 4 h of

culture. This seeding density does not depend on the

topography of the substrate but strongly depend on the

chemistry of the surface. The pristine PLA and the

hydrolysed PLA showed similar efficiency when compared to

the surfaces in which collagen has been physisorbed, while

the presence of covalently-attached collagen induces a two-

fold increase in the seeding efficiency. However, a slight

increase in cell adhesion is observed. This might be attributed

to changes in both surface roughness and wet- tability of

the substrates. However, due to the biodegrad- able nature

of PLA, it is not possible to decouple both effects, once

the substrates have been hydrolysed.

Moreover, the cells are able to proliferate on these

substrates over a period of 7 days (Fig. 7). HDF were allowed

to adhere on the substrates for 4 h in a serum-free medium.

After this incubation time, the medium was changed for

complete medium and let the cells proliferate, changing the

medium every two days. In the case of the biofunctionalised

films, the cells proliferate constantly over time, while for the

fibers, the cells start to proliferate after day 5. However,

the biofunctionalised fibers seem to increase cell

proliferation. This effect has been attributed to the

difference of topography and to the difference in sur- face

area between the fibers and the films [29]. It is important

to notice that the substrates in which collagen has been

physisorbed do not promote cell proliferation, the cell

population remained constant over this period of time,

behaving similarly to the pure PLA films and fibers (data not

shown).

The cell morphology over these substrates was studied by

means of fluorescence microscopy (Fig. 8). After 4 h of

adhesion time, the cells seeded over the biofunctionalised

films present a circular morphology, while the cells seeded

over the fibers already presented a more elongated mor-

phology. This fact is even more pronounced after one day,

where the cells seeded on the films present normal fibro-

blast morphology and are more stretched over the bio-

functionalised fibers. To analyze the changes in

Fig. 7 Cell proliferation over

7 days on PLA films (left) and

fibers (right)

adhere, and thus they are able to spread. Other interesting

aspect while culturing cells over the biofunctionalised

fibers is that they do not only adhere over the surface of the

mat, they are able to penetrate around 100 lm into the

fibers mat, one of the main limitations to the use of elec-

trospinning mats as scaffolds for tissue engineering

applications.

4 Conclusions

Fig. 8 Fluorescence microscopy images of HDF seeded over func-

tionalised films (a: after 4 h, b: after 24 h) and fibers (c: after 4 h; d:

after 24 h). The scale bar represents 100 mm. Circularity index

(e) and cell area (f) of HDF seeded over collagen I-functionalised

films and fibers

morphology, the cell area and circularity index (a circu-

larity value of 1.0 indicates a perfect circle, while as the

value approaches 0.0, it indicates an increasingly elongated

polygon) was quantified at different time points (3, 5, and 7

days) (Fig. 8). The data shows a higher cell area and

circularity index for the cells grow over the films compared

to the cells seeded over the fibers (at day 7, 1.0 9 104 and

1.54 , compared to 0.3 9 104

lm2

and 0.2, for the cell area

and circularity index, respectively). This behavior can be

attributed to the different topographical features of the

substrates. In the case of the fibers, the cells are forced to

adhere over the biofunctionalised fibers, therefore adopting

an elongated morphology (circularity index close to 0),

while for the flat substrates, the cells have a wider area to

In the present work, the biofunctionalisation with collagen

type I of both fibers and films have been studied and

characterized. This three step protocol led to the successful

covalent attachment of biological active molecules, in this

specific case collagen type I in two different kinds of

topologies created with a biodegradable polymer (PLA). The

biofunctionalisation led to a better biological perfor- mance

of the biodegradable polymer, both in terms of cellular

adhesion as well as cellular proliferation, com- pared with

the pristine polymer and the polymers func- tionalised by

physisorption. Moreover, by changing the topography of the

materials, it is possible to control the shape of the cells

seeded over this polymer. Therefore, the results obtained

open the possibility to use this protocol to biofunctionalise

more complex architectures and topogra- phies, i.e. 3D-

scaffolds for different applications, and to use different

biological active proteins, or peptide sequences with biological

activity.

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