Enhancement of hMSC In Vitro Proliferation by Surface Immobilization of a Heparin-Binding Peptide

The use of human Mesenchymal Stem Cells (hMSC) as therapeutic agents for advanced clinical therapies relies on their in vitro expansion. Over the last years, several efforts have been made to optimize hMSC culture protocols, namely by mimicking the cell physiological microenvironment, which strongly relies on signals provided by the extracellular matrix (ECM). ECM glycosaminoglycans, such as heparan-sulfate, sequester adhesive proteins and soluble growth factors at the cell membrane, orchestrating signaling pathways that control cell proliferation. Surfaces exposing the synthetic polypeptide poly(L-lysine, L-leucine) (pKL) have previously been shown to bind heparin from human plasma in a selective and concentration-dependent manner. To evaluate its effect on hMSC expansion, pKL was immobilized onto self-assembled monolayers (SAMs). The pKL-SAMs were able to bind heparin, fibronectin and other serum proteins, as demonstrated by quartz crystal microbalance with dissipation (QCM-D) studies. hMSC adhesion and proliferation were significantly increased in pKL-SAMs compared to controls, most probably related to increased heparin and fibronectin binding to pKL surfaces. This proof-of-concept study highlights the potential of pKL surfaces to improve hMSC in vitro expansion possible through selective heparin/serum protein binding at the cell-material interface.


Introduction
Human Mesenchymal Stem Cells (hMSC) are a heterogeneous set of adult stromal cells that can be isolated from different tissues. They have excellent capacities for regenerative medicine purposes due to their wide differentiation potential and immunomodulatory role [1]. Currently, more than 1400 clinical trials are ongoing worldwide, with very promising results [2][3][4][5]. However, independently of the tissue source, hMSC are scarce in the human body. Therefore, in vitro cell expansion before therapeutic applications is mandatory.
Cell behavior is a dynamic and complex process that depends on multiple signals from the cellular microenvironment, whose main component is the extracellular matrix (ECM). ECM is a 3D complex network of macromolecules (mainly glycosaminoglycans (GAGs), proteins and proteoglycans) where cells are organized in a specific way for each cell/tissue. The ECM structure and composition are highly dynamic, and the cell-matrix interactions directly control fundamental cellular functions, such as adhesion, migration, proliferation, and differentiation, among others [6]. ECM adhesive proteins (e.g., fibronectin, vitronectin, collagen) contain highly conserved domains that bind to cell surface integrin receptors. Moreover, GAGs, such as heparan sulfate, recruit heparin-binding molecules, including The success of pKL covalent immobilization on EG4-SAMs was confirmed by the increase in surface thickness and the water contact angle. Figure 1A shows the average thickness of the different surfaces determined by ellipsometry. In the absence of pKL, EG4-SAMs displayed a thickness ranging from 2.5 ± 0.3 nm (EG4-SAMs) to 3.5 ± 0.3 nm after CDI activation (30 CDI buffer). Following pKL immobilization, the average thickness ranged from 2.8 ± 0.3 nm when pKL was only adsorbed on EG4-SAMs (0 CDI pKL) to 7.6 ± 1.2 nm when pKL was immobilized on surfaces pre-activated with the higher CDI concentration (30 CDI pKL). These results are in accordance with those obtained in our previous work [24], confirming that pKL was covalently immobilized onto CDI-activated surfaces. Figure 1B depicts the water contact angle measurements for the different surfaces. Contact angle measurement is a qualitative way to assess surface wettability. It is based on the evaluation of the intermolecular interactions between the surface and a small drop of water when the drop meets the surface. By definition, water contact angles of hydrophilic surfaces are typically smaller than 90 • , while hydrophobic ones have contact angles greater than 90 • [26]. hydrophilic surfaces are typically smaller than 90°, while hydrophobic ones have contact angles greater than 90° [26].
Concerning surface activation with CDI, the water contact angle of surfaces varied from 38 ± 2° for EG4-SAMs to 55 ± 4° after CDI activation (30 CDI buffer). In comparison, the water contact angle of pKL-SAMs ranged from 61 ± 1° when pKL was only adsorbed (0 CDI pKL) to 66 ± 3° when pKL was grafted on CDI-activated SAMs (30 CDI pKL). These results also suggest that pKL was successfully grafted onto CDI-activated EG4-SAMs. However, with this technique, some adsorbed pKL is also detected, as previously reported [24].
Concerning surface activation with CDI, the water contact angle of surfaces varied from 38 ± 2 • for EG4-SAMs to 55 ± 4 • after CDI activation (30 CDI buffer). In comparison, the water contact angle of pKL-SAMs ranged from 61 ± 1 • when pKL was only adsorbed (0 CDI pKL) to 66 ± 3 • when pKL was grafted on CDI-activated SAMs (30 CDI pKL). These results also suggest that pKL was successfully grafted onto CDI-activated EG4-SAMs. However, with this technique, some adsorbed pKL is also detected, as previously reported [24].

Heparin and Fibronectin Adsorption on pKL-SAMs
The ability of pKL-SAMs to bind heparin, fibronectin or fetal bovine serum (FBS) was evaluated using a Quartz Crystal Microbalance with Dissipation (QCM-D). Results in Figure 2 were calculated by applying the Voigt model to QCM-D data (frequency and dissipation). Representative QCM-D data plots are shown in Supporting Information (Figures S4-S6).  A significantly higher amount of heparin was adsorbed on pKL-SAMs when compared with the controls (0.3 and 30 CDI buffer) ( Figure 2A). Moreover, twice the amount of heparin (379 ± 2 ng/cm 2 ) adsorbed on 30 CDI pKL surfaces compared to that adsorbed on 0.3 CDI pKL (182 ± 0.01 ng/cm 2 ). Heparin adsorption to the controls 0.3 and 30 CDI buffer was very low (31 ± 1 ng/cm 2 and 27 ± 2 ng/cm 2 , respectively). These results demonstrated the capacity of pKL-SAMs to bind heparin in a concentration-dependent way.
Fibronectin adsorption to pKL surfaces ( Figure 2B) also occurred in a concentrationdependent manner: 926 ± 67 ng/cm 2 of fibronectin adsorbed to 30 CDI pKL surface and 693 ± 112 ng/cm 2 to 0.3 CDI pKL surface. However, fibronectin adsorption was also detected on the control surfaces: 532 ± 65 ng/cm 2 for 30 CDI buffer and 159 ± 39 ng/cm 2 for 0.3 CDI buffer surfaces, suggesting that pKL also binds fibronectin in a non-specific way.
Protein adsorption on pKL surfaces incubated with 1% FBS was also measured by QCM-D ( Figure 2C) to evaluate whether proteins present in serum-supplemented cell culture medium would also bind to the cell culture surface. The amount of serum proteins adsorbed was 774 ± 1 ng/cm 2 on 30 CDI pKL and 281 ± 2 ng/cm 2 on 0.3 CDI pKL surfaces, against 489 ± 7 ng/cm 2 on 30 CDI buffer and 131 ± 32 ng/cm 2 on 0.3 CDI buffer. Results suggested that, in comparison with fibronectin adsorption from a pure solution ( Figure 2B), the decrease of adsorbed mass from FBS in 0.3 CDI pKL SAMs could be related to heparin adsorption. This fact was based on the described specific heparin adsorption on surfaces with lower pKL concentration, namely surfaces activated with CDI concentrations up to 0.3 mg/mL [24].

hMSC Adhesion and Morphology on pKL-SAMs
hMSC morphology on pKL-SAMs and TCPE surfaces was evaluated by fluorescent staining of F-actin, vinculin and nuclei 24 h after cell seeding ( Figure 3A). hMSC on pKL-SAMs (0.3 CDI-and 30 CDI-pKL) showed a more elongated shape (in green) and more focal adhesions formation (in red), suggesting a better cytoskeleton organization than cells on TCPE or in SAMs with adsorbed pKL (0 CDI). Metabolic activity evaluation on day 1 ( Figure 3B) suggested that pKL surfaces promoted hMSC adhesion since adhesion was significantly increased on pKL surfaces (30 CDI-pKL) as compared to TCPE.

hMSC Adhesion and Morphology on pKL-SAMs
hMSC morphology on pKL-SAMs and TCPE surfaces was evaluated by fluorescent staining of F-actin, vinculin and nuclei 24 h after cell seeding ( Figure 3A). hMSC on pKL-SAMs (0.3 CDI-and 30 CDI-pKL) showed a more elongated shape (in green) and more focal adhesions formation (in red), suggesting a better cytoskeleton organization than cells on TCPE or in SAMs with adsorbed pKL (0 CDI). Metabolic activity evaluation on day 1 ( Figure 3B) suggested that pKL surfaces promoted hMSC adhesion since adhesion was significantly increased on pKL surfaces (30 CDI-pKL) as compared to TCPE.

hMSC Metabolic Activity/Growth on pKL-SAMs
hMSC metabolic activity on pKL-SAMs and controls (buffer-SAMs) was analyzed to infer cell growth (during 7 days) ( Figure 4).
Cells' metabolic activity was significantly higher in pKL-SAMs (either 0.3 or 30 CDI pKL) than in surfaces where pKL was only adsorbed (0 CDI pKL) and in the controls (30 Data are expressed as average ± STDEV of three replicates per condition. * Statistically significant difference compared to TCPE surface (* p < 0.05).

hMSC Metabolic Activity/Growth on pKL-SAMs
hMSC metabolic activity on pKL-SAMs and controls (buffer-SAMs) was analyzed to infer cell growth (during 7 days) ( Figure 4).

Discussion
In the present work, a heparin-binding synthetic polypeptide (poly(L-lysine, L-leucine) (pKL)) was immobilized on self-assembled monolayers (SAMs) to evaluate its potential to improve hMSC in vitro expansion. SAMs were used since they are easy to produce, functionalize and characterize at the molecular level [27]. Moreover, SAMs are relatively stable in biological studies when protected from oxidation, which can be achieved by performing the assays shortly after SAMs preparation. Additionally, in the presence of serum, the adsorbed protein layer can protect SAMs from oxidation during cell adhesion assays and through the first days of cell proliferation [28]. pKL was immobilized onto tetra (ethylene glycol) terminated SAMs (EG4-SAMs) due to its non-fouling characteristics that strongly reduce nonspecific adsorption [29]. Earlier, it was demonstrated that pKL immobilized onto EG4-SAMs in low concentrations bound heparin from human plasma in a selective way [24]. The increase in surface thickness and hydrophobicity after pKL immobilization onto CDI pre-activated SAMs demonstrated that pKL was successfully immobilized onto EG4-SAMs. More details on surface characterization, namely using X-ray photoelectron spectroscopy (XPS), Fourier transform infrared reflection absorption spectroscopy (IRAS), atomic force microscopy (AFM) and electrokinetic analyzer (EKA) techniques can be found in [24,30].
pKL-SAMs improved hMSC adhesion and proliferation. Metabolic activity, evaluated as an index of cell adhesion (24 h after seeding), was significantly increased on pKL-SAMs (30 CDI pKL) as compared with tissue culture polyethylene coverslips (TCPE). Moreover, upon 24 h, hMSC on pKL-SAMs (0.3 and 30 CDI pKL) exhibited a more elongated shape and a more defined cytoskeleton organization than TCPE. In addition, the increase in hMSC metabolic activity over seven days in culture, here used as an indication of cell growth, was much more significant on pKL-SAMs as compared to TCPE and to the other controls (0 CDI pKL and 30 CDI buffer). No differences among surfaces with pKL Cells' metabolic activity was significantly higher in pKL-SAMs (either 0.3 or 30 CDI pKL) than in surfaces where pKL was only adsorbed (0 CDI pKL) and in the controls (30 CDI buffer). Moreover, all the pKL-SAMs had a better hMSC growth-promoting effect than TCPE surfaces. Statistically significant differences were only observed after 5 days in culture ( Figure 4).

Discussion
In the present work, a heparin-binding synthetic polypeptide (poly(L-lysine, L-leucine) (pKL)) was immobilized on self-assembled monolayers (SAMs) to evaluate its potential to improve hMSC in vitro expansion. SAMs were used since they are easy to produce, functionalize and characterize at the molecular level [27]. Moreover, SAMs are relatively stable in biological studies when protected from oxidation, which can be achieved by performing the assays shortly after SAMs preparation. Additionally, in the presence of serum, the adsorbed protein layer can protect SAMs from oxidation during cell adhesion assays and through the first days of cell proliferation [28]. pKL was immobilized onto tetra (ethylene glycol) terminated SAMs (EG4-SAMs) due to its non-fouling characteristics that strongly reduce nonspecific adsorption [29]. Earlier, it was demonstrated that pKL immobilized onto EG4-SAMs in low concentrations bound heparin from human plasma in a selective way [24]. The increase in surface thickness and hydrophobicity after pKL immobilization onto CDI pre-activated SAMs demonstrated that pKL was successfully immobilized onto EG4-SAMs. More details on surface characterization, namely using X-ray photoelectron spectroscopy (XPS), Fourier transform infrared reflection absorption spectroscopy (IRAS), atomic force microscopy (AFM) and electrokinetic analyzer (EKA) techniques can be found in [24,30]. pKL-SAMs improved hMSC adhesion and proliferation. Metabolic activity, evaluated as an index of cell adhesion (24 h after seeding), was significantly increased on pKL-SAMs (30 CDI pKL) as compared with tissue culture polyethylene coverslips (TCPE). Moreover, upon 24 h, hMSC on pKL-SAMs (0.3 and 30 CDI pKL) exhibited a more elongated shape and a more defined cytoskeleton organization than TCPE. In addition, the increase in hMSC metabolic activity over seven days in culture, here used as an indication of cell growth, was much more significant on pKL-SAMs as compared to TCPE and to the other controls (0 CDI pKL and 30 CDI buffer). No differences among surfaces with pKL grafted in different concentrations (0.3 CDI and 30 CDI pKL) were observed throughout the assay (seven days).
The higher hMSC adhesion and proliferation on pKL surfaces could be related to heparin adsorption since it was previously demonstrated that heparin immobilized on anti-adhesive surfaces increased hMSC adhesion [31].
Heparin, fibronectin and FBS adsorption to pKL-SAMs was followed using Quartz Crystal Microbalance with Dissipation (QCM-D) assays. QCM-D is a powerful and highsensitivity tool to quantify molecules (proteins, glycans and more) adsorption/binding to different substrates [27,29,32,33]. For data modeling, the Voigt model was selected. Although in some cases the simpler Sauerbrey equation could have been applied (buffer-SAMs; ∆D < 2 × 10 6 ), surfaces with pKL elicited dissipation shifts that were compatible with a viscoelastic protein layer being formed (Supporting information- Figures S4-S6). Additionally, using the Voigt model did not translate into significant changes in the overall mass (<1% change in modeled mass values) for surfaces where the Sauerbrey equation could have been applied.
Heparin-pKL surface recognition was herein demonstrated by QCM-D assays. The mass of heparin adsorbed was proportional to the amount of surface-bound peptide (higher heparin adsorption on 30 CDI pKL-than on 0.3 CDI pKL-SAMs). However, Martins et al. [24] stated that, in the presence of human plasma, surfaces with higher pKL concentration (30 CDI pKL) were not 100% selective for heparin, allowing the binding of other plasma proteins. Indeed, fibronectin adsorption to pKL-SAMs was very high, especially in 30 CDI pKL-SAMs. pKL-SAMs also adsorbed fetal bovine serum (FBS) in a pKL dose-dependent manner (higher FBS adsorption on 30 CDI pKL-than on 0.3 CDI pKL-SAMs). This may be related not only to the adsorption of heparin from FBS but also to the adsorption of other FBS proteins, such as albumin (the most abundant protein in FBS) or fibronectin (also present in large quantities [34]), mainly in 30 CDI pKL-SAMs that are not selective to heparin [24]. In addition, high FBS adsorption was also detected on 30 CDI buffer-SAMs. This might be related to the adsorption of other FBS proteins since very low heparin adsorption was observed on 30 CDI buffer-SAMs.
We then hypothesized that in hMSC standard growth medium (10% FBS), serum-borne heparin and other compounds, such as fibronectin, could adsorb at pKL-SAMs, exerting beneficial effects onto cell growth. Moreover, since fibronectin has multiple heparin-binding domains [35], it could be adsorbed onto pKL surfaces directly or through a heparin layer bound on pKL-SAMs. Importantly, when adsorbed through a heparin layer, fibronectin assumes the natural conformation found in the ECM [36], favoring cell adhesion. This could explain why both pKL surfaces (0.3 pKL and 30 CDI pKL) had a similar beneficial effect on hMSC growth since lower pKL concentrations could adsorb more heparin from a complex system [24]. On opposite, surfaces without pKL (0.3 and 30 CDI buffer) did not bind serum-borne heparin but could adsorb proteins. Adsorption of albumin, a non-fouling protein that hampers cell adhesion, could explain the lower index of hMSC growth on control surfaces regarding pKL surfaces. Importantly, this event was only detectable after five days in culture. During this period, cells may produce endogenous proteins and growth factors that are then adsorbed at the heparin layer on the surface. This, in turn, favors growth factors binding to cell receptors, amplifying the overall signaling cascade that generates cell proliferation. Further studies will investigate growth factor binding at the pKL surface in terms of identity and quantity.
Our data are in agreement with the literature when it was shown that hMSC proliferation was promoted in GAG-derivatized chitosan membranes containing heparin [37] or when a bioinert HS-EG3-SAMs was functionalized with a heparin-binding peptide (GGGKRTGQYKL) and an RGD peptide required for cell adhesion [23]. Nevertheless, the advantages of pKL as compared to other described heparin-binding peptides are (i) its relatively low production cost on a large scale (easy industry/clinical translation) and (ii) the immobilization simplicity that does not depend on its orientation on the surface.
As the main limitations, this study was fully conducted in model surfaces. In the future, it would be ideal to translate this strategy to polymer surfaces that can be used to coat tissue culture plates. Additionally, the differentiation potential and/or maintenance of the undifferentiated state of hMSC cultured on pKL surfaces was not addressed here and should be further investigated.
Finally, pKL surfaces could be combined with xeno-free media in hMSC expansion. The use of FBS is being banned due to its drawbacks, and chemically defined media and human plasma derivatives are being increasingly used [14,38]. However, the most general approach for hMSC xeno-free expansion requires surface coating with adhesive proteins and supra-physiological concentrations of growth factors in a cell culture medium, which results in high costs [14,39]. pKL surfaces may be used to overcome those limitations: pKL could interact with cell-surface glycosaminoglycans and amplify either fibronectin surface binding and cell-secreted growth factors signaling, resulting in a more affordable strategy.
Overall, this work is a proof-of-concept study presenting a surface-grafted synthetic heparin-binding polypeptide (pKL) able to be employed as a new substrate for hMSC culture, enhancing both cell adhesion and proliferation.

pKL Synthesis and Characterization
Poly(L-lysine, L-leucine) (pKL) was synthesized via ring-opening polymerization of the corresponding amino acid N-carboxyanydrides (NCAs), N-e-benzyloxycarbonyl-Llysine-N-carboxyanhydride (Z-Lys NCA) and L-leucine-N-carboxyanhydride (Leu NCA), according to the experimental procedure described in [24]. pKL schematic representation and further synthesis details can be found in Supporting Information ( Figure S1). pKL was characterized by Fourier Transform Infrared Reflection spectroscopy (FTIR) using a spectrophotometer (PerkinElmer, model 2000) ( Figure S2-Supporting Information). The pKL pellet was prepared by blending 2 mg of the co-polypeptide pKL with 200 mg of Potassium Bromide (KBr) previously dried at 105 • C for 24 h. The infrared spectra were obtained using 100 scans at 4 cm −1 spectral resolution. All spectra were recorded under dry nitrogen to eliminate the water vapor absorption.

Gold Substrates For Surface Characterization and Biological Assays
Gold-coated substrates (1 × 1 cm 2 ), obtained from Instituto de Engenharia de Sistemas e Computadores-Microsistemas e Nanotecnologias, Portugal (INESC-MN), were prepared by the deposition of a thin layer of chromium (2.3 nm) and gold (37 nm) on silicon wafers as previously described [40]. Immediately before use, gold substrates were cleaned with fresh "piranha" solution [7 parts concentrated sulfuric acid (95 vol.%, BDH Prolabo, VWR-International), 3 parts hydrogen peroxide (30 vol.% Merck, Darmstadt, Germany)] for 5 min (caution: this solution reacts violently with many organic materials and should be handled with care). Gold-coated substrates were then rinsed sequentially with absolute ethanol (99.8%, Merck), Type II water (purified and deionized: Resistivity > 1 MΩ/cm; Conductivity < 1 µS/cm; Total organic carbons < 50 ppb) and absolute ethanol once again in an ultrasonic bath for 2 min, followed by drying with a gentle argon stream.

For Adsorption Assays Using Quartz Crystal Microbalance with Dissipation (QCM-D)
Gold-coated quartz crystal sensors (QSX301-Standard Gold; 78 mm 2 active sensor area with a fundamental frequency of 5 MHz) were obtained from Biolin Scientific, Gothenburg, Sweden. Briefly, the sensors were cleaned for 10 min in a UV/ozone oven followed by a 5 min immersion in "piranha" solution as described above for gold-coated substrates.

EG4-SAMs
EG4-SAMs were prepared as described elsewhere [24]. Briefly, (11-mercaptoundecyl)tetra (ethylene glycol) (EG4-thiol; SensoPath Technology, Seattle, WA, USA, SPT-0011P4) was diluted in absolute ethanol to a final concentration of 0.1 mM. Gold substrates (gold-coated substrates or gold-coated quartz crystal sensors) were immersed in EG4-thiol solutions and incubated for 24 h at room temperature (RT) in a nitrogen environment. Afterwards, formed self-assembled monolayers (EG4-SAMs) were rinsed three times in absolute ethanol and dried with an argon stream.
Afterwards, pKL-SAMs and buffer-SAMs were rinsed with DMSO (three times, during 5 min in an ultrasonic bath) and Type I water (Milli-Q ® ultrapure water: Resistivity > 18 MΩ/cm; Conductivity < 0.056 µS/cm; Total organic carbons < 50 ppb) (twice for 5 min in an ultrasonic bath). Finally, samples were dried with an argon stream and stored under an inert atmosphere until further use.

Water Contact Angle
Water contact angle measurements were performed using a system from Data Physics, model OCA 15, equipped with a video CCD camera and SCA 20 software (DataPhysics Instruments GmbH, Filderstadt, Germany). SAMs were placed in a closed, thermostatic chamber saturated with water to prevent evaporation of the liquid from the drop. Measurements were performed using the sessile drop method with Type I water at 25 • C. After 4 µL drop deposition, images were taken every 2 s over 90 s, and the contact angle value was calculated at each time point according to the ellipse method. Then, the water contact angle for each surface was calculated by extrapolating the time-dependent curve to zero. Three replicates per sample were used.

Ellipsometry
Ellipsometry measurements were performed with an ellipsometer model EP 3 from Nanofilm Surface Analysis. The ellipsometer operates in a polarizer compensator-sampleanalyzer (PCSA) mode (null ellipsometry). As a light source, a solid-state laser with a wavelength of 532 nm was used. The gold substrate refractive index (n) and extinction coefficient (k) were determined using a delta and psi spectrum with an angle variation between 66 • and 76 • . To correct any instrument misalignment, measurements were made in four zones. To determine the thickness of SAMs, the same kind of spectrum was used, and n and k for the organic layer were set as 1.45 and zero, respectively, as described [24]. Results are the average of eight measurements on each of the three samples for each type of SAM.
pKL-SAMs and buffer-SAMs assembled onto QCM-D gold-coated crystals were placed in the system, followed by injection (flow rate = 0.1 mL/min) of filtered (0.22 µm pore size) phosphate buffered saline solution (PBS, 0.1 M, pH 7.4) until a stable signal was obtained (baseline) following a protocol adapted from [33]. Then, either heparin sodium salt from porcine mucosa (0.1 mg/mL, Sigma-Aldrich ref. H3393), fibronectin from human plasma (40 µg/mL, Sigma-Aldrich Ref. F0895), or a 1% FBS solution in PBS were injected in the system (flow rate 0.1 mL/min). Incubation proceeded under static conditions at 37 • C for 2 h. Finally, rinsing was performed with PBS to remove any unbound protein (flow rate 0.1 mL/min).
Data were modeled using the Voigt model. This model is indicated when high dissipation changes (∆D > 2 × 10 6 ) are observed [41]. With the Voigt model, the soft/viscoelastic properties are taken into account for the quantification of a hydrated mass adsorbed on the surfaces [33]. Data from the 3rd to the 11th harmonics were collected and used in the analysis. The density and viscosity of the protein solutions were established at 1.35 g/cm 3 and 0.0014 kg/ms (values commonly used for low-concentrated protein solutions) [42,43]. Results are presented in mass per area (ng/cm 2 ) and represent the average of three independent assays with three replicates per sample.

Statistics
To determine the statistical significance of the results from the resazurin assay, a One-Way (from day 0 to day 1) and a Two-Way ANOVA (days 1 to 7) test for multiple comparisons were performed. For ellipsometry and contact angle measurements, a One-Way ANOVA test for multiple comparisons was performed. For QCM-D data, a T-test (non-parametric-Mann-Whitney test) was performed. Statistical significance was set for p < 0.05.

Conclusions
In this work, a new strategy for hMSC in vitro expansion was developed. It relies on the immobilization of a synthetic polypeptide (pKL) on model surfaces. The resulting surfaces were able to bind the most common ECM proteins (heparin and fibronectin), as described by QCM studies. hMSC adhesion and growth were significantly increased on pKL surfaces compared to controls, most probably through selective heparin/serum protein binding at the cell-material interface.