Regenerating Bone via Multifunctional Coatings ... - ACS Publications

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Regenerating bone via multifunctional coatings: The blending of cell integration and bacterial inhibition properties on the surface of biomaterials Mireia Hoyos-Nogués, Ferran Velasco, Maria-Pau Ginebra, José Maria Manero, F. Javier Gil, and Carlos Mas-Moruno ACS Appl. Mater. Interfaces, Just Accepted Manuscript • Publication Date (Web): 08 Jun 2017 Downloaded from http://pubs.acs.org on June 10, 2017

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ACS Applied Materials & Interfaces

Regenerating bone via multifunctional coatings: The blending of cell integration and bacterial inhibition properties on the surface of biomaterials Mireia Hoyos-Nogués,1,2 Ferran Velasco,1 Maria-Pau Ginebra,1,2,3 José María Manero,1,2 F. Javier Gil1,2,4 and Carlos Mas-Moruno1,2* 1 Biomaterials, Biomechanics and Tissue Engineering Group (BBT), Department of Materials Science and Metallurgical Engineering, Technical University of Catalonia (UPC), 08019 Barcelona, Spain 2 Barcelona Research Center in Multiscale Science and Engineering, UPC, Barcelona, 08019, Spain 3 Institute for Bioengineering of Catalonia (IBEC), 08028 Barcelona, Spain 4 Universitat Internacional de Catalunya (UIC), 08195 Sant Cugat del Vallès, Spain

*Corresponding Author: E-mail: [email protected] Telephone: +34 93 413 79 43 Fax: +34 93 401 67 06

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KEYWORDS. Multifunctionality, antimicrobial peptides, cell adhesive peptides, osseointegration, surface functionalization

ABSTRACT

In dentistry and orthopedics, it is well accepted that implant fixation is a major goal. However, an emerging concern is bacterial infection. Infection of metallic implants can be catastrophic and significantly reduce patient quality of life. Accordingly, in this work we focus on multifunctional coatings to simultaneously address and mitigate both these problems.

We have developed a tailor-made peptide-based chemical platform that integrates the well known RGD cell adhesive sequence and the lactoferrin-derived LF1-11 antimicrobial peptide. The platform was covalently grafted on titanium via silanization and the functionalization process characterized by contact angle, XPS and QCM-D. The presence of the platform statistically improved the adhesion, proliferation and mineralization of osteoblast-like cells compared to control surfaces. At the same time, colonization by representative bacterial strains was significantly reduced on the surfaces. Furthermore, the biological potency of the multifunctional platform was verified in a co-culture in vitro model.

Our findings demonstrate that this multifunctional approach can be useful to functionalize biomaterials to both improve cell integration and reduce the risk of bacterial infection.


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1. INTRODUCTION Once implanted in the body, biomaterials need to fulfill many requirements: they are not only expected to mimic the mechanical and chemical properties of the target tissue, they should also be fully biocompatible and biofunctional, by supporting the adhesion, proliferation and differentiation of host tissue cells1,2. Titanium (Ti) is one of the most extensively used biomaterials due to its biocompatibility, corrosion resistance, lightweight and durability. Nevertheless, Ti-based implants are still associated with limited rates of osseointegration in certain compromised clinical situations and have shown susceptibility to bacterial infections3,4. In this regard, extensive research in the dental and orthopedic fields has focused on achieving increased rates of osseointegration5. However, implant surfaces that facilitate cell adhesion, spreading and proliferation, may also favor a similar behavior on bacterial cells and may promote biofilm formation. This is troublesome as implant-associated infections represent one of the major causes of implant failure, and the prevalence of peri-implantitis is on the rise6,7. On the other hand, research efforts devoted to inhibit bacterial colonization are frequently related to cytotoxic agents or treatments which do not positively affect host tissues8,9. Hence, ideally, to enhance the osseointegration of implant materials and their long-term success, biomaterial surfaces should reduce bacterial colonization levels without compromising the physiological functions of osteoblasts. Yet, the majority of approaches described in the literature tend to only focus on either improving cell adhesion or preventing bacterial infection, but rarely explore a combined effect. Multifunctional approaches, aiming at simultaneously tackling both issues on the surface of biomaterials, are thus emerging as a very potent strategy4.


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The use of antifouling coatings such as poly(ethylene glycol) (PEG) has been widely explored to prevent bacterial adhesion10. However, such low fouling surfaces also restrict mammalian cell adhesion and further modification to rescue osteoblast adhesion is required. For example, poly(L-lysine) (PLL)-PEG copolymers modified with a cell adhesive RGD peptide reduced bacterial adhesion on TiO2 while enhancing the interaction of the materials with osteogenic cells11. Following a similar approach, the immobilization on Ti substrates of other polymers such as poly(methacrylic acid) (PMAA) brushes or chitosan, functionalized with distinct cell adhesive sequences, has been proven useful to simultaneously inhibit bacterial adhesion and preserve (or improve) osteoblastic functions12,13. The above mentioned strategies are based on a bacteriostatic effect and they intrinsically reduce the osteoconductive capacity of Ti. Although the low fouling behavior of the resulting surfaces can be counteracted with cell adhesion molecules, the use of bactericidal approaches without compromising the biofunctionality of Ti represents an interesting alternative that is gaining popularity. For instance, titania nanotubes (TNTs), which have been shown to display beneficial interactions with osteogenic cell types, were anodized with silver nitrate to provide antimicrobial activity14. In pursuit of the same biological effect, TNTs have also been recently loaded with zinc15 or gentamicin-coated polymers16. In our group, we have also reported the use of RGD-decorated polyurethane-polyurea nanoparticles (PUUa NCs) loaded with roxithromycin. These systems were efficiently immobilized on Ti and enhanced osteoblast adhesion and proliferation as well as reduced bacterial adhesion in a dose dependent manner17. Nonetheless, it should be mentioned that these strategies may also have some disadvantages. Polymeric systems may entail difficulties in their fabrication and can be degraded over time under physiological environments4; more importantly, the uncontrolled/excessive release of antibiotics may cause


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cytotoxicity9, lower drug efficiency and the persistent problem of bacteria resistance18. The combination of synthetic peptides with cell adhesive functions and antimicrobial properties covalently anchored to the Ti surface19 would solve most of the aforementioned limitations. However, controlling the exact ratio and the presentation of the peptidic motifs on the biomaterial surface remains elusive when simple peptide mixtures are used. Herein, we present a molecular-based approach to install multifunctionality on the surface of Ti by blending the adhesive potential of the RGD peptide and the antibacterial activity of the LF1-11 peptide (Figure 1). The RGD sequence20, found in fibronectin, vitronectin, osteopontin and other adhesive glycoproteins of the extracellular matrix, has been shown to induce the adhesion of osteoblasts via cell-membrane integrin receptors. The effects of this cell-binding peptide on cell physiology are very well established21. The LF1-11 peptide is an antimicrobial peptide (AMP) derived from the protein lactoferrin (LF). The antibacterial activity of LF is attributed to its ability to sequester ferric ions, thereby depriving potential pathogens of this essential nutrient to growth, as well as its capacity to directly disrupt both gram-negative and gram-positive membranes22,23. Although many AMPs derived from LF have been isolated and characterized24, in this work we focused on LF1-11, which as its name suggests, corresponds to the N-terminal domain of LF, containing the first eleven residues of the protein. This peptide retains the antibacterial properties of LF and has been shown to interfere with the attachment of primary colonizers and early biofilm formation on Ti and other biomaterial surfaces25-28. Thus, we envision that such approach will prove useful to enhance the adhesion of osteoblasts on Ti surfaces while reducing bacterial colonization (Figure 1A). Our strategy relies on the use of a synthetic peptidic platform recently developed by us29,30, which has the capacity to simultaneously present two distinct bioactive sequences in a tightly and chemically controlled


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manner (Figure 1B). In previous studies, the integrin-binding motifs RGD and PHSRN were inserted in the synthetic platform to mimic the spatial orientation present in the cell attachment site of fibronectin. When coated on Ti, this molecule retained the potential of the native protein and improved the activity of osteoblasts and mesenchymal stem cells compared to linear controls and peptide mixtures, which failed to reproduce the optimal orientation of the two motifs. In the present work, such synthetic approach is used to ensure an equimolar presentation (1:1 ratio) of the cell binding and antibacterial sequences on the surfaces. This homogeneous distribution is no attainable with other methods, such as the use of peptide mixtures. To prove the feasibility of this strategy, the peptidic platform and the linear controls were covalently anchored to Ti surfaces, and the biological potential of the functionalized biomaterials evaluated in cellular assays with sarcoma osteogenic cells (SaOS-2) and bacterial cultures with Streptococcus sanguinis (S. sanguinis) and Staphylococcus aureus (S. aureus). Moreover, to ascertain the functionality of the coating in a more realistic scenario, cell behavior was further analyzed in cell-bacteria co-cultures.


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Figure 1. A) The concept of multifunctionality on the biomaterial surface. The combination of a cell adhesive motif (RGD) and an antimicrobial peptide (LF1-11) promotes cell adhesion but inhibits bacterial attachment. B) Schematic representation of the synthetic platform. C) Chemical structure of the platform.


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2. MATERIALS AND METHODS 2.1 Biofunctionalization process 2.1.1 Sample preparation Ti commercially pure (c.p.) grade 2 disks (10 mm diameter, 2 mm thickness) were smoothed up to a surface roughness (Ra) under 40 nm with silicon carbide grinding papers and polished with suspension of alumina particles (1 µm and 0.05 µm particle size) on cotton clothes. Once polished, samples were ultrasonically cleaned with cyclohexane, isopropanol, distilled water, ethanol and acetone (3 x 5 min each) and passivated with 65% (v/v) HNO3 for 1 h. Samples were then cleaned with distilled water, ethanol and acetone, and stored dried. 2.1.2 Synthesis of the biomolecules The platform containing the RGD and LF1-11 peptides was synthesized stepwise by solid-phase peptide synthesis (SPPS). SPPS allows for a modular and straightforward approach to build the multifunctional construct, thus representing an ideal method to synthesize these types of molecules. The linear peptides were also synthesized and used as controls in the biological studies. SPPS was performed following the Fmoc/tBu strategy and using Rink AmideChemMatrix® resin (0.2 g, loading of 0.74 mmol g-1) (Iris Biotech GmbH, Germany) as a solid support. Fmoc-Cys(Trt)-OH was introduced at the C-terminal as anchoring unit, FmocLys(Alloc)-OH as branching point and the PEG-spacer was built with two units of Fmoc-8amino-3,6-dioxaoctanoic acid (Iris Biotech GmbH). The detailed synthetic protocol has been described elsewhere29 and is represented in Figure S1 of the Supporting Information. All


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peptides were purified by means of semi-preparative HPLC and characterized by analytical HPLC and MALDI-TOF (Table S1 of the Supporting Information). 2.1.3 Surface functionalization Peptide functionalization was accomplished by means of a three-step procedure: silanization with 3-aminopropyltriethoxysilane (APTES, Sigma-Aldrich, USA), cross-linking with Nsuccinimidyl-3-maleimidopropionate (SMP, Alfa Aesar, Germany), and, finally, peptide immobilization on the titanium surface at a 100 µM concentration in phosphate-buffered saline (PBS). This protocol has been carefully described in previous studies30,31 and is illustrated in Figure S2 of the Supporting Information. Uncoated polished Ti disks (Ctrol) and fibronectincoated disks (FN) were selected as negative and positive controls, respectively. 2.2 Physicochemical characterization 2.2.1 Surface topography The average surface roughness of the samples (Ra) was determined by white light interferometry using a Wyko NT9300 Optical Profiler (Veeco Instruments, USA) in vertical scanning interferometry mode (5x objective lens and a scanning area of 736 × 480 µm). Three measurements were collected at different positions on three samples per group. Roughness data was analyzed with Wyko Vision 4.10 software (Veeco Instruments). The morphology of the samples was studied by means of scanning electron microscopy (SEM) (Zeiss Neon40 FE-SEM, Carl Zeiss NTS GmbH, Germany). Images were taken for each surface at a working distance of 7 mm and a potential of 5 kV.


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2.2 Contact angle analysis The wettability (i.e. hydrophilicity) of the samples was measured by static contact angle using the sessile drop method (Contact Angle System OCA15 plus, Dataphysics, Germany). Ultrapure distilled water (Millipore Milli-Q, Merck Millipore Corporation, USA) and diiodomethane were used as working fluids. Measurements were acquired in triplicate at room temperature, with an initial volume of 3 µL and a dose rate of 1 µL/min. The surface energy (SE) and its dispersive and polar components were determined using the Young−Laplace and Owen−Wendt equations as previously explained29. Data was analyzed with SCA 20 software (Dataphysics). 2.2.3 Surface chemical composition The chemical composition (atomic percentage) at the surface level was analyzed by X-ray photoelectron spectroscopy (XPS). XPS spectra of the samples were acquired with a nonmonochromatic Mg anode X50 source, operating at 150 W and a Phoibos 150 MCD-9 detector (D8 advance, SPECS Surface Nano Analysis GmbH, Germany). Detector pass energy was fixed at 25 eV with 0.1 eV steps to record high resolution spectra at a pressure below 7.5 × 10−9 mbar. Peak fittings and spectral analysis were conducted using Casa XPS software (Version 2.3.16, Casa Software Ltd., UK). All binding energies were referenced to the C1s signal located at 284.8 eV. Three samples of each condition were studied. 2.2.4 Peptide density on the surfaces The binding of the peptides on the surface was analyzed by quartz crystal microbalance with dissipation (QCM-D) monitoring. To this end, Ti sensors (QSX310, Q-Sense, Sweden) were first cleaned, activated with oxygen plasma for 5 min at a 12 MHz frequency in an expanded plasma


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cleaner PDC-002 (Harrick Scientific Corporation, USA), and silanized with APTES + SMP as described in section 2.1.3. QCM-D (D300, Q-sense) measurements were performed at 25º C by monitoring changes in frequency, ∆f (Hz), and dissipation, ∆D (×10−6), in real-time using Qsoft software (Q-Sense). To monitor the attachment of the peptides on the sensor surfaces, the baseline was first stabilized with PBS for 30 min, and then peptide solutions were injected at 100 µM in PBS and maintained in the sensor chamber for 60 min. Finally, the biomolecules weakly bound to the surface were rinsed with PBS for 10 min. All raw data was analyzed using QTools software (Q-Sense). 2.3 Biological characterization 2.3.1 Cell culture Human sarcoma osteogenic (SaOS-2) cells (ATCC, USA) were cultured in Mc Coy’s 5A medium supplemented with 10% (v/v) fetal bovine serum (FBS), 2% (v/v) 4-(2-hydroxyethyl)-1piperazineethanesulfonic acid (HEPES), 1% (w/v) sodium pyruvate, 50 µg/mL streptomycin, 50 U/mL penicillin and 1% (w/v) L-glutamine. Cells were maintained at 37ºC in a humidified atmosphere containing 5% (v/v) CO2 and culture medium was changed twice a week. Confluent cells were detached by trypsin-EDTA and subcultured into a new flask. The experiments were carried out with cells at passages 25-35. All reagents were purchased from Sigma–Aldrich, unless otherwise noted. 2.3.2 Cell adhesion To compare the efficiency of the different coatings in terms of cell attachment, the number of adherent cells on Ti surfaces was evaluated by means of an enzymatic assay. Firstly, samples


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were washed three times with PBS and blocked in 1% (w/v) bovine serum albumin (BSA) for 50 min at room temperature to avoid non-specific protein adsorption. Subsequently, cells were seeded at 5 × 104 cells/mL (25.000 cells per disk) in serum free medium and incubated at 37 ºC. After 4 h of incubation, samples were rinsed with PBS to remove non-adherent cells, and remaining cells were lysed with 350 µL/disk of mammalian protein extraction reagent (M-PER). Enzymatic activity of lactate dehydrogenase (LDH) was quantified by means of a conventional colorimetric assay (Cytotoxicity Detection Kit (LDH), Roche Diagnostics, Germany) using a multimode microplate reader (Infinite M200 PRO, Tecan Group Ltd., Switzerland). To obtain cell numbers from the absorbance read-out of the test, a standard curve of defined cell concentrations was applied. 2.3.3 Immunofluorescence analysis of cell morphology Cell spreading on the biofunctionalized samples was studied by immunofluorescence analysis. To this end, samples were treated as described in the previous section and cells (5 × 104 cells/mL) were allowed to attach for 4 h in serum-free medium. After this time, cells were fixed with paraformaldehyde (PFA, 4% w/v in PBS) for 20 min, permeabilized with 0.05% (w/v) Triton X-100 in PBS for 20 min and blocked with 1% BSA (w/v) in PBS for 30 min. Actin fibers were stained by incubating with TRITC-conjugated phalloidin (1:300, in permeabilizing buffer) for 1 h and nuclei were stained using 4’,6-diamidino-2-phenylindole (DAPI) (1:1000, in PBSglycine 20 mM) for 2 min, both in the dark. Between all steps, samples were rinsed three times with PBS-glycine for 5 min. Ti disks were mounted and examined under a fluorescence inverted microscope (AF7000, Leica, Germany) and images processed using Fiji/Image-J package to calculate cell-shape parameters.


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2.3.4 Cell proliferation and mineralization For long term analysis of cellular behavior samples were washed and blocked as described for the cell adhesion experiments. Next, SaOS-2 cells were plated at a concentration of 2 × 104 cells/mL in serum-free medium and incubated for 4 h. After this time, medium was aspired and FBS-supplemented medium was added. For proliferation studies, on days 1, 14 and 27, medium was replaced with Alamar Blue (AB)-containing medium (10% (v/v), (ThermoFisher, Belgium) for 3 h and fluorescence of the dye quantified by a microplate reader (λex = 560 nm; λem = 590 nm). To obtain cell number from the fluorescence read-out, a standard curve of defined cell concentrations was applied. Alternatively, to study the process of mineralization, the extracellular calcium deposits produced by osteoblast-like cells were stained using Alizarin Red S (ARS, Sigma-Aldrich). In this case, cells were incubated for 27 days in osteogenic medium, supplemented with 10 mM β-glycerophosphate, 50 µg/mL ascorbic acid, and 100 nM dexamethasone, and fixed with 4% (w/v) PFA. Ti disks were then washed twice with Milli-Q water and 500 µL/well of 40 mM ARS (pH 4.2) were added. Plates were incubated with the dye for 20 min while gently shaking. Prior to quantification, excess dye was washed off using copious washings with Milli-Q water. Cetylpyridinium chloride (CPC) buffer (10% (w/v) in 10 mM NaH2PO4, pH 7) was added (300 µL/well) for 30 min to elute stain. Supernatant was then collected, diluted 1:2 with CPC buffer and 100 µL aliquots were plated to measure absorbance at 570 nm. 2.3.5 Bacterial culture S. sanguinis and S. aureus were used to study the antibacterial properties of the divalent platform. S. sanguinis was chosen as a model of primary colonizer in biofilm formation and was


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obtained from Colección Española de Cultivos Tipo (CECT 480, Spain). S. sanguinis was grown overnight at 37 ºC in Todd-Hewitt (TH) broth (Scharlab SL, Spain). S. aureus was obtained from Culture Collection University of Göteborg (CCUG 15915, Sweden) and was selected as a model of implant-associated infection. S. aureus was grown in Heart Infusion Broth (BHI) (Scharlab SL, Spain). For both bacterial strains, optical density at 600 nm (OD600) was measured and adjusted to around 0.2 ± 0.01, corresponding to a bacterial concentration of 108 colony forming unit (CFU)/mL. The assays were performed in static conditions and performed using three replicates for each condition. 2.3.6 Inhibition of bacterial adhesion and early stages of biofilm formation Functionalized samples were sterilized with ultraviolet (UV) irradiation for 10 min and then washed twice with PBS. After transferring the samples into a new 48-well plate, samples were incubated with bacterial suspensions of either 40µL of S. sanguinis or 5µL of S. aureus at a concentration of 1 × 108 CFU/mL for 4 h at 37 ºC. After this time, the medium was aspired, the samples washed once with PBS and incubated with 10% (v/v) AB medium solutions for 2h (at 37 ºC). The use of AB to measure bacterial adhesion was based in studies where this method has been employed for several different types of cells, such as bacteria, lymphocytes, or hepatocytes32-34. Well contents were then transferred to a 96-well plate and fluorescence measured at 590 nm. The percentage reduction of the AB signal was calculated using the formula: % reduction of AB = [(FI590 test - FI590 untreated)/(FI590 AB control -FI590 untreated)] × 100


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where FI590 is the fluorescence measured at 590 nm, untreated corresponds to Ti control, AB control corresponds to medium with AB and without bacteria, and test corresponds to our target samples. To evaluate the inhibition of the early stages of biofilm formation, the remaining AB solution was aspired, and samples were washed once with PBS and further incubated for 24 h (S. sanguinis) or 36 h (S. aureus) (incubation time was adjusted for each bacterial strain depending on the growth rate) with 1 mL of bacterial culture medium. After this incubation time, the AB assay was repeated as previously explained but with an incubation time of 30 min. 2.3.7 Bacterial viability by live/dead assays The viability of bacteria was analyzed using a LIVE/DEAD BackLight Bacterial Viability Kit (ThermoFisher). The red-fluorescent nucleic acid staining agent propidium iodide, which only penetrates damaged cell membrane, was used to label dead bacterial cells on the modified samples. SYTO 9 green-fluorescent nucleic acid staining agent, which can penetrate cells both with intact and damaged membranes, was used to label all bacterial cells. After incubation time (4h), the supernatant was removed and the samples washed three times with PBS and incubated with 200 μL of a solution containing the two dyes at room temperature in the dark for 15 min. The dyes-containing solution was prepared by adding 3 μL of SYTO and 3 μL of propidium iodide to 2 mL of PBS buffer. A Zeiss LSM 800 confocal microscope (Carl Zeiss, Jena, Germany) was used to measure bacterial viability. Images of the attached bacteria were acquired using Zen 2.3 software (Carl Zeiss) and the specimens were observed with a 10 lens. The confocal LIVE/DEAD images were


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analyzed and quantified using ImageJ software. The volume ratio of red fluorescence (dead cells) versus green (all cells) indicated the portion of killed cells for each treatment: volume ratio of dead cells = red bacteria / green bacteria 2.3.8 Bacterial attachment and morphology by SEM Visualization of bacteria on Ti surfaces was also performed by SEM. The morphology and integrity of bacteria was analyzed at different concentrations (O.D = 0.1; 0.2; 0.4). To this purpose, bacteria were fixed with 3% (v/v) glutaraldehyde in phosphate buffer 0.1 M for 30 min. Prior to SEM analysis, samples were dehydrated in graded alcohol (ranging from 30% to 100% (v/v)) and sputter coated with carbon (Sputter Coater SCD005, BAL-TEC, Liechtenstein). 2.3.9 SaOS-2 cells and bacteria co-culture assays Biofunctionalized surfaces and control samples were exposed to 500 µL/disk of bacterial suspensions (concentration of 1 × 108 CFU/mL) at 37 ºC for 2 h. Subsequently, non-adhered bacteria were removed washing three times in sterile PBS. Next, SaOS-2 cells (1.3 × 104 cells/cm2) were seeded on bacteria-coated surfaces with complete Mc Coy’s 5A medium supplemented with 2% (v/v) bacterial growth medium. After 16 h of incubation cells were fixed and stained with Alexa Fluor 546 Phalloidin (Invitrogen) and DAPI. The outcome of the assay was expressed as the percentage of surface coverage by SaOS-2 cells in the presence or absence of bacteria. The analysis of cell viability and proliferation at longer time points (24 h and 68 h) was performed in a similar manner, but using a lower cell seeding density (6.9 × 103 cells/cm2). Bacterial attachment was confirmed by SEM visualization. All experiments were performed in


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triplicate for each condition. The medium conditions and the protocol used were adapted from published studies35. 2.4 Statistical analysis Statistically significant differences between groups were assessed by 1-way ANOVA followed by post hoc pairwise comparisons using Tamhanne and Scheffe post hoc test depending on the homogeneity of the variance. Differences were also analyzed by non-parametric Kruksall-Wallis test. Values of all graphs are reported as mean ± standard deviation. The software used for statistical analysis was SPSS statistics (IBM, USA). 3. RESULTS AND DISCUSSION 3.1 Design of the multifunctional coating The first step to define the multifunctional coating was an appropriate selection of the bioactive motifs (Figure 1). The RGD sequence has a well characterized affinity for integrins αvβ3 and α5β1, which are known to have important roles in bone biology36. Accordingly, RGD-based molecules have been widely used to functionalize biomaterials for directing cell adhesion and improving their integration with surrounding tissues21,37. To install antibacterial properties on the synthetic system, the LF1-11 peptide was selected. This peptide has shown potent antibacterial effects on several bacterial strains responsible of implant-associated infections, but showed no toxic effects on other cells such as fibroblasts, erythrocytes and bone cells25,26,38,39. Both peptide sequences were thus ideal candidates to construct the multifunctional platform. As previously introduced, the use of this synthetic platform, unlike other methods, allows an equimolar


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presentation (1:1 ratio) of the two bioactive sequences, which will ensure a homogeneous cell adhesive and antibacterial potential over the surfaces. The use of spacing and anchoring units is another important aspect to take into account in the design of the coating molecules. To ensure a proper separation and accessibility of the RGD and LF1-11 peptides, a PEG-based spacer was included (Figure 1C, brown colored); whereas the covalent binding of the biomolecules to the surfaces was mediated via the thiol group of the cysteine residue, which served as anchoring moiety (Figure 1C, green colored). Silanization was chosen as immobilization protocol because it is based on a well-established method and ensures a more stable binding of the molecules to the surfaces compared to other methods (e.g. physical adsorption)40. The details of this functionalization process are shown in Figure S2 of the Supporting Information. 3.2 Physicochemical characterization An exhaustive characterization of the surfaces is not only important to monitor and asses the success of the functionalization process, but also to analyze surface features that are critical in mediating protein adsorption and cell interactions with the materials. Rough surfaces (Ra at micro-level) have been proved to stimulate a positive response of osteoblast cells on Ti and Ti alloys1,5,41 and to significantly increase bacterial adhesion42,43. To minimize the influence of this factor in our study, samples were polished until achieving homogeneous smooth surfaces with Ra values below 40 nm (Table 1). This roughness value is too low to significantly affect the attachment of bacteria (the so-called “threshold Ra” is determined at 0,2 µm43) or to enhance the adhesion of osteoblast cells41. Moreover, the roughness of the samples, once reduced by the polishing process, was not altered by the


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functionalization steps, in agreement with previous studies31,44. Surface analysis by SEM also confirmed that the morphology of the samples remained unaffected through the biofunctionalization process (Figure S3 of the Supporting Information). These observations allowed us to exclude the influence of surface topographical features on cell and bacterial behavior. The effect of the functionalization of Ti samples on the wettability of the surfaces was monitored by contact angle measurements (Table 1). This is an important aspect, since hydrophilic surfaces (i.e. surfaces with high water wettability) can facilitate protein adsorption and initial cell adhesive events, relevant factors in the process of osseointegration1,45. Moreover, it has been reported that hydrophobic surfaces may lead to higher bacterial colonization46. Hence, wettability represents an important parameter of study, which, ideally, should favor osteoblast adhesion while not promoting biofilm formation. Furthermore, changes in wettability are good indicators of the presence of hydrophilic or hydrophobic chemical groups on a surface. In our samples, passivation of the surface with HNO3 significantly increased the wettability and surface energy (SE) of the samples, which may result from the removal of carbonaceous contaminants and the increase in the Ti oxide layer. Conversely, silanization with APTES reduced the hydrophilicity of the samples, in accordance with the hydrophobicity of the silane chains. The subsequent addition of the SMP crosslinker and the bioactive molecules significantly reduced the water contact angle, which is expected owing to the hydrophilic functional groups of the peptides and the presence of PEG moieties. The contact angle values for diiodomethane remained almost constant through the functionalization steps, illustrating that the changes observed were based on the hydrophilicity of the samples. Accordingly, the higher values of SE were also associated with an augment in its polar component (Table S2 Supporting Information). Interestingly, contact


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angle values for the biomolecule-coated samples were statistically comparable, thereby allowing us to discard the influence of the wettability factor if differences in the biological responses of the coating molecules were observed. Table 1. Physicochemical characterization of the surfaces θ(a) H2O (º)

θ CH2I2 (º) SE(b) (mJ/m2)

Ra(c) (nm)

NP(d)

n.m.(e)

n.m.

n.m.

856 ± 410

Ctrol

84 ± 4 ∆

38 ± 0.7

43 ± 1

28 ± 6

HNO3(f)

64 ± 3 *, ∆

38 ± 2

49 ± 2 *, ∆

29 ± 10

APTES

74 ± 3*

42 ± 3

44 ± 2

34 ± 9

SMP

64 ± 2 *, ∆

46 ± 6

48 ± 2 *, ∆

28 ± 6

RGD

58 ± 5 *, ∆

39 ± 1

53 ± 1 *, ∆

27 ± 11

LF1-11

59 ± 2 *, ∆

38 ± 1

54 ± 3 *, ∆

29 ± 9

PLATF

58 ± 5 *, ∆

39 ± 1

54 ± 3 *, ∆

32 ± 8

(a)

θ = contact angle; (b) SE = surface energy; (c) Ra = average roughness; (d) NP = Ti surfaces not polished; (e) n.m. = not measured; (f) HNO3 = Ti surfaces after passivation treatment * Statistically significant differences vs. control Ti (p < 0.05). ∆ Statistically significant differences vs. APTES condition (p < 0.05).

The success in the immobilization strategy was further demonstrated by means of XPS studies analyzing the atomic composition of the surface of the Ti samples (Table 2). The silanization step was confirmed by the appearance of silicon (Si 2p), an increase in the percentage of N 1s and a reduction in O 1s and Ti 2p signals, in comparison with control titanium. The immobilization of the peptides resulted in further changes in the chemical composition of the samples: the percentages of C 1s and N 1s augmented, while the Si 2p and Ti 2p signals were


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reduced, supporting the presence of the biomolecules and in good agreement with literature reports29,30,31,44. To further characterize the process of biofunctionalization and confirm XPS data, the immobilization of the peptides on the surfaces was also analyzed by QCM-D. As shown in Table 2, binding of the molecules resulted in a measurable layer of the bioactive peptides, which was below 1 nm in thickness in all cases. The highest density of peptide attachment was found for the RGD peptide (53 pmol/cm2), followed by the LF1-11 peptide (37 pmol/cm2) and the platform (13 pmol/cm2). This behavior might be explained by differences in the molecular weight (MW) of the molecules, i.e. MWRGD (810 g/mol) < MWLF1-11 (1752 g/mol) < MWPLATF (2464 g/mol), and the fact that bulkier peptides may encounter more sterical hindrance when binding to the silane layer. Table 2. Surface chemical composition and characterization of peptide immobilization Chemical composition (atomic percentages)(a) C 1s

N 1s

O 1s

Si 2p1/2

Peptide immobilization(b) Ti 2p

Thickness

Density

(nm)

(pmol/cm2)

Ctrol

32 ± 3

1.3 ± 0

39 ± 2

0.6 ± 0.6

28 ± 2

-

-

APTES

39 ± 4

4.1 ± 0.1

29 ± 4

15 ± 4

14 ± 6

-

-

RGD

42 ± 0.7

4.2 ± 0.1

30 ± 0.3

11 ± 0.3

13 ± 0.2

0.43

53

LF1-11

47 ± 3

7.9 ± 0.3

26 ± 0.9

10 ± 0.4

9±2

0.65

37

PLATF

45 ± 0.4

7.4 ± 0.1

27 ± 0.1

10 ± 0.4

10 ± 0.8

0.33

13

(a)

Chemical composition was measured by X-ray photoelectron spectroscopy (XPS) The thickness and density of the peptide layer was characterized by quartz crystal microbalance with dissipation (QCM-D) monitoring (b)


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3.3 Osteoblast-like cells behavior on biofunctionalized surfaces To demonstrate the capacity of the synthetic platform to support and influence osteoblast-like cell behavior, biofunctionalized samples were incubated with human osteoblast-like SaOS-2 cells for 4 h and the number, spreading and morphology of adherent cells was quantified by LDH (Figure 2A) and immunofluorescence (Figure 2B, Figure 2C and Figure S4), respectively. As expected, the presence of the platform (PLATF) yielded a significant increase in both cell attachment and spreading (p < 0.05) compared to uncoated Ti (Ctrol) or samples coated with the human LF1-11 peptide. These values were statistically comparable to those observed in RGDcoated samples, thereby proving the platform retains the full potential of the original cell adhesive peptide. Cell spreading and morphology were further analyzed in terms of cell circularity, aspect ratio values, percentage of spread cells versus round cells and percentage of surface coverage (Figure S4). These studies confirmed the positive effect of the PLATF on cell adhesion: this molecule yielded a lower circularity (higher aspect ratio) and an increased percentage of spread cells and surface coverage compared to controls. Although the density of PLATF on the surfaces was lower than that of RGD, the amount of peptide attached (13 pmol/cm2, Table 2) is well above the minimum density required to support cell adhesion and spreading (i.e. 1-10 fmol/cm2)47. The increase in cell spreading observed on LF1-11 in comparison with Ctrol samples was not expected, but might be attributed to changes in wettability upon addition of the molecule (see Table 1) or unspecific interactions. Fibronectin was used as positive control in all cellular studies. Coating the samples with this protein (FN) resulted in the highest values of cell spreading, but lower levels of cell attachment than RGD and PLATF (p

Regenerating Bone via Multifunctional Coatings ... - ACS Publications

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