JJOD 2476 1–13 journal of dentistry xxx (2015) xxx–xxx

Available online at www.sciencedirect.com

ScienceDirect journal homepage: www.intl.elsevierhealth.com/journals/jden 1 2 3

Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications

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Elisa Fernandez-Garcia a, Xi Chen b, Carlos F. Gutierrez-Gonzalez a, Adolfo Fernandez a, Sonia Lopez-Esteban c, Conrado Aparicio b,* a

Nanomaterials and Nanotechnology Research Center (CINN), Spanish National Research Council (CSIC)-University of Oviedo-Principado de Asturias, Spain b MDRCBB-Minnesota Dental Research Center for Biomaterials and Biomechanics, Department of Restorative Sciences, University of Minnesota School of Dentistry, USA c Materials Science Institute of Madrid (ICMM), Spanish National Research Council (CSIC), Spain

article info


Article history:

Objective: Titanium materials have been functionalized with biomolecules as a modern

Received 30 January 2015

strategy to incorporate bioactive motifs that will expand and improve their biomedical

Received in revised form

applications. Here, we have biofunctionalized biomaterials based on zirconia of much interest

8 June 2015

for dentistry: the widely used bioceramic 3Y-TZP and a newly developed 3Y-TZP/Ti biocermet.

Accepted 14 June 2015

Methods: The biosurfaces were activated, silanized, and functionalized with coatings made

Available online xxx

of oligopeptides. Surface activation by plasma or alkaline-etching was optimized. The surfaces were coated by tethering a purposely-designed RGD-containing peptide. We


selected this oligopeptide as a model peptide to validate the effectiveness of the biofunc-


tionalization process. Successful treatments after each step of the process were assessed by


surface physical and chemical characterization with water contact angles and XPS, respec-


tively. Coatings’ stability was evaluated after 2 h sonication in water. Pre-osteoblasts

Dental implants

adhesion on the functionalized surfaces was also studied.

Surface modification

Results: 10-min air-plasma treatment effectively activated all types of materials with no


detrimental effects on the material structure and hardness. Nitrogen XPS-peak confirmed that RGD-peptides were chemically-attached on the silanized samples. This was further confirmed by visualizing the functionalized surfaces with flourescence-labelled RGD-peptides before and after ultrasonication. Furthermore, RGD-functionalized surfaces significantly enhanced osteoblast adhesion on all types of substrates, which demonstrated their successful bioactivation. Conclusions: We successfully developed stable functional biocoatings on zirconia and biocermets made of oligopeptides. Surface bioactivation of zirconia-containing components for dental implant applications will enable their improved clinical performance by incorporating signalling oligopeptides to accelerate osseointegration, improve permucosal sealing,

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and/or incorporate antimicrobial properties to prevent peri-implant infections. # 2015 Published by Elsevier Ltd.

17 * Corresponding author at: Minnesota Dental Research Center for Biomaterials and Biomechanics, Department of Restorative Sciences, University of Minnesota School of Dentistry, 16-250a, 515 Delaware St. SE. Minneapolis, MN 55455, USA. Tel.: +1 612 625 4467; fax: +1 612 626 1484. E-mail address: [email protected] (C. Aparicio). http://dx.doi.org/10.1016/j.jdent.2015.06.002 0300-5712/# 2015 Published by Elsevier Ltd.

Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

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Q2 The replacement of hard tissues by synthetic devices is in

many cases the last and most costly solution to deal with the depletion of bone quality and/or the inability to preserve the natural hard tissues. Biomaterials used nowadays to replace those structures, such as those used for making dental implants are not fully compatible with the host bone.1,2 Mechanical incompatibility between the implant and the surrounding bone and lack of bioactive interactions between the inert biomaterial surface and the natural tissues are among the most significant causes for implant failures at mid and long term.3 Thus, the development of materials with improved biomechanics that have additional osseostimulative surfaces are of need for better addressing the clinical demands.4,5 Surface functionalization is currently considered as an effective and modern approach to design new multifunctional materials by changing composition, structure and/or morphology of surfaces without altering the bulk properties.6,7 Physical (increased roughness8,9), chemical (deposition of calcium phosphate phases, chemical etching or incorporation of specific ionic species10,11) and/or biological (bioinspired coatings12,13) approaches can be pursued to functionalize surfaces. Specifically, covalent bonding of biomolecules at surfaces has become pivotal to induct bioactivity on biomaterials14–16 and also for other strategies like assay technologies, biosensors, imaging devices, therapies, etc.17 One of the approaches used in dentistry to covalently immobilize biomolecules on inorganic materials involves methodologies based on silane-chemistry.18–20 Biomolecules with varied complexities like arginine-glycin-aspartate (RGD) sequence ,21,22 other oligopeptides,23,24 proteins,25–27 aptamers,28 recombinamers13,29 or even multiple peptides with cooperative activities30 have been tethered to metal surfaces using silanes as coupling agents. We have extended this approach to develop mechanically and thermochemically stable biofunctional coatings on titanium which resist surgical shear stresses during implantation as well as the challenging bioenvironments derived from the contact with physiological fluids.31 Thus, titanium has been functionalized with different bioactivities either to influence on the physiological paths involved in bone regeneration or to prevent bacterial colonization of the surface.32 Although the methodologies of biofunctionalization are developed for titanium materials with promising results, the potential of chemically biofunctionalizing other biomaterials of more recent interest in dentistry, particularly those that are based on zirconia, is less known.17 Recently, the biofunctionalization of fully-dense zirconia-based materials has been achieved, but coatings were made of adsorbed (physically adsorbed) organic-molecules.33,34 Thus, the strong and stable immobilization of covalently tethered biomolecules on materials containing zirconia is a topic of notable interest still to be investigated. Yttria-stabilized tetragonal zirconia (TZ-3Y-E) presents the best fracture toughness among oxide ceramics, optical properties (colour, translucency) that fit to the aesthetic dental standards and negligible wear debris.35–37 Moreover, ceramic/metal composites (cermets) have been

designed as a new generation of materials that pursue to combine synergistically the dissimilar properties of the monolithical components.38,39 As novel cermets sintered by spark plasma sintering (SPS), ceramic/titanium composites 40,41 displayed tailored mechanical properties as well as a set of surface features that provided biocompatibility parallel or enhanced to the one of its pure counterparts.42 In this work, we aimed biofunctionalizing zirconia (TZ-3Y-E) and a biocermet containing zirconia (zirconia/titanium composite with 75% vol. Ti) using silane chemistry. We based the biofunctionalization route in our already developed method for titanium surfaces.24,30 First, two different methods of activation with different experimental conditions were investigated to optimize this very important step of the process. Then, we used (3chloropropyl) triethoxysilane (CPTES) as covalent linker between the activated surface and the oligopeptide (KKKGGGGRGDS) containing the RGDS-sequence. This amino acid sequence is the most extensively studied cell adhesion motif to functionalize all types of biomaterials,43,44 including those used in dentistry, such as dental implants.45 Indeed, the literature is abundant of examples of substrates coated with RGD peptides, which facilitate the recruitment of cells involved in the formation of bone (osteoblasts), soft tissues (epithelial cells) or pulp regeneration.46 This short cell-binding sequence is expressed in extracellular matrix proteins, such as fibrin, collagen, fibronectin, vitronectin, osteopontin and bone sialoprotein.47 The cell adhesive properties of the RGD sequence rely on being recognized by integrins, the most prominent family of cell membrane receptors. Among them, osteoblasts mostly express a5b1 and avb3.48,49 In addition, our tailored peptide incorporates four glycines and three lysines in the N-terminus for providing spacing between the active RGDS-motif and the inorganic substrate to enable the correct accessibility of the RGD sequence to the integrins;46 and extra groups with free amines to facilitate the nucleophilic reaction with the chlorine-group of the silane molecules, respectively. Thus, we used this bioactive coating as a model to validate our route of bio-immobilization by assessing the enhancement of cellular adhesion of murine pre-osteoblasts in vitro. Coated titanium surfaces were also tested as a reference material.


Materials and methods


Materials preparation

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Tetragonal zirconia stabilized with 3% mol. yttria (TZ-3Y-E), zirconia/titanium (75% vol. Ti) composites and titanium (commercially pure, grade I) samples were obtained following our previous protocols40 and named as ZrO2, Z-75Ti and Ti, respectively. Briefly, stable suspensions were prepared by a wet-processing route of powders with an organic surfactant. After homogenization, drying and sieving, the starting powders were spark plasma sintered (SPS; FCT Systeme GMBH, HPD 25, Germany) at 1250 8C, 80 MPa for 5 min within vacuum, heating at 100 8C/min, to produce composites with high density avoiding side products. The specimens were machined to 2-mm thick square samples (5 mm  5 mm), grinded with SiC discs and finally mirror-polished with 25.5 mm and 1 mm alumina suspensions. Samples were

Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

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ultrasonically cleaned for 25 min in isopropanol, ultrapure distilled (DI)-water and acetone, and stored in a vacuum chamber for at least 2 h.



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Anhydrous pentane, (3-chloropropyl)triethoxysilane (CPTES), diisopropylethylamine (DIEA) and antivinculin primary antibody were purchased from Sigma–Aldrich (St. Louis, MO). 1Ethyl-3-[3-dimethylaminopropyl]carbodiimide hydrochloride (EDC) and N-hydroxysulfosuccinimide (sulfo-NHS) were purchased from Thermo Scientific (Rockford, IL). Alpha minimum essential medium (a-MEM), foetal bovine serum (FBS), trypsin-EDTA (0.5%), Alexa Fluor 488 secondary antibody, 40 ,6-diamidino-2-phenylindole (DAPI) and rhodamine phalloidin were purchased from Invitrogen (Grand Island, NY). 5-FAM cadaverine was purchased from Anaspec (Fremont, CA). Float-A-Lyzer1 dialysis tubes were purchased from Spectrum (Rancho Dominguez, CA). NH2-KKKGGGGRGDS-COOH oligopeptides (purity >98%) were produced by solid-phase peptide synthesis by AAPPTec (Louisville, KY).



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2.3.1. Surface activation: chemical etching and plasma treatments



Alkaline etching and plasma cleaning were investigated as alternative treatments for activating the studied surfaces (ZrO2, Z-75Ti and Ti). This was performed by immersing the samples in a 5 M NaOH solution overnight at 60 8C. The NaOHactivated substrates were rinsed in DI-water for 1 h and dried in a desiccator under vacuum. Otherwise, plasma cleaning is a solvent-free technology, economical and effective for complex-shaped surfaces.50 Plasma treatments (PTs) were applied using a PDC-32G (Harrick Plasma, NY) plasma cleaner. A set of conditions were investigated in order to optimize the treatment: O2-plasma for 5 min (PT1), O2-plasma for 10 min (PT2), air-plasma for 5 min (PT3) and air-plasma for 10 min (PT4). The activated samples were named as summarized in Table 1: ZrO2/NaOH, Z-75Ti/NaOH and Ti/NaOH for alkaline etched-groups and ZrO2/OH, Z-75Ti/OH and Ti/OH for plasmacleaned groups.


Surface silanization of activated surfaces

(3-Chloropropyl) triethoxysilane (CPTES) was used as linker because this avoids additional coupling agents (HBTU, EDC, glutaraldehyde) and allows a high surface coverage with the silanes.13 The activated samples were placed in a fully-dried N2-saturated glass chamber to hinder oligomerization of the organosilane. 7 ml pentane, 1.2 ml CPTES and 0.6 ml DIEA were sequentially added to the chamber. Silanization took place for 1 h at room temperature with application of ultrasounds every 10 min to minimize nonspecific reactions. Further curing process of the peptide coating was not applied. The CPTES-silanized samples were rinsed within ethanol, isopropanol, DI-water, acetone and, finally, dried with a N2-stream. Silanized groups are named CPTES (Table 1).

2.3.3. Immobilization of RGDS-containing oligopeptides on silanized surfaces A peptide solution was obtained by dissolving the oligopeptides (0.5 mg/ml, KKKGGGGRGDS) in a 5 mM Na2CO3 buffer solution (pH 9.5). These oligopeptides have an isoelectric point (IEP) of 10.9 (calculated from Ref. [52]). The CPTES-silanized surfaces were immersed in the peptide solution overnight, in an Ar-atmosphere, to enable the bimolecular nucleophilic substitution. The different samples with the tethered peptides were rinsed in DI-water and stored in a vacuum chamber. These biofunctionalized surfaces were named CPTES/RGD. Furthermore, mechanical-stability of the covalently-immobilized peptide coatings was assessed by applying ultrasonication to the substrates in DI-water for 2 h. These surfaces were named CPTES/RGD/S. Moreover, control group samples with physisorbed RGD-containing peptides were produced on the 3types of substrates by direct immersion of activated samples in the peptide solution under the previously described conditions. These physically-coated surfaces were coded as RGD. The codification for these diverse treatments is also summarized in Table 1.

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Surface topographical inspection

Any significant microstructural and topographical modification on the materials after performing the activation treatments was assessed by scanning electron microscopy (SEM, Analytical TableTop TM3000; Hitachi High Technologies,

Table 1 – Stages involved in biofunctionalization of materials containing 3Y-TZP. The codification and description for each treatment involved in the biological functionalization is detailed. Stage Untreated Surface activation

Silanization Peptide coating

Code Plain NaOH OH


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Treatment Polished surfaces down to 1 mm Chemical etching (5 M NaOH solution; 60 8C, overnight) Plasma-cleaning treatment (PT) O2 plasma for 5 min O2 plasma for 10 min Air plasma for 5 min Air plasma for 10 min Silanization with (3-chloropropyl)triethoxysilane CPTES-silanized and RGD-peptide coated surfaces (chemisorption) CPTES/RGD surfaces after mechanical challenge (ultrasonication for 2 h) Surfaces with adsorbed RGD-peptides (physisorption)


Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

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Japan) operating at 15 kV and 300. The table-top SEM was operated in the combo mode, which gives not only topographical details but also contrast to the image due to different average atomic number composition within the sample. The lower the atomic number the darker the image.



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A microdurometer (5100 Series-Version 2.10, Buehler Micromet, IL) was used to measure Vickers hardness (Hv) of plain (PT0) and plasma-activated substrates (PT1–PT4) 4 samples per group were tested and 15 indentations per sample were evaluated and quantified (Hv). Load of 103 gf for 10 s were applied on ZrO2 and Z-75Ti surfaces and 102 gf for 10 s on Ti surfaces, as the latter is notably softer than the other two materials containing zirconia.



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Surface wettability of plain (ZrO2, Z-75Ti, Ti) and modifiedactivated (OH), silanized (CPTES), coated (CPTES/RGD or RGD), coated and sonicated (CPTES/RGD/S)-substrates was assessed by the sessile drop method using a Dropmaster Series Contact Angle analyzer DM-CE1 (Kiowa Interfaces Co, Ltd., Japan) with Interface Measurements and Analysis System FAMAS V.3.0.0 (Kiowa Interfaces Co, Ltd., Japan) software. 1.5 ml drops of DI-water were dispensed on the tested surfaces. Water contact angles (CA) were measured using the tangent method after drop stabilization on the surfaces for 10 s. 9 samples per group were tested on peptidemodified surfaces and 3 samples per group were tested on activated surfaces.



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Plain and modified surfaces (CPTES, CPTES/RGD and CPTES/ RGD/S) were chemically characterized with an EPS SSX 100 (Surface Science Instruments, USA) spectrometer (Al Ka X-ray, 1 mm2 spot size, 200 W, 358 take-off angle). The neutralizer was used for analyzing the non-conductive ZrO2 samples. 4 survey-scans per sample were recorded (0–1100 eV) at 1 eV step-size. Peak fitting and quantification were performed using ESCA 2005 (Surface Science Instruments, USA) software provided with the XPS system.



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KKKGGGGRGDS peptides were conjugated with a greenfluorescent probe to directly visualize the organic coatings on the surfaces.30 The probe 5-FAM cadaverine reacts with – COOH groups in the oligopeptides, leaving the –NH2 groups available for attaching at CPTES-silanized surfaces. EDC/NHS crosslinking of carboxylates with primary amines in the peptides was performed with the NHS and Sulfo-NHS kit (Thermo scientific) following manufacturer’s instructions. Briefly, RGD-containing peptides (1 mg) were dissolved in activation buffer (0.1 MES, 0.5 M NaCl; pH 6.0). 2 mM EDC and 5.0 mM NHS were subsequently added to the peptide solution for 15 min (room temperature). Fluorescent 5-FAM cadaverine (1.5 mM) was added to the solution for conjugating with

Surface micro-hardness

Apparent water contact angles (CA)

X-ray photoelectron spectroscopy (XPS)

Fluorescent-labelling of RDG oligopeptides

oligopeptides (2 h; pH 7.2–7.5. The non-reacted reagents were removed by dialysis through a cellulose acetate membrane (MWCO: 500–1000 Da) in PBS-buffer overnight. The product collected was firstly frozen (80 8C, 48 h) and then lyophilized for 24 h using a Freeze Dryer 4.5 (Labconco, US). The fluorescent oligopeptides were immobilized on the studied surfaces following the aforementioned procedures. Micrographs of the labelled surfaces were taken using a fluorescence microscope (Eclipse E800; Nikon, Japan) with b-2E/C FITC fluorescence filter.

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Cellular in vitro tests



Cell Culture

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Murine pre-osteoblasts (MC3T3-E1; ATCC, Manassas, VA, USA) were cultured in a-MEM minimum essential medium supplemented with 10% FBS and 1% penicillin/streptomycin in routine passaging. Cells were stored at 37 8C in a 5% CO2 humidified incubator. Media were changed every 2–3 days. Sub-confluent cells were detached using 0.05% trypsin-HBSS (Hanks Buffered Saline Solution, Sigma Chemicals) and seeded at a density of 104 cell/ml on each sample. Samples were immersed in ethanol (70%, v/v), sterilized with UV light and PBS-rinsed before in vitro tests. For cell adhesion experiments serum-free medium was used.


Cell adhesion

Triple-immunostaining of cells at 4, 6 and 18 h of cell culturing allowed to study their adhesion on the different non-treated (ZrO2, Z-75Ti and Ti) and modified (CPTES, CPTES/RGD, CPTES/RGD/S and RGD) surfaces. At each experimental time, cells were fixed (4% PFA; 20 min), permeabilized (0.25% Triton X-100; 5 min) and blocked within 6% BSA-FBS solution (1 h). Vinculin in focal adhesions was labelled using Anti-vinculin (1:2000) and Alexa Fluor 488 secondary antibody (1:500), F-actin was labelled with rhodamine phalloidin (1:500) and cellular nucleus with DAPI (1:1000). At least 3 images per sample on 5 different samples per group were taken on randomly-selected locations using a fluorescence microscope (E600, Nikon, Japan) and processed with Cell F (Olympus, Japan) software. The number of adhered cells per image field (10) was counted using appropriate image analysis software (ImageJ, http://rsbweb.nih.gov/ij/, NIH, USA). This was also used to analyze cell spreading and focal adhesion points (20). In particular, differences in cellular spreading on the coated substrates were determined by measuring average cell area of all cells included in a field view (n = 6–37 cells).

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Statistical analysis

Data were presented as mean value with the standard error of the mean (SEM) or standard deviation (SD) (as indicated in each figure caption). Significant differences were assessed by ANOVA Tables with multiple comparisons post hoc: TukeyHSD or Dunnett’s C for comparing groups with assumed and not-assumed equality of variances (Levene’s test, p < 0.05), respectively. Statistical analyses were performed with SPSS statistics v20 (IBM, NY, USA) at two significance levels (p < 0.05, p < 0.01).

Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

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Results and discussion

Table 2 – Apparent water contact angles (CA) on the untreated (plain) and activated surfaces. n = 3 samples; mean W SD.


Surface properties of activated surfaces

Activation treatment


Surface visualization

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Fig. 1 displays SEM micrographs of plain, NaOH-etched and plasma-treated surfaces of ZrO2, Z-75Ti and Ti samples. After NaOH-etching, materials showed alterations of their original surface topography (ZrO2/NaOH and Ti/NaOH) or structure (Z-75Ti/NaOH). Notable topographical or microstructural differences were not found between plain substrates and those subjected to the diverse plasma treatments; except in the case of the cermets treated with PT2 (O2plasma for 10-min). PT2 treatment preferentially oxidized the cermet at the interface of the two components, as assessed by energy-dispersive X-ray spectroscopy (results not shown).

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Water contact angles

Table 2 shows CA values of DI-water on original (plain) and all activated surfaces. The activated surfaces were more

Plain NaOH-etching Plasma-cleaning PT1 PT2 PT3 PT4


Apparent water contact angle (CA) [8] ZrO2



58.1  8.0 22.5  4.6

64.7  8.4 24.3  4.7

55.2  7.3 31.6  11.4

7.5  3.8 3.6  3.2 7.6  3.9 2.1  1.8

5.3  2.4 5.5  3.0 6.7  2.9 5.0  2.2

2.3  2.0 2.9  2.8 3.7  1.1 4.0  2.7

hydrophilic (lower CA values) than the non-treated surfaces. Hence, the main purpose of the activation process was achieved as a result of removing hydrophobic contaminants that exposed preexisting hydroxyl groups and/or by forming new reactive polar groups on the treated surfaces. The increased hydrophilicity was significantly higher for all

Fig. 1 – SEM micrographs of the topography/microstructure of untreated (first row) and activated by NaOH-etching (second row) or O2-plasma for 10 min, PT2, (third row) surfaces. Microstructures for PT1, PT3 and PT4 are not displayed because they did not show any further difference related to the untreated (plain) ones. Insert is a close-up of the Z-75Ti/OH microstructure showing physical modifications due to oxidative processes during the PT2 treatment (some of them are also marked with arrows). Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

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Fig. 2 – Time course of water CA values on surfaces activated by different conditions of plasma cleaning and stored in vacuum for 3 h. Each curve shows the results for one of each of n = 3 samples per treatment.

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plasma cleaning treatments than for the NaOH-etching treatment on each studied substrate, which suggests that the activation process with plasma cleaning provided with an increasing number of reactive OH polar groups. Since the activation methodology aims to create a consistent reactive surface that enable the posterior chemical functionalization, plasma cleaning is revealed as the most efficient treatment. Thus, we decided to discard the alkaline etching treatment and select the best plasma cleaning process. To do so, we evaluated the ability of the differently-treated plasma activated surfaces to retain their activation/hydrophilic state for at least 3 h in storage under vacuum (Fig. 2). Conditions for PT4 (10 min air-plasma) were optimal as this treatment produced surfaces with water-CA below 208 (high hydrophilicity) for at least 180 min of storage in all activated materials.

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Microhardness, a surface mechanical property, was measured to investigate detrimental effects by the surface treatments on

the mechanical performance of these materials. If the surface of the biomaterial is weakened by the treatments applied, some relevant clinical mechanical properties can be affected, most importantly the fatigue resistance of the material. Also, hardness is of important applicability to dentistry because it prevents wear of a biomaterial by increasing its resistance to abrasion and scratching during application of loads.52 Since hardness can be correlated with a number of mechanical properties, the indentation tests have been traditionally used to evaluate not only hardness but also elastic modulus, creep and fracture of brittle materials.53,54 We compared values of microhardness (Hv) of the untreated (PT0) and plasma-treated (PT1–PT4) substrates (Fig. 3). O2-Plasma treatments (PT1 and PT2) significantly ( p < 0.01) reduced hardness of the materials containing zirconia (ZrO2 and Z-75Ti). Due to its high electron energy these non-thermal plasma cleaning treatments might induce physical chemical changes in the substrates.50 Airplasma did not significantly affect the mechanical properties of any of the tested surfaces. While the O2-saturated chamber

Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

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Fig. 3 – Microhardness of untreated (PT0) and plasmatreated surfaces. n = 60 indentations; error bars are SD; statistically significant differences with respect to untreated groups (PT0) are denoted **p < 0.01 and *p < 0.05.

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in PT1 and PT2 led to chemical oxidation of surfaces containing zirconia, a lower oxygen concentration in air plasma (approx. 21% for environmental values) did not change the mechanical properties of the studied substrates. Overall, applying plasma treatment conditions for PT4 (air plasma for 10 min) resulted in the best set of surface properties obtained for the diverse methods of activation that were investigated. Therefore, we selected those plasma conditions to activate tested and coated substrates in all further experiments.



Surface properties of biofunctionalized surfaces

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Water contact angles

Wettability of the surfaces changed due to modifications of the surface chemistry produced with the different steps of the biofunctionalization process (Fig. 4). Untreated surfaces presented a moderate hydrophilic character (CA = 58.098  8.038, 64.708  8.408 and 55.208  7.308 for ZrO2, Z-75Ti and Ti, respectively). The surfaces were highly hydrophilic after plasma cleaning (PT4) due to the surface reactivity induced by this treatment. Silanization of the activated surfaces notably increased their CA values above the original values for all materials (CA = 63.908  4.608, 70.608  3.108 and 73.008  5.608 for ZrO2/CPTES, Z-75Ti/CPTES and Ti/CPTES,

Fig. 4 – Water contact angles (CA) on the modified surfaces after each step of the biofunctionalization process. n = 9; error bars are SD; symbols indicate statistically significant differences (2 symbols p < 0.01; 1 symbol p < 0.05): (#) compared to OH; (*) compared to CPTES; (*) compared to CPTES/RGD; (^) compared to CPTES/RGD/S; (*) compared to RGD.

respectively) because CPTES has an alkyl chain that confers hydrophobic character to the molecule. When RGD-containing oligopeptides were immobilized on the silanized substrates (CPTES/RGD groups), the biofunctionalized surfaces were significantly more hydrophilic than the corresponding CPTES-silanized surfaces. The oligopeptides used here contained abundant aminoacids with charged side chains (lysines, arginine and aspartate) so they are notably hydrophilic. After ultrasonication the biofunctionalized surfaces of all materials (CPTES/RGD/S groups) retained their wettability state, demonstrating mechanical stability. Moreover, when the RGD-containing oligopeptides were physisorbed on the different materials, the surfaces were more hydrophilic than

Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

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those with peptides immobilized on the silanized surfaces, thus suggesting that the silanized and biofunctionalized substrates showed a combination of the hydrophobic character of the silane molecules and the hydrophilic character of the oligopeptides.

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Chemical composition

Fig. 5 and Table 3 show respectively XPS-spectra and elemental quantification of plain (ZrO2, Z-75Ti and Ti), silanized (CPTES), RGD-peptide biofunctionalized (CPTES/ RGD) and RGD-peptide biofunctionalized and sonicated for 2 h (CPTES/RGD/S) surfaces. All characteristic elemental XPS peaks of each non-treated substrate as well as of the substrates after each step of the functionalization process were identified. Si 2s and Cl 2s peaks were distinctive after silanization on CPTES-samples, indicating that the silanes were successfully retained on the different materials. Titanium and zirconia have very similar points of zero charge (PZC). PZC for zirconia powders has been assessed at pH 4–6,55 and at the PZC for nanocrystalline zirconia is at pH 4.5.56 The latter is the same PZC than the one for titania (TiO2).57,58 These similar electrostatic conditions may account for the parallel results obtained after the silanization process.17 N 1s peak appeared after immobilizing the RGD-containing peptides on the CPTES/RGD specimens. The N 1s peak was also detected in the spectra for biofunctionalized surfaces after sonication. The latter suggested that the peptides were strongly anchored to the surface and thus, the CPTES molecules successfully acted as covalent linkers between the inorganic substrates and the oligopeptides. The higher N% on the biofunctionalized surfaces after ultrasonication can be a consequence of the removal from the surfaces of loosely bound CPTES molecules. This is further confirmed by the decrease of the Cl/Me and Si/Me ratios and/or the increase of the N/Me ratio with respect to the corresponding base metal (Me) for each substrate on the sonicated surfaces (Table 3).


Fluorescent visualization of biofunctional coatings

449 450 451 452 453 454 455 456 457 458 459 460 461 462

Fig. 6 shows fluorescent images of green-labelled KKKGGGGRGDS oligopeptides that were physically (f-RGD) or covalently (CPTES/f-RGD) immobilized on ZrO2, Z-75Ti and Ti surfaces. The images were taken before (f-RGD and CPTES/fRGD) and after ultrasonication in DI-water for 2 h (f-RGD/S and CPTES/f-RGD/S). Unlike f-RGD-surfaces, CPTES/f-RGD-surfaces showed strong green signal after peptide immobilization on the 3-types of materials. Only a small portion of the peptides on the biofunctionalized surfaces was lost after ultrasonication. The so far discussed results confirmed that we successfully obtained a robust coating of covalently-immobilized oligopeptides on ZrO2, Z-75 Ti and Ti surfaces.



464 465 466 467 468

Fig. 5 – XPS-spectra of ZrO2, Z-75Ti and Ti plain and modified surfaces after each step of biofunctionalization process. Characteristic peaks on plain samples: C 1s (285 eV), Ti 2p (460 eV), O 1s (530 eV) and Zr 3d (182 eV).

Cell Adhesion on biofunctionalized surfaces

Fig. 7 shows the number of MC3T3 osteoblasts and fluorescence images of the adhered cells on all modified surfaces after different culture periods and Table 4 summarizes the mean values of cell area of the osteoblasts adhered to the modified substrates at the different periods in culture. Fig. 7

Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

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Table 3 – Surface chemical composition of modified surfaces. Elemental quantification from XPS-survey scans in Fig. 5. Me: metallic content (Zr in zirconia, Ti in titanium, and the sum of Zr and Ti for Z-75Ti substrates). Substrate

Elemental quantification (atom %) O 1s


469 470 471 472 473 474

Ti 2p

N 1s

C 1s

Cl 2p

Zr 3d

Si 2p





22.4 35.3 26.0

– – –

– 3.3 5.1

62.5 45.4 57.4

6.4 3.8 2.8

4.4 7.7 6.3

4.6 4.5 2.3

1.0 0.6 0.4

1.5 0.5 0.4

– 0.4 0.8

– 1.2 0.6

28.8 23.6 36.3

3.2 3.0 3.4

– 2.8 5.3

52.8 63.4 53.4

7.3 2.5 0.0

1.0 1.1 1.6

7.0 3.7 0.0

1.7 0.9 0.0

1.7 0.6 0.0

– 0.7 1.1

– 1.1 0.0

41.5 18.4 37.7

7.4 2.1 5.6

– 1.8 4.3

39.9 71.9 52.6

3.9 1.2 0.0

– – –

7.3 2.9 0.0

1.0 1.4 0.0

0.5 0.6 0.0

– 0.9 0.8

– 0.7 0.0

reveals no significant differences in cell number between glass (positive control) and untreated (ZrO2, Z-75Ti and Ti) surfaces. Silanized substrates (CPTES) significantly reduced cell number (Fig. 7) and cell spreading (Table 4) after all periods of cell culture. Neither significant differences in number of adherent cells (Fig. 7) nor apparent variations in cellular spreading

(Table 4) were assessed between untreated surfaces and surfaces coated with physisorbed RGD-containing peptides. This matched with lower peptide retention on the corresponding fluorescence-labelled coatings both before and after sonication (Fig. 6). On the contrary, chemically biofunctionalized surfaces (CPTES/RGD) significantly increased the number

Fig. 6 – Visualization of the green fluorescently-labelled RGD-coatings on ZrO2, Z-75Ti and Ti surfaces. The labelled peptides were physisorbed ( f-RGD; first and third columns) or covalently-bonded (CPTES/f-RGD; second and fourth columns) on the surfaces. Micrographs were taken before ( f-RGD and CPTES/f-RGD; first and second columns) and after 2 h of ultrasonication ( f-RGD/S and CPTES/f-RGD/S; third and fourth columns). Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

475 476 477 478 479 480

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Fig. 7 – Cell adhesion on modified surfaces after different periods of cell culture. Left: number of adhered cells on the modified surfaces. n = 5; error bars are SEM. Symbols indicate statistically differences (p < 0.05): compared to CPTES (4 h), & compared to CPTES/RGD (4 h), ^ compared to CPTES/RGD/S (4 h), * compared to CPTES/RGD (6 h), ^ compared to CPTES/ RGD/S (6 h), * compared to CPTES (6 h), * compared to CPTES (18 h). Right: fluorescent images of triple-stained cells adhered on the different modified surfaces after the different periods in culture.

Table 4 – Cellular spreading of cells attached on the investigated substrates after different periods in culture. Data are mean W standard error of the mean of cell area (n = 6–35 cells) quantified from fluorescence images of the immunolabeled osteoblasts. Substrate




Surface modification


Time in cell culture (hours) 4



929  82 405  74 1460  179 1626  126 1006  94 1034  195 488  84 1225  151 1398  125 1181  204 1004  105 349  164 1614  164 1503  144 1014  205

1502  183 1373  183 2030  219 2180  305 1577  154 1294  125 905  186 1711  136 1577  168 1272  109 1356  207 1057  146 1973  272 1819  292 1621  297

2178  302 1117  180 2856  399 2671  594 2089  250 1300  121 935  186 2266  254 2298  285 1213  156 2221  240 1040  132 2653  504 2451  558 1645  263

Please cite this article in press as: Fernandez-Garcia E, et al. Peptide-functionalized zirconia and new zirconia/titanium biocermets for dental applications. Journal of Dentistry (2015), http://dx.doi.org/10.1016/j.jdent.2015.06.002

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of adhered cells when compared with the cells adhered on untreated surfaces. The significantly highest number of adhered cells on all biofunctionalized surfaces was reached after 6 h of cell culture. The decreased cell number after 18 h of cell culture may be caused by the absence of serum in the medium which, at those longer periods of culture can compromise cell viability. Moreover, the cellular spreading evaluated from measurements of cell area increased over time on all modified substrates but, most notably, on biofunctionalized (CPTES/RGD and CPTES/RGD/S) surfaces for all materials (Table 4). On the fluorescent images, the number of focal adhesion points and average cell area on the biofunctionalized substrates before and after mechanical challenge was notably higher than on all other surfaces at all periods of cell culture. Proteins (laminin, fibronectin, vitronectin) of the extracellular matrix (ECM) of bone have the well-known RGDS motif. This motif is recognized by cell membrane receptors12; providing binding sites for integrins that facilitate the attachment and growth of osteoblasts, both in vitro and in vivo.21,22,26 Our results demonstrated that the biofunctionalized surfaces with covalently-bonded oligopeptides displayed the expected bioactivity on all different materials; that is, increasing cell adhesion and spreading by cell recognition of the RGDS motif on the purposely-designed oligopeptidecoated surfaces. Thus, the RGDS-oligopeptides were a successful model to validate our biofunctionalization process to obtain stable oligopeptide coatings on materials based on 3Y-TZP.



511 512 513 514 515 516 517 518 519 520 521

We have demonstrated physical, chemical, and biological activation of biofunctionalized zirconia and zirconia/titanium (75% vol. Ti) biomaterials using a technology that consists of (1) 10-min air-plasma surface activation; (2) CPTES silanization; and (3) peptide coating. The ability of functionalizing zirconia and other materials containing zirconia for dental implant applications can enable their improved clinical performance by incorporating signalling oligopeptides to accelerate osseointegration, improve permucosal sealing, and/or incorporate antimicrobial properties to prevent peri-implant infections.


522 523 524 525 526 527 528 529 530 531 532 533 534 535

Acknowledgments Q3 This work was supported by the Office of the Vicepresident for

Research at the University of Minnesota (Project #55466 of the Grant-in-Aid of Research, Artistry and Scholarship Program), the Spanish Ministry of Economy and Competitiveness (MINECO) under the project [MAT2012-38645]; the Spanish Ministry of Science and Innovation (MICINN) under the projects [MAT2009-14542-C02-02 and SAF2011-27863] and [CONSOLIDER CSD2009]; and the Government of the Principality of Asturias through PCTI. Parts of this work were carried out in the University of Minnesota I.T. Characterization Facility, which receives partial support from NSF through the MRSEC program.



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titanium biocermets for dental applications.

Titanium materials have been functionalized with biomolecules as a modern strategy to incorporate bioactive motifs that will expand and improve their ...
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