Physicochemical and in-vitro biological analysis of bio-functionalized titanium samples in a protein-rich medium

Journal of the Mechanical Behavior of Biomedical Materials 96 (2019) 152–164

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Physicochemical and in-vitro biological analysis of bio-functionalized titanium samples in a protein-rich medium

T

Shradha Raoa, Sarah Hashemi Astanehb, Jose Villanuevac, Filipe Silvae, Christos Takoudisb,d, ⁎ Divya Bijukumara, Júlio C.M. Souzaa,f, Mathew T. Mathewa,d, a

Department of Biomedical Science, University of Illinois College of Medicine at Rockford, USA Department of Chemical Engineering, the University of Illinois at Chicago, (UIC), Chicago, USA c Department of Restorative Dentistry, College of Dentistry, UIC, Chicago, IL, USA d Department of Bioengineering, University of Illinois at Chicago (UIC), Chicago, USA e Center for MicroElectromechanical Systems (CMEMS-UMINHO), Universidade do Minho, Portugal f Department of Dental Sciences, University Institute of Health Sciences (IUCS-CESPU), Grandra 4585-116, Portugal b

A R T I C LE I N FO

A B S T R A C T

Keywords: Osseointegration Titanium implants Fibrinogen Surface modification

The long-term survivability of the implants is strongly influenced by the osseointegration aspects of the metalbone interface. In this study, biological materials such as fibrinogen and fibrin are used to functionalize titanium surfaces to enhance the ability of implants to interact with human tissues for accelerated osseointegration. The biofunctionalized samples that were assessed by White Light Microscope, Scanning Electron Microscope and Water Contact Angle for surface properties proved samples etched with HF/HNO3 to be better than HCl/H2SO4 in terms of having optimum roughness and hydrophilicity for our further experiments. To further investigate the in vitro osseointegration of the biofunctionalized samples, Osteoblasts were cultured on the surfaces to assess cell proliferation, adhesion, gene expression as well as the mineralization process. Further bacterial adhesion (Enterococcus faecalis) and electrochemical evaluation of surface coating stability were carried out. Results of the study show that the biofunctionalized surfaces provided high cell proliferation, adherence, gene expression, and mineralization compared to other control surfaces hence proving them to have efficient and enhanced osseointegration. Also, bacterial adhesion studies show that there is no augmented growth of bacteria on the biofunctionalized samples in comparison to control surfaces. Electrochemical studies proved the existence of a stable protein layer on the bio functionalized surfaces. Such a method can reduce the time for osseointegration that can decrease risks in early failures of implants due to its enhanced hydrophilicity and cytocompatibility.

1. Introduction

Various metals used in the manufacturing of medical implants are Aluminum (Al), stainless steel (SS), Cobalt (Co), Chromium (Cr), Molybdenum (Mo) alloys, and Titanium (Ti) [3]. Titanium is extensively used due to their excellent corrosion resistance, biocompatibility, and mechanical strength when compared to the other metallic implants [4]. Commercially pure titanium that is available in four grades depending upon the amount of carbon, nitrogen, oxygen, and iron content are used to manufacture a variety of medical dental implants [5]. Oxidation of titanium occurs when titanium comes in contact with room temperature air and normal tissue fluids [6]. This condition minimizes biocorrosion phenomena and hence is favorable for dental implant devices [7]. Ti is not a bioactive material. Hence surface modification of the metal implant was considered as an efficient method to reduce the time required for osseointegration. The

The artificial devices that are placed inside or on the surface of the body are called medial implants. The major reason for prosthetics is intended to replace missing body parts. Metallic implants are used by the clinicians to treat bone defects, imperfections and fractures [1]. Over the past years, the usage of implants has grown dramatically driven by the increasing population in the developed countries and the need of the patients to maintain the activity of good quality of life. The unique challenges addressed in cardiology, vascular therapy, orthopedics, trauma, spine, dental and wound care is addressed by high-performance implantable biomaterials [2]. Hence researches focusing on improving the performance of the biomedical implants has been constantly increasing [1].

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Corresponding author at: Department of Biomedical Science, University of Illinois College of Medicine at Rockford, IL 61107, USA. E-mail address: [email protected] (M.T. Mathew).

https://doi.org/10.1016/j.jmbbm.2019.03.019 Received 7 December 2018; Received in revised form 17 March 2019; Accepted 19 March 2019 Available online 13 April 2019 1751-6161/ © 2019 Elsevier Ltd. All rights reserved.

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osteoblast integrin receptors to these bioactive molecules happen, stimulation of interaction of cells with their extracellular matrices (ECM) are triggered and cell survival, proliferation, differentiation, and mineralization are promoted [25]. When a foreign body enters the human system, it first comes in contact with the present biological fluids and its surface is quickly covered by proteins. Different proteins in the extracellular matrix like the albumin, collagen, elastin, and fibrinogen are present; but Human Plasma Fibrinogen (HPF) is the only protein that adsorbs rapidly onto any thrombogenic surface [26]. It is also estimated that the concentration of fibrinogen in the blood plasma is around 1.8–3.5 mg/ml [27]. To promote adhesion between cells and biomaterials, fibrinogen is used. Fibrinogen and fibronectin have the arginine-glycine-aspartic acid (RGD) motif, that binds effectively on the surface and hence these proteins have higher adherence to cells than any other proteins in the ECM. Hence there have been recent researches on the adsorption of fibrinogen on metals and its role in improving osseointegration. It is studied that the conversion of soluble fibrinogen to the insoluble clotforming fibrin is the terminal stage of blood clotting [28]. This is enabled by the addition of thrombin that clots fibrinogen to form fibrin. Fibrin has remarkable advantages over biomaterials, which makes it an ideal candidate for biomedical engineering applications. A temporary structure is provided by fibrin that facilitates cellular activities and also aids in deposition of a new extracellular matrix (ECM) that will enhance the osseointegration at bone-metal interface [29]. Despite high success rates of implants, implant fixture failures may occur due to the inadequacy of the host tissue to establish or maintain osseointegration with the biological surrounding. The major problem is related to the time of osseointegration. If there is an increase in the time taken for achieving osseointegration, the risk of infection also increases parallelly [12]. Several complications are linked with metal implants such as implant loosening and infections [15]. Thus, these cases remain a great challenge in dental implantology and to accelerate osseointegration after placement of implants, bioactive surface modifications are necessary. The main aim of this study is to bio-functionalize titanium surfaces in a medium enriched with fibrinogen and check for the in vitro osseointegrational behavior and bacterial adhesion property. This project also aims at modifying surfaces to beneficially alter the morphology, hydrophilicity, surface chemistry by focusing on dental implant applications which enhances the process of osseointegration.

osseointegration of titanium dental implants with the bone is critically dependent on the implant surface properties. Osseointegration can be defined as the direct structural and functional connection between living bone and the surface of a load-bearing artificial implant without the presence of a connective tissue at the bone-implant contact [8,9]. When an osseous interface is reproducibly formed and maintained at the surface of titanium of the load bearing implants, this marks the main clinical advantage of osseointegration [10]. However, mechanical and biological factors such as infection, tissue inflammation, fractures, peri-implant bone resorption and loosening of the implant, pain, and allergic responses could cause early and late complications in orthopedic and dental implants [11–13]. Within the first few months, 1–2% of patients are affected by primary implant failure due to insufficient osseointegration [14]. In about 5% of patients, secondary implant failure develops several years after successful osseointegration and is commonly caused by peri-implantitis [14,15]. Characteristic of an implant is the risk of infection and inflammation, which can significantly undermine the performance of an implant [16] and lead to a significant loss of tissue in the proximity of the implant [17]. For load-bearing implants, the transfer of forces may be interfered by septic or aseptic loosening of the implant, such as the incorrect transfer of the biting force to the dental implant and the surrounding bone, leading to mechanical failure once the fatigue strength of the dental implant is reached [2]. Hence, to accelerate osseointegration and to ensure a long-term bone-to-implant contact without substantial marginal bone loss, the aim of designing new bioactive surface properties and early loading protocols are studied. Roughly 1300 different biomedical implant systems in both orthopedics and dentistry exist varying in shape, dimension, bulk and surface material, thread design, implant-abutment connection, surface topography, surface chemistry, wettability, and surface modification [18]. The shapes of the implants are commonly cylindrical or tapered [19]. Topography, wettability, and coatings are the surface characteristics that contribute to the biological processes during osseointegration [20] by mediating the direct interaction to host osteoblasts in bone formation. Surface modifications can decrease the risks of early failures in dental implants provided by enhanced osseointegration [15]. Several methods of surface modifications have therefore been reported to improve the integration of implants with that of the bone. Dry process coating, electrochemical coatings of Hydroxyapatite (HA), TiO2 coating; CaTiO3 coating; micro-arc oxidation; surface modified layer formation by chemical treatment are some of the most commonly used techniques for surface modifications [21]. Acid etching is also one such technique done to modify the surface of metals by increasing the surface roughness using strong acids. Acid etching not only cleans the surface of implants, whereas it also creates a micro-textured surface forming a stable H+ layer on the surface due to the overexposure to acids. This H+ layer is used to attract negatively charged biomolecules, mainly proteins which in turn attracts more bone cells towards itself, thus increasing the osseointegration of the metal with the bone [21,22]. Hence, to increase the surface area and osseointegration potential of implants, most modern dental implants have a textured surface (through etching, anodic oxidation or various-media blasting) [23]. The surface roughness as being in the range of 1–10 µm is the defined range for the microtopographic profile of dental implants [24]. This range of roughness has proven to maximize the interlocking between mineralized bone and the surface of the implant. The process of osseointegration can be altered and molecular and cellular activities can be influenced by these interactions [10,14]. Besides optimizing the implant surface properties based on topographical and chemical modifications, much attention has been given to implant surface treatment with biologically active substances [22]. Purified proteins and synthetic peptides are used as biologically active substances that mediate the binding of osteoblast cell adhesion receptors (e.g. integrins, selectins and cadherins). Once the binding of the

2. Materials and methods 2.1. Sample preparation and acid etching Cylindrical samples with dimensions of 8 mm in diameter and 3 mm in thickness were cut from commercially pure titanium (cp Ti) grade IV rods (Metroohm Autolab, USA). To maintain surface uniformity, the cut samples were polished with Silicon Carbide (SiC) paper of mesh size 600. The polished samples were cleaned in isopropyl alcohol for 10 min and in distilled water for 5 min using an ultrasonic bath. Samples were then stored in a desiccator for performing further experiments [4]. Samples were divided into two main groups for surface modification considering two acid-etching methods: (AT1 and AT2). The samples of test group AT1 were immersed in an acidic solution containing equal volumes of 10% HF (by concentration) and 30% HNO3 (by concentration) at 60 °C for 10 min [30] while samples of test group AT2 were immersed in a solution composed of equal volumes of 48% H2SO4 (by concentration) and 18% HCl (by concentration) at 60 °C for 60 min [31]. The purity of the acids was HF: 100%, HNO3: 90%, HCl: 70% and H2SO4: 99.99%. The control group samples were immersed in distilled water at 60 °C for 10 min.

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2.2. Preparation of fibrinogen for biofunctionalization

To further check the intensity of the protein adsorbed on the surfaces, the washed-out fibrinogen and fibrin using PBS were collected and quantified using BCA assay to check and estimate the adsorption of protein on the surfaces. BCA assay is a standard bicinchoninic acid assay that is done to check the concentration of protein by the pierce BCA assay kit from thermofisher. A standard curve is plotted for albumin and with that curve as base, the concentration of the unknown can be measured with Reagents A, B and C at the ratio (25:24:1) at 562 nm.

Fibrinogen gel was prepared as previously described in previous studies [32,33]. Briefly, in Tris buffer (10 mM Tris/HCl, pH 7.4) plasminogen-rich human fibrinogen (Sigma Aldrich, USA) was dissolved to a concentration of 5 mg/mL. Saturated (NH4)2SO4 was slowly added to a final concentration of 19% (v/v) and the solution was mixed for 30 min at room temperature prior to centrifugation for 10 min at 2000×g. For obtaining High Molecular Weight (HMW) fibrinogen pellet (~ 99 % purity), repetition of this precipitation step was carried out, which was dissolved in 5 mL of 0.9 %( v/v) NaCl and then dialyzed against M199 culture (Sigma Aldrich, USA) medium. This step included dialyzing the protein in ultra-centrifugal Millipore with equal amounts of M199 culture media. Fibrinogen was centrifuged at 4000×g and 8 °C for 25 min three times. The fibrinogen gels were stored in single-use aliquots at -80 °C until further use. Titanium samples were divided into two subgroups (F and FT) considering coating with fibrinogen clotted with thrombin (FT group) and not (F group). On F groups, surfaces were etched and coated with 55 μL HMW fibrinogen (2 mg/mL) while fibrinogen was clotted on etched FT group with 0.5 mg/mL of thrombin dissolved in a 4.5 mM calcium chloride buffer in an 8.5:1 ratio. The concentrations of the proteins was selected based on the previous report [34]. Prior to physiochemical and biological assessment, titanium samples coated and not with fibrinogen gels were incubated at 6 °C for 18 h to allow completion of fibrin gelling [34]. After 24 h, the protein was washed with PBS and air dried. These samples were coated and further considered for surface characteristic studies.

2.4. Protein stability on the surfaces Further, to evaluate the stability of the adhered protein on the surfaces, an electrochemical method is used. It should be noted that, these tests are not to study the corrosion properties of the surfaces, however, the stable protein film would provide increased corrosion resistance, which could be reflected with electrochemical data as a function of exposure time. A corrosion cell with a three-electrode set-up as per ASTM standard (G 60), connected to a Gamry made potentiostat (Interface 1000) was used in this study. The working electrode (WE) consisted of a titanium sample, the reference electrode (RE) was a saturated calomel electrode (SCE), and the counter electrode (CE) was a graphite rod. The titanium specimens were categorized into acid etched, fibrinogen and fibrin coated samples, with an exposed area of 0.50 cm [2]. The electrolyte used here was 0.9% NaCl. The specific testing protocol used in this experiment consisted of alternating: 1) 3600 s of Open circuit potential (OCP/Eoc) test, 2) Electrochemical Impedance Spectroscopy (EIS) test at Eoc and with an amplitude of ± 10 mV and frequency range from 100 kHz to 0.001 Hz lasting for 1298 s for 3 cycles of each OCP and EIS. Last three cycles, the OCP period was extended around 10 hrs each with EIS in between the OCP tests. The evolution potential as a function time is monitored during the OCP test. EIS data was analyzed using the Bode and Nyquist plot and an equivalent circuit was used (modified Randel’s circuit with CPE) to determine the resistance (R) and capacitance (C). Fig. 1 shows the protocol used for this experiment.

2.3. Physiochemical characterization of the surfaces The biofunctionalized surfaces were characterized by the following methods. 2.3.1. 3D profiles of the surface (white-light interferometry microscope) The topography of the roughened samples was observed through three-dimensional (3D) metrology using a white light microscope (Nano Contour GT-K, Bruker, Cheryl Pkwy, Fitchburg, WI) over an evaluated area of the samples. Vision 64 was the software used to scan and evaluate the surface topography of the samples. The results were obtained using a filter which determined the roughness. Head of the profilometer was lowered till a few mm above-treated surface and was slowly raised until the surface was focused on, they were then scanned, and a filter was used to assess roughness and remove any tilt from the sample [35,36].

2.5. Cell culture Human Osteoblasts (MG63) were used in this study. Cells were cultured at 37 °C in an atmosphere of 5% CO2 and 100% humidity in culture medium composed of Modified Eagle's Medium with 10% Fetal Bovine serum and Antibiotic (1:100) in 25 cm2 flasks (Corning, Corning NY, USA) until reaching 80-85% confluency. Cells were then detached using trypsin-EDTA, centrifuged at 3000 rpm and re-suspended in culture media at an adequate density for the biological assays. All experiments were conducted at 37 °C.

2.3.2. Scanning Electron Microscopy To analyze the morphology of the surfaces using SEM, samples coated with protein (F and FT) were washed twice in PBS and fixed in 2% glutaraldehyde for 5 min. Then the surfaces were washed thrice in PBS and dehydrated through a series of graded ethanol solutions (50,70, 80, 90, 100%) [37]. For morphological inspection, samples were sputter-coated with Osmium tetroxide and analyzed by FieldEmission Scanning Electron Microscopy (HITACHI, SU8030 Ultra Highresolution Scanning Electron Microscope, Japan) at 5–8 kV at an angle of 60° [38,39].

2.6. In vitro analysis of the surfaces for osseointegration To check for the cell viability on different surfaces, the samples were cleaned and exposed to UV radiation for around 10 h and kept on 24 well plates. Cells were counted and dropped on the sample surfaces and allowed to grow with a periodic change in media. 10% of Alamar blue was added to the media on days 2, 3, 5, 10, 14. After 24 h, the media

2.3.3. Wettability measurements To check for the wettability of surfaces, a sessile drop method was used for carrying out water contact angle (WCA) measurements. About 2 µl of distilled water was placed on the surface and the water absorption capacity was inspected by a video-based drop shape analyzer (Ramè-Hart, NRL, C.A. goniometer, model no: 100-00). Image J software was used to analyze the angle measurements in detail. Three measurements on each sample and three samples for each material were assessed [40,41].

Fig. 1. Experimental protocol for analyzing the stability of protein coating on the surfaces. Protocol consists of 6 OCPs and 6 EIS. First three OCPs consisted of 1.16 h each and last three OCPs consisted of 10.16 h each followed by 40 min of EIS after every OCP, altogether accounting for approximately 39 h. 154

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Fig. 2. Morphology of Ti surfaces. Microscopic analysis of the samples (A) Polished (control) at 100X and 1000X, (B) etched in HF/HNO3 (AT1) at 100X and 1000 X (C) etched in HCl/H2SO4 (AT2) at 100X and 1000X. (A) Shows uniformity on the surfaces due to polishing while figures (B) and (C) show the irregularities on the surfaces due to treatment of the samples to acid etching. (n=3).

and then converted into cDNA by the process of reverse transcription. This cDNA was used to check for the GAPDH amplification as GAPDH served as the control. Also, cDNA was used to perform qPCR analysis with the mentioned genes above [46,47].

from the samples were collected and analyzed under 570–600 nm. The graphs were plotted and checked for cells proliferation [42,43]. Further to check and confirm for the adherence of the cells on the surfaces under microscopes, cells were counted and dropped on the sample surfaces and allowed to grow with a periodic change in media for analysis under SEM. On days 1 and 5, the samples were removed from the wells and washed once in PBS. The cells were fixed with 2.5% glutaraldehyde for 15 min. Then the surfaces were washed thrice in PBS and dehydrated through a series of graded ethanol solutions (50%, 70%, 80%, 90%, 100%) [37]. The surfaces were then analyzed under SEM. For analyzing the under a confocal microscope, the samples were cleaned and exposed to UV radiation for around 10 h and kept on 24 well plates. Cells were counted and seeded on the sample surfaces and allowed to grow with a periodic change in media. On days 1 and 5, the samples were removed from the wells and washed once in PBS. The surface was then fixed for 5 min in 3.7% formaldehyde solution in PBS and washed in PBS. The cells were then permeabilized with 0.1% Triton X-100 in PBS and washed again in PBS. Cells were stained with a 50 µg/ ml fluorescent phalloidin conjugate solution in PBS for 40 min at room temperature. The surfaces were washed with PBS and stained with DAPI for 5 min and then stored in dark. The surfaces were then analyzed under the confocal microscope [44,45].

2.8. Mineralization assay The samples were cleaned and exposed to UV radiation for around 10 h and kept on 24 well plates. Cells were counted and dropped on the sample surfaces and allowed to grow with a periodic change in media for a period of 21 days. After 21 days, the media was discarded, and alizarin red stain was added on the surface of the samples and kept for 10 min. After 10 min the stain was washed with PBS and images of the surfaces were taken. The staining on the samples was then quantified by adding 10% acetic acid to the samples and transferred to 96 well plates and was analyzed under 405 nm [48,49]. 2.9. Bacterial activity of E. faecalis For the biofilm assays on the surfaces, the samples were cleaned and exposed to UV radiation for around 10 h and kept on 24 well plates E. faecalis culture and TSB containing 1% glucose at a ratio of 1:100 were added to different modified samples placed on 12 well plates and allowed to form a biofilm over a span of 7 days with periodic change in the TSB media. After incubation for 24 h, the supernatant was removed, and the biofilm formed was gently rinsed in PBS once and resuspended in 1 ml of saline. Alamar blue assay was then performed to these samples at 560-600nm to find out the viability of bacteria on different surfaces [50].

2.7. Expression of osteoblast-specific genes qPCR analysis was done to osteoblast-specific genes Alkaline phosphatase, Collagen, Osteocalcin, Bone morphogenetic protein-2, and transcription factor RUNX2 to analyze its expression on the surfaces on day 7 and day 14. The samples were cleaned and exposed to UV radiation for around 10 h and kept on 24 well plates. Cells were counted and dropped on the sample surfaces and allowed to grow with a periodic change in media. On days 7 and 14, the media was discarded and Trizol reagent was added to each group of samples. The RNA was extracted and stored at −80 °C for 24 h. The extracted RNA was thawed

3. Results 3.1. Acid etching of titanium samples The surfaces of polished, and etched samples were inspected by 155

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Fig. 3. (a) 2D and 3D analysis of the surfaces: (i) indicates the roughness of the polished surfaces with no or very less surface irregularities. (ii) indicates the roughness of the surfaces etched in Hcl/H2SO4 showing surface irregularities represented by red marks on the surface. (iii) indicates the roughness of the surfaces etched in HF/HNO3 showing surface roughness represented by red and blue marks on the surface. (n=3). (b) Graphical representation of the surface roughness estimation of various modified surfaces as mentioned. There is a significant surface roughness increase in the surfaces that were acid etched compared to that of the control surfaces (*P value < 0.05; n=3).

number of peaks while it was the least for the control samples which were mirror polished. Samples etched in HF/HNO3 had the optimum amount of roughness with both peaks and valleys. Fig. 3b depicts a bar graph representation of the surface roughness from the 3D light microscopic images. The roughness of the surfaces that were polished was 0.15 µm, surfaces etched in HCl/H2SO4 were 1.04 µm and surfaces etched in HF/HNO3 were 0.44 µm. Clearly, the surfaces etched in HCl/ H2SO4 had higher roughness and control samples had the least roughness. Further the roughness of fibrinogen coating were 0.89 µm and 1.1 µm in HCl/H2SO4 and HF/HNO3 surfaces respectively. Whereas the average roughness decreases to 0.34 µm and 0.52 µm in HCl/H2SO4 and HF/HNO3 surfaces respectively. There is no significant difference in roughness observed between HCl/H2SO4 and HF/HNO3 groups with fibrinogen and fibrin coating. FEGSEM showed the magnified view of the surfaces that were etched and coated with fibrinogen and fibrin. Fig. 4 shows the images of different surfaces analyzed under Scanning Electron Microscope. The control group is the surfaces that are etched in the respective acid groups without any coating. Fibrinogen and fibrin coated surfaces are

optical microscopy (Unitron MEC 3, USA) as seen in Fig. 2. Samples from control group revealed surfaces with multi-directional grooves as a result of the cutting process (Fig. 2A). The irregular patterns noticed on the surface of the titanium samples after etching indicated that there was a change in roughness when compared to the surfaces of the control group. Differences between the surfaces were noticeable between the test groups considering the different etching processes. Samples subjected to acid treatment classified as AT1 showed a homogenous distribution of irregularities throughout the surface (Fig. 2B). AT2 samples showed a surface with a higher number of the valley- like irregularities (Fig. 2C). 3.2. Physiochemical characterization of surfaces The surfaces of the polished, and etched samples were analyzed for roughness under a 3D white light microscope. The 3D and 2D images of the surface characteristics are analyzed and depicted in Fig. 3a. The red color depicts high peaks and the blue represents low valleys. The picture also shows that the samples etched in HCl/H2SO4 had the highest 156

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Fig. 4. SEM images of surfaces. (a) indicates the surfaces etched in HCl/ H2SO4 while (b) indicates the surfaces etched in HF/HNO3. (a)i indicates irregularities on the surfaces without any coating while (a)ii and (a)iii indicates coating of fibrinogen and fibrin molecules on the etched surfaces. Similar pattern can be observed for (b)i, (b)ii and (b)iii. Therefore, through these images, coating of protein layer can be observed on the surfaces. (n=3).

to the surfaces. Thus, as the contact angle is high, the surfaces are more hydrophobic while the contact angle is low, the surfaces are more hydrophilic making it better sustainable in the biological environment. Since both optimal roughness and wettability was observed on surfaces that were etched in HF/HNO3, these surfaces were considered further for further tests.

depicted in the adjacent images. In control groups, the irregularities on the surfaces can be clearly seen, fibrinogen-coated surfaces have a layer of fibrinogen adhered on the surface while fibrin coated samples have globular fibrin molecules attached on the etched surfaces. Water contact angle is used to generate the angle of water adhesion in contact with the respective surfaces. It was found that the contact angle was greater for surfaces etched in HCl/H2SO4 while it was less for surfaces etched in HF/HNO3 proving these surfaces to be hydrophilic than the surfaces etched in HCl/H2SO4. Fig. 5a shows the measurements of water contact analyzed on different surfaces and Fig. 5b is a graphical representation of the values of the contact angles with respect

3.3. Protein stability on the surfaces The Open Circuit Potential (OCP) and Electrochemical Impedance Spectroscopy (EIS) data were analyzed. The principle of OCP is that 157

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Fig. 5. Surface hydrophilicity. (a) Images indicating contact angle of water placed on surfaces where we can analyze the adsorption rate of water on the surfaces (b) Bar graph showing mean contact angles for different surfaces indicating the hydrophilic and hydrophobic nature of the surfaces. We can observe a significant decrease in water contact angle on the surfaces that were etched and coated with protein. (*P value < 0.05; **P-value < 0.01; n=3).

resistance has decreased for the surfaces that were just etched vindicating the unstable nature of the etched surface. The adsorption intensity of the protein on the surfaces was found out by BCA assay. The graph is shown in Fig. 7 depicts the adsorption of proteins on control and etched surfaces in HF/HNO3 conditions. The adsorption of protein was found to be maximum for etched samples coated with fibrin gel and least for the control samples coated with fibrinogen.

with the increase in the potential and maintenance of its consistency with the time, there is the better resistance of the surface to corrosion tendency. Therefore, it can be estimated that with the increase in potential and time, the protein coating is stable and resistant to corrosion. In Fig. 6a, the OCP curves of etched (control), fibrinogen and fibrin coated samples can be seen. There is an increase in the potential of the curves of fibrinogen and fibrin coated samples which indicate the stability of the protein coating for a period of 39 h while the potential of the etched samples is the least resistant to corrosion. The OCP shows a slight reduction for fibrinogen group, possibly due to an environmental changes that affect the electrochemical kinetics and stability of the protein layer. The EIS data of the surfaces were modeled and the resistance of the coating and capacitance were analyzed for a series of 6 EISs and at a total period of 39 h. The purpose of the study to compare the stability of the three surfaces, hence an EIS model suitable for the three groups were used. The goodness of fit values aimed for < 0.001, however in some cases, particularly in the case of fibrin group, it was < 0.1. The increase in resistance (Rp) and a decrease of capacitance (Yo) through the EIS indicates the increase in the stability of the coating of the surfaces. In Fig. 6b it can be seen that the resistance has increased, and capacitance has decreased through the EIS in the case of fibrinogen and fibrin coated surfaces indicating a stable protein layer while the

3.4. In vitro analysis of surfaces for osseointegration Alamar blue proliferative assay showed that cells were most proliferative on the surfaces coated with protein compared to that on the surfaces of etched and control in all days- 2, 3, 5, 10, 14. This is represented as a bar graph in Fig. 8. This shows that the surfaces etched and coated with protein are better for cell proliferation than polished and etched surfaces. SEM images of cells attached on the surfaces were taken on day 1 and day 5 as shown in Fig. 9a. As seen, the cells were found least attached on the surfaces that were polished, more on etched surfaces and maximum on biofunctionalized surfaces. This shows that the surfaces coated with protein are more biocompatible. To further check the surfaces, confocal microscopic images shows 158

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Fig. 6. a: OCP analysis of surfaces. The OCP graph shows the increase in resistance of fibrin and fibrinogen-coated surfaces towards corrosion due to the increase in potential indicating the stability of the protein layer. (n=3). b: EIS analysis of surfaces. (i) The EIS graphs show the decrease in resistance and a decreasing trend in capacitance indicating instability of the surfaces of etched samples. (ii) and (iii) The EIS graphs show the increase in resistance and decrease in capacitance for the surfaces etched and coated with fibrin and fibrinogen over the period of 6 EIS indicating the stability of the protein layer on the surfaces when exposed to saline. Each EIS started from 100,000 Hz to 0.01 Hz for a total period of approximately 39 h. (n=3).

biofunctionalized surfaces on day 14; for BGLAP there was a significant 2.5 fold and 4 fold increase of gene on the biofunctionalized surfaces on day 14; for BMP2 there was a significant 4 fold and 5 fold increase of gene on the biofunctionalized surfaces on day 14; for ALP there was a significant 3 fold and 6.5 fold increase of gene on the biofunctionalized surfaces on day 14 and for COL there was a significant 1.5 fold and 2 fold increase of gene on the biofunctionalized surfaces on day 14 in comparison to control surfaces. Overall, the upregulation of the genes and transcription factor was found in all surfaces that were coated with fibrin compared to other surfaces. The graphs indicate the fold increase compared to that of the control group.

staining of DAPI and rhodamine phalloidin in Fig. 9b. As seen, the cells were found least attached on the surfaces that were polished, more on etched surfaces and maximum on etched surfaces coated with protein. This shows that the surfaces coated with protein are more biocompatible.

3.5. qPCR- expression of osteoblast-specific genes Fig. 10 shows the expression of different genes ALP, COL, BGLAP, BMP-2, and transcription factor RUNX2 on days 7 and 14 of cell attachment that is specific to osteoblasts on the respective surfaces. For RUNX2, there was a significant 7 fold and 3 fold increase of gene on the 159

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3.6. Mineralization assay Alizarin red was seen stained on the surfaces that had calcium and phosphate deposits. On comparing the surfaces for staining, the surfaces that were etched and coated with fibrin had the most staining which indicated most deposits of calcium and phosphate while minimum stain was observed on the control surfaces as shown in Fig. 11. The stain was also quantified and represented as a graph. 3.7. Bacterial activity of E. faecalis The Alamar blue assay confirmed the proliferation of bacteria when compared to different modified surfaces. There was no significant increase or decrease in bacterial proliferation when compared to all surfaces. This shows that there is no significant difference in bacterial growth on all the surface studied. Fig. 12 shows the Alamar blue proliferation of bacteria on the different surfaces of titanium.

Fig. 7. Estimation of protein adsorption. The graph shows the adsorption of fibrinogen on control surfaces, fibrinogen adsorption on etched surfaces and fibrin adsorption on etched surfaces. There is a significant increase in the adsorption of fibrin on etched surfaces compared to that on the control surfaces (*P value < 0.05; n=3).

4. Discussion Giving an overview of the results, the surface of grade IV titanium was biofunctionalized by increasing the roughness and coating it with fibrinogen or fibrin. This modification has proven to be beneficial for the osteoblastic cells to adhere, proliferate and mineralize more compared to its activity on polished and normal etched surfaces. The surfaces of titanium were successfully modified by increasing the surface roughness by acid etching. Acid etching has proven to create micro textured homogenous roughening on metallic surfaces. Also, since acid generates H+ ions on the surface, this layer is beneficial for attracting negatively charged molecules such as protein (fibrinogen in particular) [51]. This layer of fibrinogen on titanium proves to be effective in osteoblastic proliferation and enhanced osseointegration in the human body as fibrinogen is one of the first proteins that attach on the surface of any metallic implant and helps in aiding osseointegration [31]. Surface characterization was done to check for the surface morphology and characteristics of the bio functionalized samples. To confirm this, the surface roughness was assessed by a white light microscope. The white light microscopic results proved that surfaces etched in HCl/H2SO4 had more roughness compared to surfaces etched in HF/ HNO3. This is due to the exposure of the surfaces for around 60 min in the condition HCl/H2SO4 while just 10 min in the condition HF/HNO3 [14]. Longer the exposure to the acid, better is the surface irregularities.

Fig. 8. Cell proliferation. In the graphical representation of osteoblasts proliferating on respective surfaces on days 2,3,5,10 and 14; we can observe a significant increase in the cell proliferation on etched surfaces coated with fibrinogen. (*P value < 0.05; **P-value < 0.01; n=3).

Fig. 9. a: Cell adhesion on surfaces by SEM. (a and b) i surfaces indicate less amount of cells attached on the control surfaces, (a and b) ii surfaces indicate a higher number of cells attached on the etched surfaces (a and b) iii and iv indicate maximum number of cells attached on the biofunctionalized surfaces with the highest number of cells on surfaces that were etched and coated with fibrinogen on both days 1 and 5. (n=3). b: Cell adhesion on surfaces by confocal imaging. (i) surfaces indicate less amount of cells attached on the control surfaces, (ii) surfaces indicate a higher number of cells attached on the etched surfaces (iii) and (iv) indicate maximum number of cells attached on the biofunctionalized surfaces with the highest number of cells on surfaces that were etched and coated with fibrinogen on both days 1 and 5. (n=3). 160

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Fig. 9. (continued)

implant enters the human body. SEM images proved the presence of fibrinogen and fibrin on the surfaces by the look at the images compared to the uncoated surfaces that were etched. WCA data proved that HF/ HNO3 etched surfaces coated with protein were the most hydrophilic compared to that of the surfaces etched in HCl/ H2SO4. Since these surfaces etched in HF/ HNO3 have both optimum roughness and hydrophilicity, these surfaces were further considered best and suitable for the biological environment. MG63 cells were placed on the biofunctionalized samples and Alamar blue assay indicated that the biofunctionalized surfaces had more cells proliferating on them compared to that of control and etched surfaces. This is due to the presence of fibrinogen on the surfaces that attract more cells onto them. Fibrinogen has the tendency to attract more osteoblasts [31]; similarly, the presence of cells attaching on the

However, when comparing HCl/H2SO4 and HF/ HNO3 groups, the roughness after fibrin and fibrinogen coating did not alter significantly. To determine the stability of the protein layer attached to the surfaces, electrochemical studies were performed. There was an increase in the potential of the curves of fibrinogen and fibrin coated samples which indicated the stability of the protein coating for a period of 39 h while the OCP values of the etched samples was the least indicating its least resistance to corrosion. The EIS data also proved that the resistance had increased, and capacitance had decreased through the 6 EISs in the case of fibrinogen and fibrin coated surfaces indicating a stable protein layer while the resistance had decreased for the surfaces that were just etched indicating the unstable nature of the etched surface. This is an effective way to prove that the protein coating on the surfaces was stable for a period of 39 h after the implantation indicating the exact time accomplished by the protein to remain stable once the

Fig. 10. Expression of genes on days 7 and 14. Significant upregulation of genes specific to osteoblasts was observed on surfaces etched and coated with fibrin on both days 7 and 14 for the genes RUNX2, BGLAP, BMP2, ALP and COL. This shows that osseointegration is acting fast and effective on the biofunctionalized surfaces compared to that of the control and etched surfaces. (*P value < 0.05; **P-value < 0.0; n=3). 161

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Fig. 11. Mineralization assay. Alizarin red staining is observed brighter and more on the biofunctionalized surfaces and was quantified to be maximum on surfaces that were etched and coated with fibrin. There is a significant increase in the staining of surfaces that were biofunctionalized as well as etched surfaces compared to that of the control group. (*P value < 0.05; **P-value < 0.01; n=3).

adhesion of bacteria to the samples coated with a protein that indicated that there would be no extreme cause of bacterial infection produced by the samples due to protein compared to control uncoated surfaces. Fig. 13 depicts the importance of the biofunctionalization and the need to biofunctionalized the titanium samples. As it can be seen from the mechanism, with the increase in surface roughness by acid etching (step 1 and 2), there will be a presence of H+ layer on the surfaces of titanium which allows the adherence of negatively charged fibrinogen on the surfaces easily and more in number compared to the unmodified surfaces allowing lesser adherence of fibrinogen on them (step 3). Due to the presence of more fibrinogen on biofunctionalized surfaces, there is more attraction and adherence of osteoblasts on the surfaces while there is lesser attraction and adherence of osteoblasts on the surfaces that are unmodified (step 4). Finally, due to the increased adherence of protein on biofunctionalized surfaces, there is more production of bone and ECM that provide better and enhanced osseointegration on them. The clinical implication of this project is that, surface modification is an effective way to enhance osseointegration of metals with the human body. Hence, by performing such biofunctionalization by increasing surface roughness and coating with fibrinogen, it is proved that there is an enhancement of osseointegration as there is enhanced expression of genes and also increase in mineralization of metal to the bone. This will in turn reduce complications and other infections produced by our body as an immune response. This study demonstrates a method of increasing osseointegration of an implant surface by using an innovative Surface modification- biofunctionalization approach. The shortening of time for osseointegration is desirable for clinical treatment in oral implantology that can decrease the risks of peri-implant infections and loosening of implants due to early complications related to prosthetic rehabilitation. This approach proposed in the present project is easily reproducible and it involves low costs for titanium functionalization. It is also important for the industrial production of implants as well as for patient satisfaction concerning costs for treatment. However, the limitations associated with this project is that animal studies have not been carried out to check for the in vivo analysis of osseointegration of this biofunctionalized surface. Our future direction could be to coat the surface of titanium with fibrinogen and an antibacterial agent together as an alternative to reduce bacterial growth. Mechanical stability of the protein layer is a concern, it will be

Fig. 12. Alamar blue analysis of bacterial adhesion. This analysis shows the adhesion of bacteria on different surfaces of titanium. (P value of etched samples are 0.30, fibrinogen coated samples are 0.200 and fibrin coated samples are 0.66; n=3).

surfaces was observed by SEM and confocal microscope. More the number of cells were seen attached on the surfaces of etched titanium coated with protein than compared to the control samples. This indicated that the presence of protein does attract more cells. Expression of genes specific to osteoblasts was studied and it was proven to havesignificant upregulation of genes on etched surfaces coated with protein proving it to enhance the process of osseointegration. This again is due to the presence of fibrinogen on the surfaces [31]. Mineralization assay proved that the etched surfaces coated with protein had more calcium and phosphate deposits that aids to the bone formation that will, in turn, lead to the bonding of the implant and the bone quickly, proving it to be a better surface for implant-bone integration. Due to the increase in the mineralization by the cells attached to the surfaces coated with fibrin, hence it can be proven that this surface is best suited for enhancing and quickening the process of osseointegration. By doing so, the complications during early implantation can be tremendously reduced. In order to check if any augmentation of bacterial growth has been occurred due to the protein surface functionalization we have quantified the bacterial count. There was no significant difference in the 162

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Fig. 13. Overall schematic representation of biofunctionalization. The left section of the picture represents surface which doesn’t undergo modification while the right section depicts the surfaces that are biofunctionalized leading to better osseointegration of the implant.

stability of the implants and decrease the risks and complications produced by the body due to the failure of early osseointegration of the metal-bone interface.

addressed in future investigations. In addition, the possible effect of UV light on WCA is neglected in the study. Chemical crosslinkers could be used to bond the protein on the surfaces to enhance its mechanical stability.

Acknowledgements 5. Conclusions Acknowledgements to Blazer Foundation (Rockford, IL, USA), MBT program at Biomed Science, UICOM at Rockford.

From this study, the following conclusions can be drawn:

• Biofunctionalized Ti surfaces were developed by increasing surface roughness and coated it with fibrinogen. • The biofunctionlized surface was characterized through for wettability and surface morphology • Stability of the adhered protein coating was evaluated by electro• • • • • •

References Ibrahim, M.Z., Sarhan, A.A.D., Yusuf, F., Hamdi, M., 2017. Biomedical materials and techniques to improve the tribological, mechanical and biomedical properties of orthopedic implants – a review article. J. Alloys Compd. 714, 636–667. Prasad, K., et al., 2017. Metallic biomaterials: current challenges and opportunities. Materials 10. Gotman, I., 1997. Characteristics of metals used in implants. J. Endourol. 11, 383–389. Souza, J.C.M., et al., 2015. How do titanium and Ti6Al4V corrode in fluoridated medium as found in the oral cavity? An in vitro study. Mater. Sci. Eng. C Mater. Biol. Appl. 47, 384–393. Dental implant. Wikipedia (2018). Oshida, Y., et al., 2010. Dental Implant Systems. Int. J. Mol. Sci. 11, 1580–1678. Gemelli, E., Camargo, N.H.A., 2007. Oxidation kinetics of commercially pure titanium. Matéria 12, 525–531. Albrektsson, T., Brånemark, P.I., Hansson, H.A., Lindström, J., 1981. Osseointegrated titanium implants. Requirements for ensuring a long-lasting, direct bone-to-implant anchorage in man. Acta Orthop. Scand. 52, 155–170. Parithimarkalaignan, S., Padmanabhan, T.V., 2013. Osseointegration: an update. J. Indian Prosthodont. Soc 13, 2–6. Mendonça, G., Mendonça, D.B.S., Aragão, F.J.L., Cooper, L.F., 2008. Advancing dental implant surface technology – from micron- to nanotopography. Biomaterials 29, 3822–3835.

chemical techniques and found it is stable for the initial period of the implantations. Samples etched in HF/HNO3 showed higher hydrophilicity and can easily be used in the biological environment. It was also observed that osteoblasts proliferate and adhere more on the etched surface coated with fibrinogen. Upregulation of genes specific to osteoblastswas observed on surfaces etched and coated with fibrin. Mineralization of osteoblasts were found to be greater on the surfaces etched and coated with fibrin. No potential risk of bacterial infection was identified on the biofunctionalized surfaces. Therefore, these biofunctionalized samples can enhance the primary 163

Journal of the Mechanical Behavior of Biomedical Materials 96 (2019) 152–164

S. Rao, et al.

titanium metal via anodic oxidation treatment. Biomaterials 25, 1003–1010. Holm, B., Nilsen, D.W., Kierulf, P., Godal, H.C., 1985. Purification and characterization of 3 fibrinogens with different molecular weights obtained from normal human plasma. Thromb. Res. 37, 165–176. Kaijzel, E.L., Koolwijk, P., van Erck, M.G.M., van Hinsbergh, V.W.M., de Maat, M.P.M., 2006. Molecular weight fibrinogen variants determine angiogenesis rate in a fibrin matrix in vitro and in vivo. J. Thromb. Haemost. 4, 1975–1981. van der Stok, J., et al., 2015. Full regeneration of segmental bone defects using porous titanium implants loaded with BMP-2 containing fibrin gels. Eur. Cell. Mater. 29, 141–153 discussion 153-154. Gaddam, R., Sefer, B., Pederson, R. & Antti, M.-L. Study of alpha-case depth in Ti-6Al-2Sn4Zr-2Mo and Ti-6Al-4V. IOP Conf. Ser. Mater. Sci. Eng. 48, 012002 (2013). Hung, K.-Y., Lin, Y.-C., Feng, H.-P., 2017. The effects of acid etching on the nanomorphological surface characteristics and activation energy of titanium medical materials. Materials 10. Souza, J.C.M., et al., 2013. Corrosion behaviour of titanium in the presence of Streptococcus mutans. J. Dent. 41, 528–534. Sela, M.N., Badihi, L., Rosen, G., Steinberg, D., Kohavi, D., 2007. Adsorption of human plasma proteins to modified titanium surfaces. Clin. Oral Implants Res. 18, 630–638. Jia, S., Zhang, Y., Ma, T., Chen, H., Lin, Y., 2015. Enhanced hydrophilicity and protein adsorption of titanium surface by sodium bicarbonate solution. J. Nanomater. https://doi.org/10.1155/2015/536801. Bai, Z., Filiaggi, M.J., Dahn, J.R., 2009. Fibrinogen adsorption onto 316L stainless steel, Nitinol and titanium. Surf. Sci. 603, 839–846. Xu, L.-C., Siedlecki, C.A., 2007. Effects of surface wettability and contact time on protein adhesion to biomaterial surfaces. Biomaterials 28, 3273–3283. Divieto, C., Sassi, M.P., 2015. A first approach to evaluate the cell dose in highly porous scaffolds by using a nondestructive metabolic method. Future Sci. OA 1. Bermudez-Reyes, B., et al., 2017. Effect on growth and osteoblast mineralization of hydroxyapatite-zirconia (HA-ZrO2) obtained by a new low temperature system. Biomed. Mater. https://doi.org/10.1088/1748-605X/aaa3a4. (11) A nanoparticulate injectable hydrogel as a tissue engineering scaffold for multiple growth factor delivery for bone regeneration. ResearchGate Available at: 〈https:// www.researchgate.net/publication/234071146_A_nanoparticulate_injectable_ hydrogel_as_a_tissue_engineering_scaffold_for_multiple_growth_factor_delivery_for_ bone_regeneration〉. (Accessed: 16.08.2018). Velázquez-Cayón, R., et al., 2017. Production of bone mineral material and BMP-2 in osteoblasts cultured on double acid-etched titanium. Med. Oral Patol. Oral Cir. Bucal. 22, e651–e659. Yamamichi, N., et al., 2008. Gene expression monitoring in osteoblasts on titanium coated with fibronectin-derived peptide. Dent. Mater. J. 27, 744–750. Pozio, A., Palmieri, A., Girardi, A., Cura, F., Carinci, F., 2012. Titanium nanotubes activate genes related to bone formation in vitro. Dent. Res. J. 9, S164–S168. Yusa, K., Yamamoto, O., Takano, H., Fukuda, M., Iino, M., 2016. Zinc-modified titanium surface enhances osteoblast differentiation of dental pulp stem cells in vitro. Sci. Rep. 6. Subramani, K., Pandruvada, S.N., Puleo, D.A., Hartsfield, J.K., Huja, S.S., 2016. In vitro evaluation of osteoblast responses to carbon nanotube-coated titanium surfaces. Prog. Orthod. 17, 23. Ma, J., Tong, Z., Ling, J., Liu, H., Wei, X., 2015. The effects of sodium hypochlorite and chlorhexidine irrigants on the antibacterial activities of alkaline media against Enterococcus faecalis. Arch. Oral Biol. 60, 1075–1081. Zareidoost, A., Yousefpour, M., Ghaseme, B., Amanzadeh, A., 2012. The relationship of surface roughness and cell response of chemical surface modification of titanium. J. Mater. Sci. Mater. Med. 23, 1479–1488.

Osseointegration: Examining the Pros and Cons. Amputee Coalition. (13) Biological factors responsible for failure of osseointegration in oral implants. ResearchGate Available at: 〈https://www.researchgate.net/publication/283857577_ Biological_factors_responsible_for_failure_of_osseointegration_in_oral_implants〉. (Accessed: 24.07.2018). Mouraret, S., et al., 2013. Development of mice models to study implant osseointegration and failure in alveolar bone. BioScientifica. https://doi.org/10.1530/boneabs.1. PP183. Chrcanovic, B.R., Albrektsson, T., Wennerberg, A., 2014. Reasons for failures of oral implants. J. Oral Rehabil. 41, 443–476. Smeets, R., et al., 2016. Impact of dental implant surface modifications on osseointegration. BioMed. Res. Int. https://doi.org/10.1155/2016/6285620. Bazaka, K., Jacob, M.V., Chrzanowski, W., Ostrikov, K., 2015. Anti-bacterial surfaces: natural agents, mechanisms of action, and plasma surface modification. RSC Adv. 5, 48739–48759. Cappiello, M., et al., 2008. Evaluation of peri-implant bone loss around platform-switched implants. Int. J. Periodontics Restorative Dent. 28, 347–355. Junker, R., Dimakis, A., Thoneick, M., Jansen, J.A., 2009. Effects of implant surface coatings and composition on bone integration: a systematic review. Clin. Oral Implants Res. 20 (Suppl 4), 185–206. Esposito, M., Ardebili, Y., Worthington, H.V., 2014. Interventions for replacing missing teeth: different types of dental implants. Cochrane Database Syst. Rev. CD003815. https://doi.org/10.1002/14651858.CD003815.pub4. Dohan Ehrenfest, D.M., Coelho, P.G., Kang, B.-S., Sul, Y.-T., Albrektsson, T., 2010. Classification of osseointegrated implant surfaces: materials, chemistry and topography. Trends Biotechnol. 28, 198–206. Hanawa, T., 2010. Biofunctionalization of titanium for dental implant. Jpn. Dent. Sci. Rev. 46, 93–101. Le Guéhennec, L., Soueidan, A., Layrolle, P., Amouriq, Y., 2007. Surface treatments of titanium dental implants for rapid osseointegration. Dent. Mater. Off. Publ. Acad. Dent. Mater. 23, 844–854. Nanomaterials and Nanosystems for Biomedical Applications | M. Reza Mozafari | Springer. Available at: 〈https://www.springer.com/us/book/9781402062889〉. (Accessed: 16.08.2018). Gittens, R.A., Olivares-Navarrete, R., Schwartz, Z., Boyan, B.D., 2014. Implant osseointegration and the role of microroughness and nanostructures: lessons for spine implants. Acta Biomater. 10, 3363–3371. Lutz, R., et al., 2010. Biofunctionalization of titanium implants with a biomimetic active peptide (P-15) promotes early osseointegration. Clin. Oral Implants Res. 21, 726–734. Hemmersam, A.G., Foss, M., Chevallier, J., Besenbacher, F., 2005. Adsorption of fibrinogen on tantalum oxide, titanium oxide and gold studied by the QCM-D technique. Colloids Surf. B Biointerfaces 43, 208–215. Marucco, A., Fenoglio, I., Turci, F., Fubini, B., 2013. Interaction of fibrinogen and albumin with titanium dioxide nanoparticles of different crystalline phases. J. Phys. Conf. Ser. 429 012014. Horasawa, N., Yamashita, T., Uehara, S., Udagawa, N., 2015. High-performance scaffolds on titanium surfaces: osteoblast differentiation and mineralization promoted by a globular fibrinogen layer through cell-autonomous BMP signaling. Mater. Sci. Eng. C Mater. Biol. Appl. 46, 86–96. Lollobrigida, M., et al., 2018. Biomimetic implant surface functionalization with liquid LPRF products: in vitro study. BioMed. Res. Int. 2018. Orsini, G., Assenza, B., Scarano, A., Piattelli, M., Piattelli, A., 2000. Surface analysis of machined versus sandblasted and acid-etched titanium implants. Int. J. Oral Maxillofac. Implants 15, 779–784. Yang, B., Uchida, M., Kim, H.-M., Zhang, X., Kokubo, T., 2004. Preparation of bioactive

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