Novel doped hydroxyapatite thin films obtained by pulsed laser deposition

Applied Surface Science 265 (2013) 41–49

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Novel doped hydroxyapatite thin films obtained by pulsed laser deposition L. Duta a , F.N. Oktar b,c,d , G.E. Stan e , G. Popescu-Pelin a , N. Serban a , C. Luculescu a , I.N. Mihailescu a,∗ a

National Institute for Lasers, Plasma and Radiation Physics, Lasers Department, 409 Atomistilor Street, Magurele, Romania Department of Bioengineering, Faculty of Engineering, Marmara University, Goztepe, Istanbul 34722, Turkey c Department of Medical Imaging Technics, Vocational School of Health Services, Marmara University, Uskudar, Istanbul 34668, Turkey d Nanotechnology and Biomaterials Application & Research Centre, Marmara University, Istanbul, Turkey e National Institute of Materials Physics, 105 Bis Atomistilor Street, Magurele, Romania b

a r t i c l e

i n f o

Article history: Received 4 July 2012 Received in revised form 5 September 2012 Accepted 12 October 2012 Available online 12 November 2012 Keywords: Doped hydroxyapatite Ovine/bovine bones Highly adherent thin films Pulsed laser deposition

a b s t r a c t We report on the synthesis of novel ovine and bovine derived hydroxyapatite thin films on titanium substrates by pulsed laser deposition for a new generation of implants. The calcination treatment applied to produce the hydroxyapatite powders from ovine/bovine bones was intended to induce crystallization and to prohibit the transmission of diseases. The deposited films were characterized by scanning electron microscopy, X-ray diffraction, Fourier transform infrared spectroscopy, and energy dispersive X-ray spectroscopy. Pull-off adherence and profilometry measurements were also carried out. X-ray diffraction ascertained the polycrystalline hydroxyapatite nature of the powders and films. Fourier transform infrared spectroscopy evidenced the vibrational bands characteristic to a hydroxyapatite material slightly carbonated. The micrographs of the films showed a uniform distribution of spheroidal particulates with a mean diameter of ∼2 ␮m. Pull-off measurements demonstrated excellent bonding strength values between the hydroxyapatite films and the titanium substrates. Because of their physical–chemical properties and low cost fabrication from renewable resources, we think that these new coating materials could be considered as a prospective competitor to synthetic hydroxyapatite used for implantology applications. © 2012 Elsevier B.V. All rights reserved.

1. Introduction Calcium phosphate (CaP) coatings for metallic implants are the subject of the recent researches in the attempt to improve their biointegration. The mineral constituent of vertebrate skeletal systems mainly consists of calcium-deficient hydroxyapatite (HA) doped with various ions [1]. Pure HA, with chemical formula Ca10 (PO4 )6 (OH)2 , is one of the most well known and studied biocompatible materials, used as implantable ceramic due to its chemical and structural similarity with human hard tissues (50% mass, 70% volume) [2]. However, HA presents some disadvantages. For instance, HA ceramics are very brittle and break easily in bulk [3], exhibiting poor mechanical properties, especially in wet environments. Therefore, HA cannot be used in orthopedic devices that must withstand substantial loadings during their expected lifetimes. To overcome this drawback, HA can be applied in the form of a very thin coating on metallic implants or areas where no loads are exerted. The aim is to significantly improve the metallic implant’s osteoconductivity and to combine the mechanical advantages of

∗ Corresponding author. Tel.: +40 214574491; fax: +40 214574243. E-mail address: ion.mihailescu@inflpr.ro (I.N. Mihailescu). 0169-4332/$ – see front matter © 2012 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.apsusc.2012.10.077

metals with the excellent bioactivity of the ceramic [3,4]. In this respect, HA can be reinforced with various secondary phases of oxides and glasses, such as SiO2 , MgO, Al2 O3 and ZrO2 [5], commercial inert glass (CIG) [6], Li2 O [7], Si [8], or others [1], to constitute a strong composite appropriate for load-involving applications. Bioceramic thin films are nowadays used for medical prostheses to modify the implant surface [9]. HA coatings on medical implants deposited by conventional thermal plasma spraying function as an intermediate layer between human tissues and the metallic implant [10]. However, this method supplies very thick films which might crack, peel or dissolve in biological fluids. These shortcomings supported the use of alternative coating techniques for the deposition of films, such as liquid plasma spraying [11], radio-frequency and direct current magnetron sputtering [11,12], ion implantation [3,11], ion beam sputtering [11], pulsed laser deposition (PLD) [3,11,13] or combination of different techniques [11]. In the field of thin film growth, PLD has proven to be among the most versatile methods [14], with characteristics superior to conventional deposition techniques (fast processing, reliability and low production cost). In this technique, a pulsed laser beam is focused onto a target in order to evaporate its surface under vacuum or different gas atmospheres. The vaporized material consisting of

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ions, atoms, molecules or clusters is subsequently deposited onto a neighbor, generally parallel substrate. PLD technique ensures the stoichiometric transfer of the material from target to collector, an excellent adherence of the deposited structures to substrate and a high rate film growth [14]. Post-deposition thermal treatments (generally in water vapors) are of key importance. Their role is to restore the stoichiometry of the synthesized compound and to improve the overall crystalline status of the coating. Ovine derived HA (SHA) and bovine derived HA (BHA) are natural materials which can be produced by calcination methods [15,16]. Ozyegin et al. reported that sintered BHA is safe from potentially transmitted diseases. Besides, the manufacturing of these powders is very simple and economic [16]. Recently, prions have been discovered as the smallest proteinaceous infectious particles, which resist inactivation by procedures that modify nucleic acids. The finding that these proteins alone can transmit an infectious disease and cause harm to the host tissue has come as a considerable surprise to the scientific community [17]. The temperature applied for calcination by Goller et al. [15] and Ozyegin et al. [16] was around 850 ◦ C for 4 h and it was demonstrated that no disease transmission agents (including prions) can survive to such high temperatures. All materials used in our experiments were prepared with respect of these regulations. This study aimed to compare the physical–chemical properties of different natural (ovine/bovine bones)-derived HA composites and commercially bioactive HA for obtaining a new generation of metal implant coatings. BHA was doped with Li2 O (1%) to improve the biocompatibility/bioactivity action or with CIG (10%) to reinforce mechanical properties [6,18]. 2. Materials and methods 2.1. Powders preparation The ovine and bovine femoral bones were purchased on market. Before usage, they were submitted to a veterinary control. The head of the bones were cut off and only the femoral shafts were further used. Bone marrows were extracted and all other soft tissue residues were removed from the shafts. The femoral shafts were cut into small pieces and washed with distilled water very neatly. Cleaned parts were deproteinized for 14 days in an alkali solution of 1% sodium hypochlorite. After washing and drying, the bone pieces were calcinated at 850 ◦ C for 4 h in air, in order to completely eliminate the organic components of bone, as proved by the absence in FTIR spectra of amide, C C and C H bonds’ vibration modes (belonging to collagen). The calcination treatment applied to produce the HA powders from ovine and bovine bones was intended also to induce crystallization and to totally prohibit risks of diseases transmission. The resulting calcinated bone specimens were first crushed with a mortar and pestle and then ball-milled to fine powders (i.e., with particles of submicron mean size). The fine powder of calcinated BHA was admixed with Li2 O (1%) (BHA:Li) or fine CIG (10%) (BHA:CIG) powders. The glass is produced by Pas¸abahc¸e A.S¸., Istanbul, and has the composition: 73.65 wt.% SiO2 ; 11.67 wt.% Na2 O; 5.89 wt.% CaO; 2.60 wt.% Al2 O3 ; 1.91 wt.% MgO [19]. Commercial HA powder, further noted as HAsyn , from Calbiochem, Merck, Germany, with a 10 ␮m mean size was used for comparison. 2.2. PLD experiment SHA, BHA:Li, BHA:CIG and HAsyn powders were pressed at 5 MPa in a 2 cm diameter mold. The resulting pellets were heat-treated for

4 h at 1000◦ C in air, with a heating rate of 20 ◦ C/min and a cooling ramp of 5 ◦ C/min, and used as targets in PLD experiments. The ablated material was collected either onto titanium (Ti) substrates (Dentaurum GmbH, Politehnico di Milano), of 1.2 cm diameter and 0.15 cm thickness, or on flat Si(1 0 0) wafers (1 cm × 1 cm), that were placed parallel to the targets, at 5 cm separation distance. The experimental conditions were identical for both deposition substrates. Prior to introduction into the deposition chamber, the substrates were successively cleaned using a laboratory protocol described elsewhere [20]. The PLD experiments were conducted in a stainless steel deposition chamber, fitted with a rotary pump (model Alcatel SD2033). The ablation of the targets was produced with a KrF* excimer laser source, model COMPexPro 205 from Lambda Physik/Coherent Radiation ( = 248 nm,  FWHM ≤ 25 ns), running at a repetition rate of 10 Hz. The laser beam was incident at 45◦ on target surface. The laser fluence on target was set at ∼3.3 J/cm2 (the corresponding pulse energy was of 330 mJ). Before deposition, the targets were submitted to a “cleaning” process with 1000 laser pulses. During this procedure, a shutter was interposed between the target and the substrate to collect the flux of expulsed impurities. By the application of the multi pulse laser irradiation, the targets were continuously rotated with 0.3 Hz and rasted along two orthogonal axes in order to avoid target piercing and improve the overall quality of the deposited films. During the experiments, the substrates were heated and maintained at a constant temperature of 500 ◦ C using a PID-EXCEL temperature controller. A heating rate of 25 ◦ C/min and a cooling ramp of 10 ◦ C/min have been applied. For the growth of one thin film, 15.000 consecutive laser pulses were applied. All depositions were performed in 50 Pa water vapors. In order to restore the stoichiometry of the synthesized structures and to improve the crystalline status of the coatings, all samples were submitted to a 6 h post-deposition thermal treatment (in water vapors), at the same temperature used during the deposition process (500 ◦ C). 2.3. Characterization of deposited structures (a) The surface morphology of the deposited films was investigated by scanning electron microscopy (SEM) with a FEI Inspect S electron microscope. The measurements were performed at 20 kV acceleration voltages, in high vacuum, under secondary electrons acquisition mode. The samples were coated with a thin Au film in order to prevent electrical charging. Cross-section SEM images were recorded on samples deposited on Si wafers in order to estimate the HA films thickness. Surface roughness studies were performed with a Dektak6M Stylus profiler, 2.5 ␮m indenter tip, weight 5 mg, acquisition time 30 s, measurement length 2.000 ␮m. (b) Identification of crystalline phases in the PLD thin films was performed by X-ray diffraction (XRD) using a Bruker D8 Advance ˚ radiation. The scattered diffractometer, with Cu K␣ ( = 1.5418 A) intensity was scanned in the 2 range 20–60◦ , with a step size of 0.04◦ , and 20 s/step. For the XRD patterns, Rietveld quantitative analysis has been carried out using a MAUD v.2.31 software [21]. (c) Fourier transform infrared (FTIR) spectroscopy was performed with a Perkin Elmer BX Spectrum spectrometer, in attenuated total reflection mode using a Pike-MIRacle diamond head of 0.18 cm diameter. The spectra were recorded in the range 4000–400 cm–1 , with a resolution of 4 cm−1 and a total of 50 scans/experiment. (d) Energy dispersive spectroscopy (EDS), with a SiLi detector (model EDAX Inc.) operated at 20 kV, was conducted in duplicate, on different relatively large regions of 250 ␮m × 250 ␮m. (e) Films adherence to the Ti substrates was tested by pull-out measurements. The bonding strength (tensile) at the biofunctional

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Fig. 1. SEM micrographs recorded in top-view (a, d, g, and j) and cross-view modes (b, e, h, and k) for SHA (a and b), BHA:Li (d and e), BHA:CIG (g and h) and HAsyn (j and k). EDS spectra for HA films (c, f, i, and l).

film–substrate interface is considered a critical factor in fabricating high-quality implants and long-term stability of these devices in situ [4,22,23]. The investigations were carried out using a DFD Instruments PAT MICRO adhesion tester AT101 equipped with a 0.28 cm diameter stainless steel test elements, glued to films surface with a cyano-acrylate one-component Epoxy adhesive, E1100S. After gluing, the samples were placed in an oven for thermal curing 1 h at 130 ◦ C. Each test element was pulled-out vertically with a calibrated hydraulic pump until detachment. The experimental procedure was conducted in accordance with the ASTM D4541 and ISO 4624 standards. For each type of HA film, 5 assays were run. The statistical significance was determined using an unpaired Student’s t-test. The differences were considered relevant when p < 0.05. In case of films, the bonding strength values should be considered as relative values, as the pulling force was divided to

Table 1 Particulates density on HA films. Sample type SHA BHA:Li BHA:CIG HAsyn

Density (particulates × cm−2 ) (5.5 (11.2 (13.3 (1.4

± ± ± ±

0.8) × 10−7 0.7) × 10−7 1.1) × 10−7 0.5) × 10−7

the test dollies active surface, and not to the actually film delaminated area. Both adhesive fracture (at the interface), in the center of the tested area, and cohesive fracture (near the interface or in the film volume), at the tested area edges, were involved in the film mechanical failure.

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3. Results 3.1. SEM and profilometry SEM examination of the HA thin layers was carried out under three magnifications (1000, 2500, and 10,000×). Typical SEM micrographs of HA films show a rather irregular morphology. Particulates (droplets) were found on the surface with sizes in the range of 0.2–5.9 ␮m. Fig. 1 displays typical SEM images for all types of PLD films, consisting of merged spheroid grains with different densities (Table 1). Fig. 2 shows the histograms with size distribution of particulates. One observes that the size distribution of particulates has an extended maximum plateau between 0.3 and 0.7 ␮m for SHA, 0.5 and 0.7 ␮m for BHA:Li, 0.4 and 0.8 ␮m for BHA:CIG, and 0.5 and 0.8 ␮m for HAsyn , respectively. The decrease to large size values is gradual in case of SHA and HAsyn , and rather steep for BHA:Li and BHA:CIG films. It is interesting to correlate the values of particulates density in Table 1 and the surface roughness measured by profilometry. Thus, the values of the roughness obtained as mean ± standard deviation (SD), were 0.61 ± 0.05 ␮m for SHA, 0.71 ± 0.06 ␮m for BHA:Li, 0.77 ± 0.03 ␮m for BHA:CIG, and 0.76 ± 0.1 ␮m for HAsyn . The roughness of the bare Ti of 0.47 ± 0.02 ␮m is about half of the roughness of coated samples which is obviously related to the presence of particulates in the second case. Also, the surface with the largest density of particulates, BHA:CIG, exhibits the largest roughness. This trend is generally preserved with the other samples as well. The cross-section SEM images (Fig. 1b, e, h, and k) revealed HA films with a rather dense microstructure having the following average thicknesses: 2.95 ± 0.21 ␮m for SHA, 3.52 ± 0.16 ␮m for BHA:Li, 2.79 ± 0.02 ␮m for BHA:CIG, and 3.71 ± 0.15 ␮m for HAsyn . The corresponding deposition rates were of 0.2 ± 0.02 nm/pulse for SHA, 0.24 ± 0.02 nm/pulse for BHA:Li, 0.18 ± 0.01 nm/pulse for BHA:CIG, and 0.25 ± 0.01 nm/pulse for HAsyn , respectively. This rates are in relation with the ablation efficiency which is the smallest in case of BHA:CIG because of intense light scattering by glass impurities. We observed a columnar-type structure for SHA and HAsyn samples (Fig. 1b and k), while for BHA:Li and BHA:CIG a more compact structure was visible (Fig. 1e and h). 3.2. XRD The XRD patterns of the PLD targets are presented comparatively in Fig. 3. In case of targets, the XRD analysis revealed that the main crystalline phase is HA [ICDD: 9-432], along with very low concentrations (<1%) of hydrogen phosphorus fluoride hydrate [HPF6 (H2 O)6 , ICDD: 77-0136] and brushite [CaPO3 (OH)·2H2 O, ICDD: 11-0293], as residual phases (inset to Fig. 3). We note that low concentration of fluorine for both targets and films was visible in XRD spectra. The XRD patterns of the PLD films are presented comparatively in Fig. 4. For an eloquent comparison with the targets’ XRD analyses and in order to evidence the possible texturing of the films, symmetric geometry (–) measurements were also performed. The most intensive peaks of the patterns in Fig. 4 are the Ti peaks of the substrate (ICDD: 44-1294). The intensity scales of the graphical representations were therefore chosen so that to emphasize the lower intensity lines originating from deposited films. The films remarkably consisted of a HA phase having different degrees of crystallinity. This was evident from the diffracted intensity variation and the mean crystallite sizes estimated from the FWHM of (0 0 2) and (3 0 0) XRD lines by applying the Scherrer equation. The lines width was corrected for instrumental broadening using a corundum standard reference (NIST SRM 1976). The crystallite growth of films was anisotropic in shape (Table 2), with the

Fig. 2. Histograms of size distribution of particulates on the surface of HA films.

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Table 2 Lattice parameters for targets and films, extracted by Rietveld confinement. Crystallites size were determined at the (0 0 2) and (3 0 0) crystal planes by applying the Scherrer equation. Sample type

Space group

Lattice parameter, a (Å)

Lattice parameter, c (Å)

c/a ratio

D0 0 2 (nm)

D3 0 0 (nm)

D0 0 2 /D3 0 0

SHA target SHA film BHA:Li target BHA:Li film BHA:CIG target BHA:CIG film HAsyn target HAsyn film

Hexagonal, P63 /m (176)

9.4186 ± 9.4080 ± 9.4342 ± 9.4017 ± 9.4412 ± 9.3946 ± 9.418a 9.4025 ±

6.8831 ± 6.8814 ± 6.8832 ± 6.8924 ± 6.8880 ± 6.8864 ± 6.884a 6.8871 ±

∼0.73080 ∼0.73144 ∼0.72960 ∼0.73310 ∼0.72957 ∼0.73301 0.73094 ∼0.73247

130.8 71.5 115.9 83.5 117.7 48.9 177.4 59.2

131.7 53.3 46.2 56.0 46.0 23.4 146.8 47.4

∼0.99 ∼1.34 ∼2.51 ∼1.49 ∼2.56 ∼2.1 ∼1.21 ∼1.25

a

0.0002 0.0020 0.0006 0.0029 0.0003 0.0177 0.0018

0.0002 0.0022 0.0004 0.0019 0.004 0.0190 0.0025

Theoretical values (ICDD: 9-432).

c-axis direction 1.25–2.1 times longer than the a-axis (D0 0 2 /D3 0 0 ). One can see that a more pronounced crystallinity corresponds to BHA:Li films (Fig. 4 and Table 2). In case of BHA:CIG, one can notice an intense amorphous halo, centered at 2 ≈ 30◦ , suggesting the incorporation of an important amount of glassy phase. If for HAsyn films a smaller crystallites’ shape anisotropy was determined (D0 0 2 /D3 0 0 = 1.25), in case of BHA:CIG films the crystallites are the smallest, D0 0 2 = 48.9 nm and D3 0 0 = 23.4 nm, and posses the highest degree of shape anisotropy (D0 0 2 /D3 0 0 = 2.1). There were also identified several weak lines that might be associated with small percents of oxygen deficient titanium oxides, whose formation is due to the substrate oxidation during the postdeposition heat-treatment [24]. Rietveld quantitative analysis revealed that the HA cell parameters present slight variations (Table 2), an indicative of structural modifications associated to the incorporation into the crystalline network of various species (as e.g. Li, Na, Mg, Si, F and Cl). 3.3. FTIR Fig. 3. XRD patterns of HA targets.

Fig. 4. XRD patterns of HA films synthesized by PLD onto Ti substrates (䊉, HA; , TiO and Ti2 O sub-oxides).

The recorded FTIR spectra are given in Fig. 5a–c, and the vibration bands assignation are collected in Table 3. The FTIR spectra of targets (Fig. 5a) displayed the vibration modes of hydroxyl group in the adsorbed water: a broad band in range of 3700–2500 cm−1 generated by 1 and 3 stretching modes of hydrogen-bonded H2 O molecules, and by the O–H 2 bending mode centered at 1647 cm−1 . The presence of hydroxyl group in the structure of HA target materials is confirmed by the presence of the librational mode at 633 cm−1 and of the stretching and translational modes of the (OH)− ions at 3573 cm−1 . These modes were present also for the HA films (at ∼630 and 3597 cm−1 ) (Fig. 5b, inset). The animal origin target materials only present two well-defined bands positioned at higher wavelengths (3448 and 3501 cm−1 ), which can be ascribed to the hydrogenic bond with fluoride ions (OH F). Fluorine is present as an ubiquitous element in bone (∼1–3 wt.%). The presence of OH F stretching modes was difficult to detect for the films (Fig. 5b, inset). Additional IR bands have been observed at 3627 cm−1 (HAsyn , SHA and BHA-based films), 3649 cm−1 (SHA and BHA-based targets), 3698 cm−1 (SHA target), 3731 cm−1 (HAsyn , SHA and BHA-based films) and 3740 cm−1 (BHA-based targets), assignable to the O H stretching vibrations of surface P OH groups (Fig. 5a and b-inset). The target material is not only strongly hydroxylated, but also carbonated, as confirmed by sharp C–O asymmetric stretching (3 ) lines at (∼1419, 1457, and 1544 cm−1 ) (Fig. 5a), characteristic to the biological bone mineral phase [25]. Besides, the vibrational band peaking at ∼877 cm−1 could be the result of the overlapping of the signature band of (HPO4 )2− with the (CO3 )2− (2 ) bending modes. The position of the (CO3 )2− stretching bands denote a mix “B-type” (the replacement of (PO4 )3− ions in the HA lattice) and “A-type” (the replacement of (OH)− ions in the HA lattice) carbonation of the bone HA target material [26]. The presence of the carbonate bands was also visible in case of films (Fig. 5b).

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Fig. 5. FTIR spectra of PLD targets (a) and films (b, overall spectra and (3400–3800 cm−1 ) range zoom – inset; c, (PO4 )3− stretching bands range detail).

Table 3 Assignment of FTIR vibration bands for the PLD targets and films. Vibration band

Sample type SHA Target

(PO4 )3− (4 ) asymm. bending (PO4 )3− (4 ) asymm. bending (OH)− libration PØP Q3 , Q2 and Q1 units symm. stretching (HPO4 )2− ions vibrations and (CO3 )2− (2 ) asymm. bending (PO4 )3− (1 ) symm. stretching

BHA:Li Film

Target

BHA:CIG Film

Target

HAsyn Film

Target

Film

572 607 633 805 877 961

565 600 629 799 872 961

572 607 633 805 877 961

559 599 626 800 873 961

572 607 633 805 877 961

564 590 631 – 862 934

572 607 633 – 875 961

556 598 628 – 877 957

(PO4 )3− (3 ) asymm. stretching

1022 1052 1087

1002 1027 1053

1022 1052 1087

993 1018 1061

1022 1052 1087

995 1023 –

1022 1052 1087

975 1014 1060

(HPO4 )2− ions vibrations (PO2 )− (Q2 ) species asymm. stretching

1159 1270

1089 –

1159 1270

1089 –

1159 1270

– –

1159 –

1089 –

(CO3 )2− (3 ) asymm. stretching

1419 1457 1544

1417 1470 –

1419 1457 1544

1417 1470 –

1419 1457 1544

– – –

1419 1457 –

1417 1470 –

H O H (2 ) bending Hydrogen-bonded H2 O molecules (1 and 3 ) stretching OH F stretching

1647 2500–3800

1647 2500–3800

– –

1647 2500–3800

−

(OH) stretching Surface P OH bands

– –

1647 2500–3800

– –

– –

3448 3501 3573

– – 3597

3448 3501 3573

– – 3597

3448 3501 3573

– – 3597

– – 3573

– – 3597

3649 3698

3627 3731

3649 3740

3627 3731

3649 3740

3627 3731

– –

3627 3731

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Table 4 Chemical compositions in at.% for the PLD targets and for the HA films deposited onto titanium substrates. Sample type

Element Ca

P

Na

Mg

Si

Composition [mean ± SD (at.%)] SHA target SHA film BHA:Li target BHA:Li film BHA:CIG target BHA:CIG film

57.1 54.3 56.7 54.5 54.6 50.7

± ± ± ± ± ±

0.32 0.19 0.10 0.20 0.04 0.36

37.8 39.5 38.1 39.4 35.2 35.9

± ± ± ± ± ±

0.25 0.29 0.30 0.34 0.07 0.26

The strong vibration modes of the phosphate functional groups were evidenced for both HA targets and films (Table 3 and Fig. 5a–c). In a HA crystalline structure, the symmetry of the phosphate tetrahedron lowers and there could exist up to nine IR active vibration modes [27]. Absorption bands at lower wave numbers are ascribed to the bending mode of PØP Q3 , Q2 , Q1 and Q0 units (600–550 cm−1 range) and to symmetric stretching bands of PØP Q3 , Q2 and Q1 units (800–650 cm−1 range), respectively (Table 3). The sharp bands peaking at 1089–1053, 1052–1014, 1022–975 and 961–934 cm−1 ranges belong to 3 asymmetric and 1 symmetric stretching modes of (PO4 )3− groups [4] (Table 3). In case of biological materials (SHA and BHA-based targets) only, an additional band, centered at ∼1270 cm−1 , due to (PO2 )− (Q2 species) asymmetric stretching vibrations, was noticed [28]. The splitting of the 3 asymmetric stretching IR absorption band of phosphate is a fingerprint of a wellcrystallized HA structure, in agreement with the XRD observations. The broader aspect of the IR absorption spectrum of BHA:CIG is assigned to the juxtaposition of the HA phosphate bands with the strong vibration modes of the SiO4 tetrahedra of the glass phase present in a significant amount for these type of PLD structures (Fig. 5c). This behavior is in good agreement with the XRD results, and one can deduce that the glass phase acts as a crystallization inhibitor for the HA phase. 3.4. EDS EDS quantitative analyses were performed for the PLD films in order to estimate the Ca/P ratio. The EDS mapping of the coatings (data not shown here) evidenced a homogeneous distribution of the Ca and P elements, proving the absence of the segregation phenomena, which are usually responsible for the formation of the undesired secondary phases. The presence in low content of other chemical elements, typically found in a healthy bone chemical composition, such as Na, Mg, F, C, was revealed (Fig. 1c, f, i and l). Similar compositions were evidenced for the PLD targets used in experiments. Moreover, for BHA:CIG samples, the presence of Si, Na and Cl, pertaining to the glassy phase doping, was emphasized. No other impurities were detected. The results of the quantitative compositional analysis are presented in Table 4. In case of light elements (Li, C, O, F, and Cl), because of their lower determination accuracy, no data are provided. The sub-stoichiometric Ca/P ratios obtained for both target and films (Table 5) indicate the formation of a calcium-deficient HA Table 5 Ca/P atomic ratio for the PLD targets and films, as determined by EDS analysis. Ca/P molar ratios

Target Film a

SHA

BHA:Li

BHA:CIG

HAsyn

1.51 ± 0.02 1.37 ± 0.02

1.49 ± 0.01 1.38 ± 0.02

1.55 ± 0.01 1.41 ± 0.01

1.67a 1.57 ± 0.01

Theoretical value.

3.2 3.7 2.9 3.7 3.7 5.1

± ± ± ± ± ±

0.10 0.05 0.14 0.08 0.10 0.35

1.9 2.5 2.3 2.4 2.3 2.8

± ± ± ± ± ±

0.09 0.15 0.24 0.07 0.26 0.16

– – – – 4.2 ± 0.39 5.5 ± 0.37

Table 6 Pull-out bonding strength values for the HA films prepared by PLD. Sample type

Pull-out adherence value (MPa)

SHA BHA:Li BHA:CIG HAsyn

66.5 74.8 49.3 42.6

± ± ± ±

8.3 3.8 10.0 7.1

compound or important substitution of the Ca ions by other alkali or rare-earth elements, during powder synthesis. This process seems to be enhanced during the PLD transfer.The compositional data (Tables 4 and 5) have for both targets and films SD < 0.5 at.%, which is within the acceptance limit. 3.5. Pull-out adherence measurements Control tests regarding the quality of the bonding adhesive have been carried out under the same conditions on bare Ti substrates. The adherence values estimated at the stainless steel dolly/Ti substrate interface were of ∼85 MPa. This value is in good agreement with the specification of DFD Instruments® manufacturer [29]. The pull-out adherence values are collected in Table 6 and presented as mean ± SD. The two tailed t testing, assuming unequal variances, showed statistically significant differences between the recorded adherence values (p < 0.05) of samples. The exception was met in case of HAsyn comparation with BHA:CIG samples, when no statistically significant differences have been found (p = 0.28). The adherence values recorded for the HA films are generally similar to the ones often reported in literature for this type of PLD films [30]. Significantly higher pull-out adherence values have been obtained in case of SHA and BHA:Li. The decreased value of adherence recorded for the BHA:CIG structure could be attributed to the intrinsic friability of the glassy doping phase under mechanical stress. We note that solely pull-out adherence values of the order of 40 MPa are considered acceptable for these types of implant coatings [23,31,32]. Therefore, the excellent value of adherence obtained for BHA:Li films should be emphasized. The increased values of adherence obtained for BHA films could be explained in terms of intentional (Li) and unintentional (Na, Mg, and F) doping of these structures. Cai et al. [33] have observed similar effects with Mg and F doping. Moreover, Galca et al. [34] have noticed that when substituting larger radius atoms with Li, in case of ZnO, the crystalline network is less perturbed, therefore limiting the internal stress caused by neighboring crystallites or substrate nature. 4. Discussion According to the biomimetic approach, a material meant to repair the skeletal system must be similar to the biological one in respect with morphology, structure, composition and functionality. As known [1], the actual inorganic phase present in human bone tissues is carbonated, partially crystallized and non-stoichiometric,

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due to the presence of a number of foreign ions. Presently, it exists therefore a large interest for synthesis and deposition of apatites enriched with biologically active ions and molecules, as well as in more resorbable (soluble) CaPs. Cotell [35] demonstrated that the amorphous-to-crystalline HA conversion depends both on temperature and water vapor pressure, which has the role to fix P in HA deposited structures. The FTIR spectra and XRD patterns emphasized the presence in all deposited films of a crystallized HA structure. In our case, the highest degree of crystallization was obtained in case of BHA:Li films. The phosphate bands position in FTIR was characteristic in all cases to PLD HA films [36]. The shift of these bands is generally associated with significant structural changes induced by different levels of doping (substitutional or interstitial positions of the doping ions in the HA structure), and the different crystalline/amorphous phase ratio. The existence of ions not originating from either HA or Ti was confirmed by EDS studies of coatings of SHA or BHA (doped with either Li2 O or CIG), where signals of Na, Mg, Cl, F, O or Si were recorded, besides the majority Ca, P and Ti ions. The presence of the ions in the HA structure makes the coating closer to the actual formula of bone and is known to have a beneficial influence over the HA bioreactivity [1]. Very important, all films contained C. HA containing C is known to be more osseoconductive and more resorbable than stoichiometric HA, allowing the osteoclasts to dissolve the ceramic rapidly, determining a faster osteosynthesis and bone remodeling [4,37]. The origin of C (in HAsyn films) can be partly due to atmospheric pollution during samples storage after deposition. For animal origin materials, C is a ubiquitous element, as the mineral component of bone is in fact a non-stoichiometric carbonated HA. As a consequence, a part of C content was transferred in the films. Changes in crystallinity and/or phosphate groups symmetry distortion by the insertion of various foreign impurities into the HA films structure was also hinted by the XRD results. A low concentration of brushite, as residual phase, was detected in biological targets. This is not unexpected as more than three decades ago Roufosse et al. [38] observed its presence in the embryonic chick bone by X-ray and electron diffraction. Brushite was found to be the major CaP crystalline solid phase in the lowest density bone fractions containing the least mineralized osseous tissues. Their studies showed that brushite is a component of the earliest deposited mineral phase, being a precursor of HA, the CaP phase characteristic to the highly mineralized bone tissue. From the chemical point of view, HA is characterized by the Ca/P atomic ratio, which has a value of 1.67 in case of the stoichiometric compound. On the other hand, non-stoichiometric HA, presenting higher solubility rates is more similar to biological apatite and, as a consequence, is more appropriate for coating applications [1]. We therefore consider that the biological materials we worked with, SHA and BHA (simple or doped), which have a Ca/P atomic ratio of ∼1.5 in targets and only ∼1.4 in films, should ensure high osseoconductivity values and fast osseointegration rates of implants covered with these coatings. SEM examination revealed the significant presence of particulates on the deposited film surface, with sizes in micronic range. Their presence is a common characteristic of the PLD films. It was demonstrated that, in case of biomedical applications, the presence of these particulates (causing a significant increase of surface roughness) proves benefic because they confer a good adhesion and a high proliferation rate to the osteoblast cells cultivated on the surface of the coatings [39]. Moreover, the density of particulates was larger in case of HA films of biologic nature with important expected beneficial effects. At elevated temperatures, the high thermal diffusion coefficients of Ca and P in Ti (±10−8 cm2 s−1 ) improve the interface adherence of synthesized thin films [40]. The highest

adherence values were recorded by pull-out tests in case of SHA and BHA doped films. These films are also expected to get a larger hardness as an effect of the reinforcement with Li2 O and CIG. 5. Conclusions We conducted a comparative study of animal (ovine or bovine) origin and commercial hydroxyapatite (HA) thin films obtained by pulsed laser deposition technique. The bovine HA thin films were doped with Li2 O (1%) or commercial inert glass (10%). All powders of animal origin were calcinated according to a protocol which was demonstrated to guarantee the full security against disease transmission or other contaminations. The presence of HA was clearly detected in all deposited films. The films of animal origin were carbonated and contained Na, Mg, Cl, C, and Si besides the majority Ca and P ions. They were deficient in Ca which is usually associated to a higher solubility (resorbability). They have been also found to be more adherent to Ti substrate and much rougher (up to two times) than the films deposited under identical conditions from commercial HA. The reinforcement with metals or glass is expected to bring them a larger specific micro hardness. We consider that these improved performances should allow for obtaining implant coatings with shorter osseointegration time and better osseoconductive characteristics. Last but not the least, these films are obtained from quite cheap and renewable resources. Acknowledgements All authors acknowledge with thanks the support of this research by the bilateral Romanian-Turkish collaboration agreement/theme 109T6141GR. FNO acknowledges with thanks to Turkish Republic Government Planning Organization in the framework of the Project 2003K120810 for the equipment used in these studies. The authors thank to Dr. I. Pasuk for the XRD measurements and for useful discussions, to Dr. L.S. Ozyegin for the delivery of BHA powders, to Assoc. Prof. U. Karacayli for the delivery of SHA powders, and to Dr. C. Aktas for cross-section SEM measurements. References [1] I.N. Mihailescu, C. Ristoscu, A. Bigi, I. Mayer, Advanced biomimetic implants based on nanostructured coatings synthesized by pulsed laser technologies, in: A. Miotello, M. Ossi (Eds.), Laser–Surface Interactions for New Materials Production, Tailoring Structure and Properties, Springer, New York, 2010, pp. 235–268. [2] C.V.M. Rodriguesa, P. Serricella, A.B.R. Linhares, R.M. Guerdes, R. Borojevic, M.A. Rossi, M.E.L. Duarte, M. Farina, Characterization of a bovine collagenhydroxyapatite composite scaffold for bone tissue engineering, Biomaterials 24 (2003) 4987–4997. [3] B. León, J. Jansen, Thin Calcium Phosphate Coatings for Medical Implants, Springer Science + Business Media, New York, 2009. [4] L.E. Sima, G.E. Stan, C.O. Morosanu, A. Melinescu, A. Ianculescu, R. Melinte, J. Neamtu, S.M. Petrescu, Differentiation of mesenchymal stem cells onto highly adherent radio frequency-sputtered carbonated hydroxylapatite thin films, Journal of Biomedical Materials Research A 95A (2010) 1203–1214. [5] F.N. Oktar, S. Agathopoulos, L.S. Ozyegin, O. Gunduz, N. Demirkol, Y. Bozkurt, S. Salman, Mechanical properties of bovine hydroxyapatite (BHA) composites doped with SiO2 , MgO, Al2 O3 , and ZrO2 , Journal of Materials Science. Materials in Medicine 18 (2007) 2137–2143. [6] O. Gunduz, Z. Ahmad, N. Ekren, S. Agathopoulos, S. Salman, F.N. Oktar, Reinforcing of biologically derived apatite with commercial inert glass, Journal of Thermoplastic Composite Materials 22 (2009) 407–419. [7] F.N. Oktar, M.R. Demirer, O. Gunduz, Y. Gen, S. Agathopoulos, I. Peker, L.S. Ozyegin, S. Salman, Sintering effect on mechanical properties of composites of bovine hydroxyapatite (BHA) and Li2 O, Key Engineering Materials 309–311 (2006) 49–52. [8] Q. Tang, R. Brooks, N. Rushton, S. Best, Production and characterization of HA and SiHA coatings, Journal of Materials Science. Materials in Medicine 21 (2010) 173–181. [9] I.N. Mihailescu, P. Torricelli, A. Bigi, I. Mayer, M. Iliescu, J. Werckmann, G. Socol, F. Miroiu, F. Cuisinier, R. Elkaim, G. Hildebrand, Calcium phosphate thin films

L. Duta et al. / Applied Surface Science 265 (2013) 41–49

[10]

[11]

[12]

[13]

[14] [15]

[16] [17] [18]

[19]

[20]

[21]

[22]

synthesized by pulsed laser deposition: physico-chemical characterization and in vitro cell response, Applied Surface Science 248 (2005) 344–348. C. Renghini, E. Girardin, A.S. Fomin, A. Manescu, A. Sabbioni, S.M. Barinov, V.S. Komlev, G. Albertini, F. Fiori, Plasma sprayed hydroxyapatite coatings from nanostructured granules, Materials Science and Engineering B 152 (2008) 86–90. R.A. Surmenev, A review of plasma assisted methods for calcium phosphatebased coatings fabrication, Surface and Coatings Technology 206 (2012) 2035–2056. G. Socol, A.M. Macovei, F. Miroiu, N. Stefan, L. Duta, G. Dorcioman, I.N. Mihailescu, S.M. Petrescu, G.E. Stan, D.A. Marcov, A. Chiriac, I. Poeata, Hydroxyapatite thin films synthesized by pulsed laser deposition and magnetron sputtering on PMMA substrates for medical applications, Materials Science and Engineering B 169 (2010) 159–168. L. Duta, G. Socol, F. Sima, I.N. Mihailescu, G.E. Stan, D.A. Marcov, L.E. Sima, S.M. Petrescu, A. Melinescu, A. Ianculescu, A. Chiriac, I. Poeata, Increased bioactivity of cranio-spinal implants functionalized with hydroxyapatite nanostructured coatings: morpho-structural characterization and in-vitro evaluation, IEEE Computer Society 11694245 (2010) 127–131. R. Eason, Pulsed Laser Deposition of Thin Films – Applications-Led Growth of Functional Materials, Wiley-Interscience, USA, 2007. G. Goller, F.N. Oktar, S. Agathopoulos, D.U. Tulyaganov, J.M.F. Ferreira, E.S. Kayali, I. Peker, Effect of sintering temperature on mechanical and microstructural properties of bovine hydroxyapatite (BHA), Journal of Sol–Gel Science and Technology 37 (2006) 111–115. L.S. Ozyegin, F.N. Oktar, G. Goller, S. Kayali, T. Yazici, Plasma-sprayed bovine hydroxyapatite coatings, Materials Letters 58 (2004) 2605–2609. A. Aguzzi, Prion diseases of humans and farm animals: epidemiology, genetics, and pathogenesis, Journal of Neurochemistry 97 (2006) 1726–1739. A.P.M. Shainberg, P. Valerio, A. Zonari, F.N. Oktar, L.S. Ozyegin, M.P.F. Graca, M.F. Leite, A.M. Goes, Attachment and proliferation of osteoblasts on lithiumhydroxyapatite composites, Advanced Materials Science and Engineering (2012), http://dx.doi.org/10.1155/2012/650574. B. Engin, C. Aydas, H. Demirtas, ESR dosimetric properties of window glass, Nuclear Instruments and Methods in Physics Research Section B 243 (2006) 149–155. G. Socol, A.C. Galca, C.R. Luculescu, A. Stanculescu, M. Socol, N. Stefan, E. Axente, L. Duta, C.N. Mihailescu, V. Craciun, D. Craciun, V. Sava, I.N. Mihailescu, Tailoring of optical, compositional and electrical properties of the Inx Zn1−x O thin films obtained by combinatorial pulsed laser deposition, Digest Journal of Nanomaterials and Biostructures 6 (2011) 107–115. L. Lutterotti, Total pattern fitting for the combined size–strain–stress–texture determination in thin film diffraction, Nuclear Instruments and Methods in Physics Research Section B 268 (2010) 334–340. G.E. Stan, I. Pasuk, M.A. Husanu, I. Enculescu, S. Pina, A.F. Lemos, D.U. Tulyaganov, K.E.L. Mabrouk, J.M.F. Ferreira, Highly adherent bioactive glass thin films synthesized by magnetron sputtering at low temperature, Journal of Materials Science. Materials in Medicine 22 (2011) 2693–2710.

49

[23] Draft International Standards [ISO/DIS], Implants for surgery – hydroxyapatite ceramic: Part 1 and 2, 1999, 13779. [24] G.E. Stan, C.O. Morosanu, D.A. Marcov, I. Pasuk, F. Miculescu, G. Reumont, Effect of annealing upon the structure and adhesion properties of sputtered bio-glass/titanium coatings, Applied Surface Science 255 (2009) 9132–9138. [25] E. Landi, G. Celotti, G. Logroscino, A. Tampieri, Carbonated hydroxyapatite as bone substitute, Journal of the European Ceramic Society 23 (2003) 2931–2937. [26] A. Kaflak, A. Slosarczyk, W. Kolodziejski, A comparative study of carbonate bands from nanocrystalline carbonated hydroxyapatites using FT-IR spectroscopy in the transmission and photoacoustic modes, Journal of Molecular Structure 997 (2011) 7–14. [27] W.E. Klee, G. Engel, I.R. spectra of the phosphate ions in various apatites, Journal of Inorganic and Nuclear Chemistry 32 (1970) 1837–1843. [28] D. Muresan, M. Vasilescu, I. Balasz, C. Popa, W. Kiefer, S. Simon, Structural investigation of calcium-soda-phosphate glasses with small content of silver oxide, Journal of Optoelectronics and Advanced Materials 8 (2006) 558–560. [29] http://www.dfdinstruments.co.uk/topics/Study5-ASTM-D4541.htm (last accessed 30.08.2012). [30] Q. Bao, C. Chen, D. Wang, Q. Ji, T. Lei, Pulsed laser deposition and its current research status in preparing hydroxyapatite thin films, Applied Surface Science 252 (2005) 1538–1544. [31] American Society for Testing and Materials [ASTM], Standard specification for composition of ceramic hydroxylapatite for surgical implants, F 1185-03 (2009) 514-515. [32] Food and Drug Administration [FDA], Calcium phosphate (Ca–P) coating draft guidance for preparation of FDA submissions for orthopedic and dental endooseous implants, 1997, pp. 1–14. [33] Y. Cai, S. Zhang, X. Zeng, M. Qian, D. Sun, W. Weng, Interfacial study of magnesium-containing fluoridated hydroxyapatite coatings, Thin Solid Films 519 (2011) 4629–4633. [34] A.C. Galca, M. Secu, A. Vlad, J.D. Pedarnig, Optical properties of zinc oxide thin films doped with aluminum and lithium, Thin Solid Films 518 (2010) 4603–4606. [35] C.M. Cotell, Pulsed laser deposition and processing of biocompatible hydroxylapatite thin films, Applied Surface Science 69 (1993) 140–148. [36] H. Zeng, W.R. Lacefield, XPS, EDX and FTIR analysis of pulsed laser deposited calcium phosphate bioceramic coatings: the effects of various process parameters, Biomaterials 21 (2000) 23–30. [37] G. Spence, S. Phillips, C. Campion, R. Brooks, N. Rushton, Bone formation in a carbonate-substituted hydroxyapatite implant is inhibited by zoledronate: the importance of bioresorption to osteoconduction, Journal of Bone and Joint Surgery: British Volume 90 (2008) 1635–1640. [38] A.H. Roufosse, W.J. Landis, W.K. Sabine, M.J. Glimcher, Identification of brushite in newly deposited bone mineral from embryonic chicks, Journal of Ultrastructure Research 68 (1979) 235–255. [39] G. Socol, P. Torricelli, B. Bracci, M. Iliescu, F. Miroiu, A. Bigi, J. Werkmann, I.N. Mihailescu, Biocompatible nanocrystalline octacalcium phosphate thin films obtained by pulsed laser deposition, Biomaterials 25 (2004) 2539–2545. [40] L. Torrisi, R. Setola, Thermally assisted hydroxyapatite obtained by pulsed-laser deposition on titanium substrates, Thin Solid Films 227 (1993) 32–36.