Characterization of Hydroxyapatite Films Prepared by Pulsed Laser Deposition Quanhe Bao,†,‡ Chuanzhong Chen,*,† Diangang Wang,† and Junming Liu§ School of Materials Science and Engineering, Shandong UniVersity, Shandong Ji’nan, 250061 P. R. China, School of Materials Science and Engineering, Anhui UniVersity of Technology, Anhui Maanshan, 243002 P. R. China, and National Laboratory of Solid State Microstructures, Nanjing UniVersity, Jiangsu Nanjing, 210093 P. R. China
CRYSTAL GROWTH & DESIGN 2008 VOL. 8, NO. 1 219–223
ReceiVed February 12, 2007; ReVised Manuscript ReceiVed September 25, 2007
ABSTRACT: Hydroxyaptite (HA) films were prepared by pulsed laser deposition from two kinds of targets using a KrF excimer laser at a substrate temperature of 575 °C. The characterization of films was studied by an electron probe microanalyzer, scanning electron microscopy, atomic force microscopy, glancing angle X-ray diffraction, and Fourier transform infrared spectroscopy. In vitro tests were performed to study the biocompatibility of the films. The results showed that the films were crystalline and there was a substitution of CO32- for PO43-. In vitro tests showed that with increased immersion time, the crystallinity of films became low, and the Ca/P ratio of films decreased first and then stayed at a stable value which is lower than that of the theoretical value of hydroxyapatite. The bands of new precipitated bonelike apatite were the same as with the original HA films, which indicated that the films had a good biocompatibility. The sintered temperature of the target had a small effect on the microstructure, while the coarse film could increase the bonelike apatite formation rate, which indicated that the bioactivity of coarse film was better.
1. Introduction Hydroxyapatite (HA) is one of the most attractive materials for human hard tissue implants because of its close resemblance to chemical composition of human bones. Nevertheless, due to the poor mechanical properties of bulk HA ceramic, it cannot be used as implant devices to replace large bony defects or for load-bearing applications. To overcome its problems of brittleness, HA was used as coatings on metallic substrates such as titanium or Ti-6Al-4V. The HA coatings permit a fast osteointegration and a protective layer against the release of metal ions. Various deposition techniques such as plasma spraying,1,2 ionbeam sputtering,3,4 and sol–gel5,6 have been applied to obtain HA coatings. The crystallinity of HA coatings prepared by these techniques were low or amorphous. The low crystallinity or amorphous led to a fast dissolution and enhanced loss of adherence of the titanium surface and dramatic implant failure. It was reported that osteoblasts cultured on the crystalline surface were more active in vitro,7 and the HA coatings having high crystallinity showed the higher shear strengths between the implant and bone in vivo studies.8 Ideally, HA coatings should have a high degree of crystallinity and a good adhesion to the substrate. Recently, pulsed laser deposition (PLD) has been applied to obtain crystalline HA films.9–12 Simply the adjustment of PLD parameters can deposit hydroxyapatite coatings of different phases and composition. In this paper, we performed a study of HA films deposited by pulsed laser and its biocompatibility in simulated body fluid (SBF). The microstructures and compositions of films were evaluated by electron probe microanalyzer (EPMA), scanning electron microscopy (SEM), atomic force microscopy (AFM), and Fourier transform infrared spectroscopy (FT-IR). * Corresponding author. Tel: +86-531-88395991. E-mail:
[email protected]. † Shandong University. ‡ Anhui University of Technology. § Nanjing University.
2. Experimental Section The HA powder (Sichuan University) was made into disks 2 mm thick and 25 mm in diameter by uniaxial and cold isostatical pressing at 300 MPa. Then one group of targets was used directly, and another group of targets was sintered at 1200 °C for 120 min. The grade two Ti was selected as a substrate material owing to its excellent mechanical properties, corrosion resistance, and good biocompatibility. They were cut into pieces 1 mm thick and 16 mm in diameter. Before deposition, they were cleaned in a sequence of ultrasonic baths of acetone and ethanol. After that, they were placed 4 cm in front of the target and kept at a temperature of 575 °C for deposition. We used a KrF (Lambda Physik Compex 201) exicmer laser source generating pulses of τfwhm e 30 ns at λ ) 248 nm. The frequency repetition rate was of 10 Hz, and the output power of laser was 170 mJ. The experiments were performed in stainless steel with water vapors of 45 Pa. To avoid target drilling during deposition, the targets were rotated at a speed of 12 rpm. To obtain a crystalline films in situ, water atmosphere 45 Pa was used during deposition. The HA films were immersed in a SBF proposed by Kokubo et al.13 to verify biocompatibility. The SBF was prepared by dissolving chemical reagents in the following, NaCl, NaHCO3, KCl, K2HPO4 · 3H2O, MgCl2 · 6H2O, CaCl2, Na2SO4 into deionized water. The fluid was buffered at pH ) 7.4 with tris-hydroxymethyl aminomethane and hydrochloric acid at 36.5 °C. HA films were immersed in a polyethylene bottle for periods of 7, 14, and 28 days. Every three days, the SBF fluid was refreshed. A JXA-8800R EPMA with a Link ISIS300 energy spectrum analyzer, JSM-6380LA SEM and AFM were used to give the secondary electron (SE) image and determine the element composition of the films. Glancing angle X-ray diffraction (GAXRD) was performed on the films using an X’Pert Philips diffractometer employing Cu KR radiation with a grazing angle of 0.5°. A NICOLET avatar370 FT-IR was used to study the functional groups of films.
3. Results and Discussion Figure 1 shows morphologies of pulsed laser deposited HA films. The films showed a typical image of pulsed laser deposited films. There were many particles with different sizes in the films. In Figure 1c,d, the scratch lines were caused by substrate but not the morphologies of deposited films. Comparing films obtained with different targets, we can see that there were little amounts of particles in films obtained with the target sintered
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Figure 3. XRD patterns of pulsed laser deposited HA films with different targets (other peaks corresponding to HA): (a) nonsintered target, (b) target sintered at 1200 °C.
Figure 1. Morphologies of pulsed laser deposited HA films with different targets: (a, b) nonsintered targets, (c, d) targets sintered at 1200 °C.
Figure 4. FT-IR spectra of pulsed laser deposited HA films with different targets: (a) nonsintered target, (b) target sintered at 1200 °C.
Figure 2. AFM topographs of pulsed laser deposited HA films with different targets: (a) nonsintered target, (b) target sintered at 1200 °C.
at 1200 °C. It may be caused by the target density increasing due to the sintered temperature increasing. The dense target was difficult to form a splashing of the molten layer, which was caused by the presence of a subsurface superheated layer. Splashing is one of the major problems of pulsed laser deposited films such as YBCO, PZT, and other films.14 However, it is not a problem if the particles adhere well to the substrate, which can increase the bonding strength between the films and new bones, owing to the increase of contract areas. Figure 2 shows AFM topographs of pulsed laser deposited HA films. It can be seen from Figure 2 that the HA films consisted of relatively dense surface layers, and particle sizes ranged from 40 nm to a few micrometers. The films deposited with targets sintered at high temperature showed a relatively smooth surface. Figure 3 shows X-ray diffraction (XRD) patterns of pulsed laser deposited HA films. It can be seen that the films consisted of high crystallinity HA, and there were some peaks corresponding to substrate. The results revealed that the high
crystallinity HA films could be prepared under 575 °C and an atmosphere of H2O at 45 Pa. Comparing Figure 3a,b, we find that density of the HA target had a little effect on the crystallinity of HA films. The films prepared by targets sintered at 1200 °C had relatively broad peaks, which means that the crystal sizes were relatively small. It verified the results of AFM. The Ca/P ratios of HA films deposited by nonsintered targets and targets sintered at 1200 °C were 1.67 and 1.82, respectively. The high Ca/P ratio of films deposited by targets sintered at 1200 °C may be caused by the decomposition of HA during sintering or the vaporing of P2O5, which was a product of HA decomposing during laser ablation. The film deposited by nonsintered targets had more particles, while the particle was caused by splashing of the target, which retains the stoichiometric ratio of the target. Hence, the Ca/P ratio of film deposited by nonsintered target was relatively low and close to the theoretical value of HA. Figure 4 shows FT-IR spectra of pulsed laser deposited HA films. The results showed that there were main vibration modes of carbonated hydroxyapatite. For PO43- groups, the υ4 vibration (asymmetric bending) was located at 565 and 603 cm-1, υ1 vibration (symmetric stretching) occurred at 960, while 1022 and 1110 cm-1 corresponded to υ3 vibration (asymmetric stretching). For CO32- groups, υ vibration (asymmetric bending) was located at 866, while 1418 and 1457 cm-1 corresponded to υ3 vibration (asymmetric stretching). In addition, 3567 cm-1 corresponding to OH- was observed, which indicated the films
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Figure 5. Morphologies of the films prepared by nonsintered HA targets after immersion in SBF for (a, b) 7 days; (c, d) 14 days; (e, f) 28 days.
were HA but not oxyapatite, owing to absence of OH- groups in oxyapatite. The existence of 1418 and 1457 cm-1 bands indicated that there was a substitution of CO32- for PO43-, which was already present in the HA target.15 The substitution could be explained as follows: Ca10-x+u(PO4)6-x(CO3)x(OH)2-x+2u where, to maintain the charge balances, the incorporation of one CO32- ion in the place of a PO43- ion causes the loss of one calcium and one OH- group; although, owing to the smaller size of the CO32- compared with the PO43-, in some lattice sites one calcium and two OH- groups are reincorporated.16 Comparing Figure 4a,b, we find that the target sintered temperature has a small effect on functional groups of the films; the FT-IR spectra of two films are similar. The bands in films are tense and sharper, which means that the crystallinity of films is high. It confirmed the results of XRD. Usually, there are two methods to evaluate the biocompatibility of HA films, one is in vitro testing17–19 and another one is in vivo testing.8,20 The in vivo testing needs an animal implantation and a long time, so in this paper we used the in vitro test to study the biocompatibility of pulsed laser deposited HA films. Figures 5 and 6 show morphologies of films after immersion in SBF for different periods of times. There were many fine crystal grains precipitated on the films. And with increased immersion time, the grains conglomerated forming relative big particles. When the films were immersed after 28 days, a new precipitated layer had covered the whole films. There were some
areas (white areas in Figure 5a) of dissolution during immersion testing, and these may be caused by the dissolution of noneHA phases and amorphous phases. Comparing Figures 5 and 6, it can be seen that there were more grains precipitated on films deposited by targets sintered at 1200 °C, owing to the rough surface of the films which can increase the dissolution rate and provide energy for crystal grain growth. Figures 7 and 8 show XRD patterns of films after immersion in SBF for different times. We can see that the films consisted of crystalline HA phase after immersion in SBF. With increased immersion time, the peaks corresponding to HA became broad indicating that the crystallinity of the new precipitated phase was low or amorphous. Kim et al.21 observed that the precursor phases of apatite such as amorphous calcium phosphate (ACP) were formed when HA was immersed in SBF, and then ACP crystallized into bonelike apatite. Figure 9 shows the changing of ratio of Ca/P during immersion. It can be seen that the ratio of Ca/P decreased first and then stayed at a stable value, which may be caused by the dissolution of ACP and tetracalcium phosphate (TetraCP) in films, which cannot be found in the XRD results. Cleries et al.18 found that calcium phosphate coatings obtained by pulsed laser deposition show different dissolution behaviors depending on the phases present in them. Highly crystalline HA coatings do not dissolve. TetraCP and R-TCP dissolve completely in multiphasic coatings, leading to microporosity, whereas HA and β-TCP remain within the coating. Amorphous coatings show complete dissolution. Once formed in SBF which is supersaturated with respect to the apatite,22 the apatite grows spontane-
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Figure 6. Morphologies of the films prepared by HA targets sintered at 1200 °C after immersion in SBF for (a, b) 7 days; (c, d) 14 days; (e, f) 28 days.
Figure 7. XRD patterns of the films prepared by nonsintered HA targets after immersion in SBF for (a) 7 days; (b) 14 days; (c) 28 days.
Figure 8. XRD pattern of the films prepared by HA targets sintered at 1200 °C after immersion in SBF for (a) 7 days; (b) 14 days; (c) 28 days.
ously consuming the calcium and phosphate ions, incorporating minor ions such as sodium, magnesium, and carbonate, and thereby developing bone mineral-like compositional and structural features. The new precipitated bonelike apatite was a Ca-poor phase, and hence the ratio of Ca/P stayed at a stable value when films
were immersed in SBF after 7 days. Kim et al.21 found that the formation of bone like apatite in sequence of Ca-rich amorphous, Ca-poor amorphous, and finally bonelike apatite. Figures 10 and 11 show FT-IR spectra of films after immersion in SBF for different times. Compared with the FTIR spectra of films before immersion (Figure 4), we find that
Pulsed Laser Deposition of Hydroxyapatite Films
Figure 9. Ca/P ratios of the films after immersion in SBF: (a) nonsintered target, (b) target sintered at 1200 °C.
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was studied by in vitro testing. The results revealed that the high crystallinity HA films could be prepared under 575 °C and an atmosphere of H2O at 45 Pa. There was a substitution of CO32- for PO43- in the films. In vitro tests showed that the bioactivity of films was good and with increased immersion time, the crystallinity of films became low and the Ca/P ratio of films decreased first and then stayed at a stable value, which is lower than that of the theoretical value of HA and was near the value of human bones. The bands of new precipitated bonelike apatite were same that of the original HA films, which indicated that the films had a good biocompatibility. The sintered temperature of the target had a small effect on microstructure, while the coarse film could increase the bonelike apatite formation rate, which indicated that the bioactivity of film was better. Acknowledgment. This work is part of a research program financed by Promotional Fund of Scientific Research for Middleaged and Youthful Scientists of Shandong Province (Project No. 02BS056).
References
Figure 10. FT-IR spectra of the films prepared by nonsintered HA targets after immersion in SBF for (a) 7 days; (b) 14 days; (c) 28 days.
Figure 11. FT-IR spectra of the films prepared by targets sintered at 1200 °C after immersion in SBF for (a) 7 days; (b) 14 days; (c) 28 days.
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4. Conclusion Pulsed laser deposition has been used to prepare HA films using two kinds of targets, and the biocompatibility of films
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