In vivo implantation of porous titanium alloy implants coated with magnesium-doped octacalcium phosphate and hydroxyapatite thin films using pulsed laser depostion  z,1 Bogusław Budner,1 Renata Syroka,1 Kryspin Niedzielski,2 Grzegorz Golan  ski,2 Waldemar Mro 3 4 5  sarczyk, Dieter Schwarze, Timothy E. L. Douglas * Anna Slo 1

Institute of Optoelectronics, Military University of Technology, 00-908 Warsaw, Poland  dz, Poland Clinic of Orthopaedics and Traumatology, Polish Mother’s Memorial Hospital Research Institute, 93-338 Ło 3  w, Poland Faculty of Material Science and Ceramics, AGH University of Science and Technology, 30-059, Krako 4 € beck, Germany SLM Solutions GmbH, Roggenhorster Straße 9c, 23556 Lu 5 Department of Biomaterials, Radboud University Medical Center Nijmegen, 6500 HB Nijmegen, the Netherlands 2

Received 25 September 2013; revised 12 February 2014; accepted 30 March 2014 Published online 00 Month 2014 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.b.33170 Abstract: The use of porous titanium-based implant materials for bone contact has been gaining ground in recent years. Selective laser melting (SLM) is a rapid prototyping method by which porous implants with highly defined external dimensions and internal architecture can be produced. The coating of porous implants produced by SLM with ceramic layers based on calcium phosphate (CaP) remains relatively unexplored, as does the doping of such coatings with magnesium (Mg) to promote bone formation. In this study, Mg-doped coatings of the CaP types octacalcium phosphate and hydroxyapatite (HA) were deposited on such porous implants using the pulsed laser deposition method. The coated implants were subsequently implanted in a rabbit femoral defect model for 6 months. Uncoated implants served as a reference material. Bone–implant contact and bone volume in the region of interest were evaluated by histopathological techniques using a tri-chromatographic

Masson–Goldner staining method and by microcomputed tomography (lCT) analysis of the volume of interest in the vicinity of implants. Histopathological analysis revealed that all implant types integrated directly with surrounding bone with ingrowth of newly formed bone into the pores of the implants. Biocompatibility of all implant types was demonstrated by the absence of inflammatory infiltration by mononuclear cells (lymphocytes), neutrophils, and eosinophils. No osteoclastic or foreign body reaction was observed in the vicinity of the implants. lCT analysis revealed a significant increase in bone volume for implants coated with C 2014 Wiley Mg-doped HA compared to uncoated implants. V Periodicals, Inc. J Biomed Mater Res Part B: Appl Biomater 00B: 000–000, 2014.

Key Words: titanium (alloys), in vivo, ceramic, coating(s), bone

 z W, Budner Bł, Syroka R, Niedzielski K, Golan  ski G, Slo  sarczyk A, Schwarze D, Douglas, TEL 2014. How to cite this article: Mro In vivo implantation of porous titanium alloy implants coated with magnesium-doped octacalcium phosphate and hydroxyapatite thin films using pulsed laser depostion. J Biomed Mater Res Part B 2014:00B:000–000.

INTRODUCTION

This article deals with development of coatings for porous titanium alloy (Ti6Al4V) implants for bone contact with an emphasis on in vivo evaluation of their osseointegration, that is, the formation of a direct interface between an implant and bone without intervening soft tissue. The highlights of the article are (1) additive manufacturing of porous Ti6Al4V implants with well-defined external dimensions and internal architecture; (2) coating of implants with different types of calcium phosphate (CaP), more specifically Mg-doped hydroxyapatite (HA) and octacalcium phosphate

(OCP), in order to facilitate osseointegration; and (3) the deposition of the coatings by pulsed laser deposition (PLD). Titanium and its alloys (e.g., Ti6Al4V) are widely used implant materials in orthopedics and dentistry. Porous Ti6Al4V scaffolds offer the following advantages over nonporous scaffolds: (i) a greater surface area for bone contact, (ii) the possibility of bone ingrowth into the pores, improving mechanical interlocking between implant and bone, and (iii) lower stiffness, reducing the mismatch in the stiffnesses of bone and implant and thus the risk of stress shieldinginduced bone loss.1,2 One method used to produce porous

*Present address: Polymer Chemistry and Biomaterials (PBM) group, Ghent University, Campus Sterre, Building S4, Krijgslaan 281, 9000 Gent, Belgium Correspondence to: T. E. L. Douglas (e-mail: [email protected]) Contract grant sponsor: Eureka (57/N-Eureka/2007 project)

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Ti6Al4V scaffolds is the rapid prototyping technique SLM. SLM involves melting of Ti6Al4V powder using a very precise nanolaser beam in order to build up a structure, layerby-layer, with highly defined external dimensions and internal architecture.3–6 Advantages of SLM include exact tailoring of: (a) implant dimensions to bone defect dimensions; (b) overall porosity and pore diameter to yield desirable mechanical properties;4 (c) pore sizes to promote bone ingrowth;9,10 and (d) strut thickness and alignment to achieve anisotropy similar to that found in natural bone tissue, which affects mechanical properties and cell colonization.11 SLM-produced scaffolds have supported bone regeneration, osseointegration, and bone ingrowth into pores in vivo.12–14 Bioceramic layers based on CaP have been applied as coatings on titanium-based implants, most often in the form of HA (Ca10(PO4)6(OH)2), which has resulted in superior osseointegration. Titanium-based implants have been coated with HA due to its bioactivity, that is, ability to form a direct chemical bond with surrounding bone.15–17 Porous titanium-based scaffolds coated with HA have shown greater bone ingrowth compared to uncoated implants.14 OCP (Ca8H2(PO4)6
5H2O) is a precursor in the formation of HA in vivo and has converted to HA after implantation in vivo and under physiological conditions in vitro.18,19 Compared to HA, OCP has shown superior ability to support new bone formation.19,20 Hence, OCP appears to be a promising alternative to HA as a coating material for titanium-based scaffolds. Indeed, there are reports of OCP coatings on titaniumbased scaffolds,21–23 but there are few, if any, reports of OCP coatings on porous titanium-based scaffolds. OCP and HA have both been deposited as coatings onto metallic implant surfaces by PLD.24–27 This method allows control of the stoichiometry of the deposited material by selection of the appropriate process parameters, especially the substrate temperature. This temperature affects the type of CaP layer produced, as the temperatures of formation of HA and OCP are 430 6 30 and 150 6 30 C, respectively.24,25,27–29 A further advantage is the ability of PLD to coat nonsmooth and porous substrates, such as porous scaffolds produced by SLM. Furthermore, the bioceramic coatings deposited by PLD can be easily doped with ions such as Mg21 without compromising the stability and biocompatibility of the coating.29,30 In vitro, osteoblast adhesion has been improved by addition of Mg21 to HA31 and carbonated apatite,32,33 while a stimulatory effect of Mg21 on proliferation has been reported for Mg-doped tricalcium phosphate,34,35 OCP,26 and HA.36–38 In vivo, Mg-substituted HA as a coating has promoted early osseointegration of titanium implants after 2 weeks compared to pure HA coatings,39 while Mg-substituted HA granules showed greater resorption and superior bone ingrowth after 2 and 3 months compared to pure HA.40 This study aimed to assess the in vivo response to SLMproduced Ti6Al4V implants coated with Mg-substituted OCP and HA in a rabbit femoral defect model. To our best knowledge, no comparative in vivo study on Mg-substituted OCP and HA has been published. A comparative analysis of the

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FIGURE 1. Porous titanium-based implant used in this study. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]

quality of the connection was performed at the implant– bone interface on the basis of histopathological examination and microcomputed tomography (lCT) imaging studies. MATERIALS AND METHODS

Preparation of implants and deposition of layers of OCP and HA In these studies titanium implants were used with a porous structure and dimensions: 10.3 mm 3 2.5 mm 3 2.5 mm (Figure 1) made by SLMV (Trademark of D. Schwarze, SLM Solutions GmbH, Luebeck, D). Pore sizes in the range 280– 420 nm were chosen, because these dimensions have been demonstrated to favor bone ingrowth into both metallic and ceramic scaffolds.9,10 Porous Ti6Al4V scaffolds were produced using SLMV using a SLM 250HL device. The desired geometry of the scaffolds was planned using computer-aided design (CAD) techniques. Scaffolds were built up layer-by-layer on a Ti6Al4V substrate platform using a cyclic process involving application of a layer of Ti6Al4V powder with average particle size 30 lm on the platform, laser irradiation of the powder layer at selective points corresponding to the CAD-derived geometry, resulting in melting of the powder and its fusion with the underlying scaffold layer, and finally lowering of the platform by the thickness of one layer to allow application of the next powder layer and repetition of the cycle. The application of R

R

TABLE I. Parameters for the Deposition of Layers of Octacalcium Phosphate (OCP) and Hydroxyapatite (HA) Parameter

OCP 1 0.6 wt % Mg

HA 1 0.6 wt % Mg

Distance from 60 6 1 mm 60 6 1 mm sample to target Target material Ti6Al4V Ti6Al4V 410 C Average temperature 186 C of implant Deposition time 144 min 122 min Number of laser shots 3000 2200 Laser repetition 3 Hz 3 Hz Average pressure 3.1 3 1022 mbar 3.2 3 1022 mbar in experimental chamber Atmosphere Water vapor Water vapor Laser pulse energy 367 mJ 367 mJ Spot size 7.9 mm2 7.9 mm2

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TABLE II. Overview of Implant Sample Groups Serial Number 0 I II

Composition of Implant Surface

Formula

Quantity (Units)

Uncoated implant OCP (octacalcium phosphate) doped with Mg HA (hydroxyapatite) doped with Mg

– Ca8H2(PO4)6
5H2O 1 (0.6 wt % Mg) Ca10(PO4)6(OH)2 1 (0.6 wt % Mg)

7 8 8

each new powder layer of thickness 50 lm lasted approximately 7 s. An yttrium-doped fiber laser with a maximum power output of 400 W and focus of 90 lm laser spot diameter was used. The process took place in a protective argon atmosphere under exclusion of air and was monitored optically. Implants, before the deposition process, were subjected to sterilization in radiation of plasma generator of the type RF-13.56 MHz in an argon atmosphere using the following settings: sterilization and activation time 5 min, RF generator power PRF  30 W, argon pressure p  9 3
1022 mbar, and argon flow v  20 cm3/min. Vapor deposition parameters for OCP and HA are shown in Table I. Three experimental groups were prepared, namely uncoated implants, implants coated with Mg-doped OCP and implants coated with Mg-doped HA, hereafter referred to as groups 0, I, and II, respectively (Table II). Mg content of OCP and HA was 0.6 wt %. OCP and HA powders for deposition were prepared as described previously.41 The individual implants within groups were designated the numbers .1, .2, 3, and so forth. For example, II.6 refers to group II, implant number 6. Group 0 served as the control group. Since the goal of the study was to compare coatings on an implant material, and not the implant material itself, a sham group (without any implant) was not considered to be indispensable. Surgical procedure In vivo studies involving placement of implants in the femoral condyles of rabbit hindpaws were conducted using 23 male rabbits of the type white New Zealand (long ears) whose mass ranged from 3.8 to 6.0 kg (mean 4.78 kg). The animals were approximately 1 year of age at the time of the surgical procedure. Ethical approval was obtained from the Local Ethical Committee for the Affairs of Experiments on Animals in Lodz (22.12.2008, decree No. 56/ŁB 440/2008). Animals were housed in individual, marked cages with free access to feed and water. Operations were performed at the Experimental Animal Laboratory at Medical University of Lodz in accordance with current surgical principles with respect to asepsis and antisepsis. Xylazine (Sedazine; Biowet, Pulawy, PL) was administered intramuscularly as premedication at a dose of 8 mg/kg. The operated region was shaved and subsequently animals were placed under general anesthesia by intramuscular administration of a mixture of ketamine [10% ketamine (50 mg/kg) and xylazine (10 mg/kg)]. After anesthesia was achieved, the field of operation was disinfected and covered with a disposable surgical barrier fabric with self-gluing edges. Cefamandol (Tarcefandol; Polfa-Tarchomin, Warszawa, PL) was administered intravenously via the marginal vein of

the ear at a dose of 30 mg/kg. Subsequently infiltration anesthesia was performed using approximately 1 mL of 1% lidocaine with added adrenaline (Xylodont 2% with adrenaline 1:80,000; Molteni Dental s.r.l., Milano, I). In individual cases where prolongation of anesthesia was necessary, a mixture of ketamine and sedazine amounting to 1/3 of the initial dose was administered intramuscularly. After completion of the operation, buprenorphine (Bunondol; Polfa, Warszawa, PL) was administered subcutaneously as an analgesic at a dose of 40 lg/kg. A skin incision of length 1.5–2.5 cm was made on the lateral side immediately above the knee joint of the hindpaw. The lateral region of the end of the distal femur was exposed by peeling off the periosteum together with the muscle attachments taking care not to open the cavity of the knee joint. An opening of depth 10 mm perpendicular to the axis of the bone was bored in the cortical substance and spongiosa using drills of diameter 1.5 mm mounted on a stomatological drive mechanism. Subsequently, the cortical substance was drilled open using a drill of diameter 2.5 mm. In order to obtain an opening corresponding to the shape of the implants, the round opening obtained was broken further open using a bone awl, whose tip had the shape of a pyramid with a square base. This was plunged in to the desired depth in order to obtain a square opening of side length 2.5 mm in the cortical substance. Subsequently, the implant was inserted (Figure 2) by pushing it in manually or by applying delicate mallet strikes. The goal was to place the implants such that they would pass through the lateral cortical substance and anchor in the trabecular bone at the base of the distal femur. After placing of the implants, the wound was rinsed with a solution of cefamandol [0.5 g/20

FIGURE 2. Placement of a titanium-based implant in the femoral condyle of a rabbit hind paw. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]

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FIGURE 3. Light microscopy images of implants of reference group 0 after Masson–Goldner staining. (a; left) Sample number 0.4, 203 magnification, bone (blue) visibly surrounds the titanium implant (black) on the outside and penetrates into the pores of the implant. (b; right) Sample number 0.1, 1003 magnification, contact between bone (blue) and implant (black) is visible. Scale bars: (a) 500 lm and (b) 100 lm. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]

mL (Tarcefandol)], and sutured layer by layer using Safil (Braun) 4/0 suture for the deep tissue and Dafilon (Braun) 3/0 suture for the skin. A sterile dressing was applied. No immobilization of the limbs was used in view of the negligible loss in mechanical strength of the femur. After the planned 6-month observation period, animals were euthanized and implants with surrounding tissue were retrieved for evaluation of the endosseous healing and osseointegration of the implants. Implants were fixed securely by immersion in polymethacrylates (Technovit 9100 NEW, Heraeus Kulzer). Subsequently, after crystallization of the methacrylate with the implant, the samples were cut to the size required for further investigations using a saw with a diamond disc Leica SP 1600 (Leica). In order to evaluate the connection at the interfaces between titanium and calcified bone and osteoid tissue, which is not calcified tissue, tests were performed using three chromatographic staining methods according to Masson–Goldner.

Planimetric analysis of integration of implant with bone Planimetric (histomorphometric) investigations and calculations were performed to evaluate the bone–implant integration and assess whether differences existed between implants with different nanolayers. Analysis of the results was performed using the statistical package StatSoft STATISTICA 9.0. Length of bone–implant interface and ratio of contact length to implant circumference were calculated. The results of histomorphometric analysis in dependence of sample group were analyzed using a Kruskal–Wallis test, and a comparison of the groups with various modifications of the titanium surface with the reference group was performed using the Dunn test. Additionally, a univariate analysis of variance (ANOVA) and a comparison of groups with various modifications of the titanium surface with the reference group were performed using Dunnett’s test. A probability level of

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p < 0.05 was considered to be significant, and a level of p < 0.01 was considered highly significant.

Qualitative and quantitative evaluation of bone–implant integration using lCT Analysis of the bone–implant integration was performed by lCT using a SkyScan 1172, equipped with a camera matrix with 11 megapixels, with a pixel size of 8.81 lm. The following parameters were applied: current radiation source 100 lA, voltage 100 kV, exposure time 560 ms, sample rotation step 0.5, averaging at 15 fps, use of aluminum 1 copper filter, output file format .tif with 16 bit depth, with a pixel size of 2000 3 1200. Files obtained from the study were reconstructed using NRecom version 1.6.4.1 software from the scanner’s manufacturer. For quantitative analysis, a space [volume of interest (VOI)] containing the implant in the center was selected from a stack of lCT scans. The VOI as determined to be a cube of side length 3.9 mm. The three-dimensional picture of the VOI, consisting of 1102 lCT scans in 256level grayscale images, was processed to determine the threshold to distinguish between titanium, mineralized bone tissue and background. Thresholds specific for titanium and calcified bone were determined by comparing the binary images with the application of a threshold to the original image on a gray scale according to the method proposed by Butz.42 In this study, the following thresholds were used: 70 and 150. Voxels with values below and above the threshold were classified as: background (150), whereas calcified bone was between these values (70–150). Subsequently, bone volume (Vk) was determined. Statistical results in dependence of sample group were analyzed using a Kruskal–Wallis test and comparison of groups with various modifications of the titanium surface with the reference group was performed using the Dunn test. A probability level of p < 0.05 was considered to be significant while a probability level of p < 0.01 was considered

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ORIGINAL RESEARCH REPORT

FIGURE 4. Light microscopy images of implants of sample group I after Masson–Goldner staining. (a; left) Sample number I.1, 403 magnification, bone (blue) surrounding the implant (black) is visible on the exterior walls. (b; right) Sample number I.5, magnification 1003, part of a titanium implant with the surrounding bone and the red infiltrate inflammation is visible. Scale bars: (a) 250 lm and (b) 100 lm. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]

highly significant. Analysis of results was performed using the statistical package StatSoft STATISTICA 9.0. RESULTS AND DISCUSSION

Histopathological evaluation of implants Representative images of implants in group 0 are shown in Figure 3(a,b). All samples showed tight adhesion of mineralized bone tissue to the implant, which in many cases was within the fatty marrow. In individual animals small fields of the connective tissue were seen directly adjacent to the implant from the internal structure of the implant. Direct penetration of regular bone into the pores of the implant was observed. No inflammatory infiltration consisting of mononuclear cells (lymphocytes) neutrophils or eosinophils was observed, nor was an osteoclastic reaction in the vicinity of the implant. Only in the case of implant number 0.4 was a slight "thinning" of bone structure observed in close proximity to of the implant, which may be due to the fact that the bone was not present throughout the thickness of the preparation [Figure 3(a)]. In addition, features of focal

fibrosis in the vicinity of the implant were only found in one of the samples evaluated (number 0.7—not shown). Representative images of implants in group I are shown in Figure 4(a,b). In the whole group, direct integration of the bone and penetration into the pores and the structure of the implant were observed. In the majority of animals tested there were no signs of inflammation or fibrosis in the vicinity of the implant. Only in sample number I.5 [Figure 4(b)] was inflammatory infiltration (red) observed in a field in which there was bone, however this was in the middle of the fatty marrow. In sample I.8 (not shown), small, "residual” outbreaks of fibrosis were seen extending from the implant. No osteoclastic or osteoblastic reactions were found adjacent to the implant in any of the cases. Representative images of implants in group II are shown in Figure 5(a,b). In many cases (II.1, II.2, II.4, and II.6) symptoms of focal fibrosis and osteoid fields of construction were observed, which form structures similar to "caps" in the immediate vicinity of the implant [II.4, II.6, II.7; Figure 5(b)]. In the majority of cases evidence of direct adhesion

FIGURE 5. Light microscopy images of implants of sample group II after Masson–Goldner staining. (a; left) Sample II.8, 1003 magnification, the interior of titanium mesh (black) with bone ingrowth (blue). (b; right) Sample II.6, 1003 magnification, outer periphery of the implant (black) is visible with fibrous tissue (red) between the implant and bone (blue). Scale bars: (a) 100 lm and (b) 100 lm. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]

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FIGURE 6. (a) Length of bone–implant interface. (b) Bone volume (Vk). Red dots indicate median values. Boxes indicate 25–75% range. Bars indicate minimum and maximum values within the sample distribution. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]

(integration) of bone to the implant was observed. In one of the samples (II.5—figure omitted) weaving bone marrow was observed. In this group, there was no sign of osteoclastic or osteoblastic reaction around the implant.

group II with the reference group 0 (p 5 0.0311). The result for group II was significantly higher than for the reference group 0. DISCUSSION

Results of planimetric analysis The length of bone–implant contact is shown in Figure 6(a). Statistical analysis revealed that implants in group II showed the best integration with bone, as this group had the highest average length of the bone–implant interface and the best reproducibility of results. Statistical results in dependence of sample group and applied tests are shown in Table III. Among the two types of implants (groups I and II), which differed in the manner by which the surface was modified, no significant differences were observed at the p < 0.05 level relative to the reference group "0" (without implant surface modification). Hence, on the basis of these data from bone histomorphometric evaluation of implants made using optical microscopy, it is not possible to show differences influencing the integration and bone–implant contact between the groups of implants with different surface layers. Results of qualitative and quantitative investigations performed using lCT The bone volume (Vk) in the VOI is shown in Figure 6(b). The only significant difference concerned the comparison of

All the three implant groups integrated well with surrounding bone, as shown by the direct formation of new bone on the implant surfaces and the lack of inflammatory or foreign body reaction (Figures 3–5). This is in accordance with the findings of other authors who have implanted Ti6AL4V scaffolds produced by SLM into rat and rabbit calvarial defects, rabbit femoral defects and the iliac crest of goats12–14,43 lCT studies indicate a significant advantage of the implants coated with Mg-doped HA (group II) compared to uncoated implants (group 0) with regard to bone volume [Figure 6(b)]. A much lower difference was observed between OCPcoated implants (group I) compared to groups 0 and II. Reports on the in vivo effects of CaP-based coatings on porous titanium-based scaffolds have been variable. Barrere et al.21 implanted porous Ti6Al4V scaffolds coated with OCP and carbonated apatite into the femoral condyles of goats and observed superior bone ingrowth into the OCP-coated implants. Subcutaneous implantation of Ti6Al4V samples coated with OCP and carbonated apatite in Wistar rats revealed a faster dissolution rate of the OCP coating during the first week after implantation.44 Biemond et al.45 implanted porous Ti6Al4V scaffolds in the iliac crest of goats and

TABLE III. Results of Statistical Histomorphometric Analysis Variable

Group

Length of bone–implant interface

Reference I II Reference I II

Ratio of contact length to the circumference of implant

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Kruskal–Wallis Test/ANOVA

Dunn’s Test

Dunnett’s Test

1.0000 1.0000

0.8819 0.3108

1.0000 1.0000

0.7623 0.1453

0.6800/0.5382

0.6349/0.3541

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reported increased bone ingrowth on implants coated with HA and brushite compared to uncoated implants. However, a later study by the same group in the same in vivo model revealed no significant differences between uncoated implants and implants coated with amorphous HA.14 There are few comparative studies in literature concerning the in vivo effect of Mg doping of HA and, to our best knowledge, no studies on Mg-doped OCP. It remains unclear why Mg21 stimulated bone formation more effectively as a component of a HA coating than as a component of an OCP coating. The Mg content was set at 0.6 wt % for three reasons, one biomimetic, one physicochemical, and one biological. First, 0.6% content is similar to the Mg contents found in mineralized human tissues in vivo. For example, the Mg contents in bone, dentine and enamel are 0.47, 1.11, and 0.3%, respectively.46 Second, the inclusion of 0.6 wt % Mg does not appreciably alter the structure of either HA or OCP. Previous work of the authors on PLD of Mg-containing HA coatings has shown that the size of the elemental cell of HA is lowest at an Mg content of 0.6 wt %, and hence, Mg was most effectively incorporated into the structure of HA at this weight percentage.30 Previous work by Boanini et al.26 on matrix-assisted pulsed laser evaporation of Mgcontaining OCP coatings showed that the presence of 0.6 wt % Mg did not alter OCP structure. Third, incorporation of 0.6 wt % Mg is sufficient to influence osteoblast behavior positively. Boanini et al.26 observed superior osteoblast proliferation, differentiation and spreading on OCP doped with 0.6 wt % Mg. Yamasaki et al.32 showed that 0.144 mol/g, equivalent to 0.35 wt % Mg promoted osteoblast adhesion to carbonated apatite. HA preparations doped with similar amounts of Mg have promoted osseointegration and bone ingrowth. Zhao et al.39 implanted titanium implants coated with pure HA or HA containing 0.78 mol % Mg in a rabbit femoral defect. Higher bone–implant contact was seen on implants coated with HA containing Mg after 2 weeks. However no differences were seen after 4 and 8 weeks. Landi et al.40 compared granules of pure HA and HA substituted with 0.7 wt % Mg in rabbit femoral defects and observed that bone ingrowth into Mgsubstituted granules was inferior after 1 month but superior after 2 and 3 months, which was ascribed to poorer initial resorbability of the Mg-substituted granules due to its more compact morphology as a result of the substitution process. Doping of HA with Mg with has been reported to accelerate dissolution.32,33,38,40 Hence, the addition of Mg to HA remains controversial, despite the aforementioned biological benefits.

CONCLUSIONS

Based on histopathological and histomorphometrical studies it can be concluded that both the bare titanium and surface modified implants are biocompatible and integrate with the bone with good bond implant–bone connection. In some cases in all groups, penetration of newly formed bone through the mesh into the center of the implant was observed. In most of the cases studied in each individual

group, characteristics of direct adhesion (integration) of bone to the implant are obvious. In no cases was inflammatory infiltration consisting of mononuclear cells (lymphocytes) observed, nor was infiltration of neutrophils and eosinophils, which demonstrates the good biocompatibility of the implants used. No osteoclastic reaction was observed in the vicinity of the implant. In the groups studied, a tight adhesion of implant to the bone was obvious without evidence of a foreign body reaction. lCT analysis revealed that the volume of bone in the VOI was highest for group II and significantly higher than for group 0. REFERENCES 1. Niinomi M. Mechanical biocompatibilities of titanium alloys for biomedical applications. J Mech Behav Biomed Mater 2008;1:30–42. 2. Long M, Rack HJ. Titanium alloys in total joint replacement—A materials science perspective. Biomaterials 1998;19:1621–1639. 3. Mullen L, Stamp RC, Brooks WK, Jones E, Sutcliffe CJ. Selective laser melting: A regular unit cell approach for the manufacture of porous, titanium, bone in-growth constructs, suitable for orthopedic applications. J Biomed Mater Res B Appl Biomater 2009;89: 325–334. 4. Mullen L, Stamp RC, Fox P, Jones E, Ngo C, Sutcliffe CJ. Selective laser melting: A unit cell approach for the manufacture of porous, titanium, bone in-growth constructs, suitable for orthopedic applications. II. Randomized structures. J Biomed Mater Res B Appl Biomater 2010;92:178–188. 5. Murr LE, Quinones SA, Gaytan SM, Lopez MI, Rodela A, Martinez EY, Hernandez DH, Martinez E, Medina F, Wicker RB. Microstructure and mechanical behavior of Ti–6Al–4V produced by rapidlayer manufacturing, for biomedical applications. J Mech Behav Biomed Mater 2009;2:20–32. 6. Sallica-Leva E, Jardini AL, Fogagnolo JB. Microstructure and mechanical behavior of porous Ti–6Al–4V parts obtained by selective laser melting. J Mech Behav Biomed Mater 2013;26:98–108. 7. Van Bael S, Chai YC, Truscello S, Moesen M, Kerckhofs G, Van Oosterwyck H, Kruth JP, Schrooten J. The effect of pore geometry on the in vitro biological behavior of human periosteum-derived cells seeded on selective laser-melted Ti6Al4V bone scaffolds. Acta Biomater 2013;8:2824–2834. 8. Warnke PH, Douglas T, Wollny P, Sherry E, Steiner M, Galonska S, Becker ST, Springer IN, Wiltfang J, Sivananthan S. Rapid prototyping: Porous titanium alloy scaffolds produced by selective laser melting for bone tissue engineering. Tissue Eng Part C Methods 2009;15:115–124. 9. Karageorgiou V, Kaplan D. Porosity of 3D biomaterial scaffolds and osteogenesis. Biomaterials 2005;26:5474–5491. 10. Ryan G, Pandit A, Apatsidis DP. Fabrication methods of porous metals for use in orthopaedic applications. Biomaterials 2006;27: 2651–2670. 11. Spoerke ED, Murray NG, Li H, Brinson LC, Dunand DC, Stupp SI. Titanium with aligned, elongated pores for orthopedic tissue engineering applications. J Biomed Mater Res A 2008;84:402–412. 12. Van der Stok J, Van der Jagt OP, Amin Yavari S, De Haas MF, Waarsing JH, Jahr H, Van Lieshout EM, Patka P, Verhaar JA, Zadpoor AA, Weinans H. Selective laser melting-produced porous titanium scaffolds regenerate bone in critical size cortical bone defects. J Orthop Res 2013;31:792–799. 13. de Wild M, Schumacher R, Mayer K, Schkommodau E, Thoma D, Bredell M, Kruse A, Gratz KW, Weber FE. Bone regeneration by the osteoconductivity of porous titanium implants manufactured by selective laser melting: A histological and microCT study in the rabbit. Tissue Eng Part A 2013. 14. Biemond JE, Hannink G, Verdonschot N, Buma P. Bone ingrowth potential of electron beam and selective laser melting produced trabecular-like implant surfaces with and without a biomimetic coating. J Mater Sci Mater Med 2013;24:745–753. 15. Chambers B, St Clair SF, Froimson MI. Hydroxyapatite-coated tapered cementless femoral components in total hip arthroplasty. J Arthroplasty 2007;22:71–74.

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PULSED LASER DEPOSITION OF MG-ENRICHED CERAMIC LAYERS ON POROUS TI6AL4V