Polycrystalline Diamond Coating of Additively Manufactured Titanium

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Polycrystalline Diamond Coating of Additively Manufactured Titanium for Biomedical Applications Aaqil Rifai, Nhiem Tran, Desmond W. M. Lau, Aaron James Elbourne, Hualin Zhan, Alastair Stacey, Avik Sarker, Edwin Lawrence Harrop Mayes, Elena P. Ivanova, Russell J Crawford, Phong A Tran, Brant C. Gibson, Andrew D. Greentree, Elena Pirogova, and Kate Fox ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b18596 • Publication Date (Web): 22 Feb 2018 Downloaded from http://pubs.acs.org on February 25, 2018

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ACS Applied Materials & Interfaces

Polycrystalline Diamond Coating of Additively Manufactured Titanium for Biomedical Applications Aaqil Rifai,1* Nhiem Tran,2 Desmond W. Lau,3 Aaron Elbourne,2 Hualin Zhan,4 Alastair D Stacey,4 Edwin LH Mayes,5 Avik Sarker,1 Elena P. Ivanova,6 Russell J. Crawford,2 Phong A. Tran,7 Brant C. Gibson,3 Andrew D. Greentree,3 Elena Pirogova,1‡ and Kate Fox1,8*‡ 1.

School of Engineering, RMIT University, Melbourne, VIC 3001, Australia

2.

School of Science, RMIT University, Melbourne, VIC 3001, Australia

3.

ARC Centre of Excellence for Nanoscale BioPhotonics, School of Science, RMIT University, Melbourne, VIC 3001, Australia

4.

School of Physics, University of Melbourne, Parkville, VIC 3010, Australia

5.

RMIT Microscopy and Microanalysis Facility (RMMF), RMIT University, Melbourne, VIC 3000, Australia

6.

School of Science, Swinburne University of Technology, Hawthorn, VIC 3122, Australia

7.

Institute of Health and Biomedical Innovation, Queensland University of Technology, Kelvin Grove, QLD 4059, Australia

8.

Center for Additive Manufacturing, RMIT University, Melbourne, VIC 3001, Australia

* Corresponding Authors (Mr. Aaqil Rifai [email protected], Dr Kate Fox [email protected]) ‡ These authors contributed equally

Keywords: diamond, adhesive coatings, TI6Al4V, additive manufacturing, biomaterials

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Abstract Additive manufacturing using selective laser melted titanium (SLM-Ti) is used to create of bespoke items across many diverse fields such as medicine, defence and aerospace. Despite great progress in orthopaedic implant applications, such as for ‘just in time’ implants, significant challenges remain with regards to material osseointegration and the susceptibility to bacterial colonization on the implant. Here, we show that polycrystalline diamond coatings on these titanium samples can enhance biological scaffold interaction improving medical implant applicability. The highly conformable coating exhibited excellent bonding to the substrate. Relative to uncoated SLM-Ti, the diamond coated samples showed enhanced mammalian cell growth, enriched apatite deposition and reduced microbial S. aureus activity. These results open new opportunities for novel coatings on SLM-Ti devices in general, and especially show promise for improved biomedical implants.


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1. INTRODUCTION

Three-dimensional (3D) printing or additive manufacturing (AM) has gained significant commercial traction in a number of biomedical applications. AM, especially selective laser melting (SLM) of metals, has revolutionised the manufacture of many devices, particularly implants, offering an efficient use of raw materials with minimal material waste, satisfactory geometric accuracy and overall mechanical integrity for implant production.1 In contrast to conventional implant manufacturing methods, AM enables complex structures to be printed on a layer-by-layer basis. As a result, SLM has the potential to become one of the most important techniques for the efficient creation of personalised biomedical implants.2-3 SLM offers a great advantage in the development of personalized healthcare products that are tailored to the specific needs of a patient; especially in terms of the shape, size and internal structure of the implant. Such personalised solutions will lead to improvements in the health and quality of life of the general population, and reducing the burden on the healthcare system.

The SLM process uses powdered metal to fuse custom geometries layer-by-layer.4 SLM can use a variety of metals such as iron, nickel, copper, aluminium and steel,5 with cobalt chromium, stainless steel and medical grade titanium (TI6Al4V) being specifically used to manufacture orthopaedic implants.6 Further, SLM provides the necessary high quality surface properties without the additional mechanical polishing or acid etching that is usually required for improving an implant-bone tissue interface. The internal geometries and lattice structures within selective laser melted titanium (SLM-Ti) are of particular importance for printing orthopaedic implants to


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achieve secure bone attachment with high strength-to-weight ratio (mechanical stability) 6 and biocompatibility7-8 required for advanced biomedical applications.9-10

The typical lifetime of an implantable prosthesis, such as a hip implant, can range from five to twenty-five years.11 One of the main reasons for titanium implant failure is aseptic loosening12-13 of the implant or the development of wear debris.14 Eventually, the sliding force between a bone and an implant causes the implant to wear, leading to its failure.15-16 A method for improving the interface between an implant and the surrounding hard tissue is applying a secondary coating to the implant. The coating of 3D structures is, however, challenging for a number of reasons due to the complexity of underlying structures and printing parameters that affect coating adherence, thickness, roughness and morphology. As a result, these modification parameters require optimization.17-19 Ineffective coating often leads to the failure of the implant. This is particularly true where heat is applied for the coating as thermal mismatch between the substrate and the coating leads to induced stress and delamination.20-21 The lack of a well-adhered and uniform surface coating hinders osseointegration and causes an increase in the likelihood of bacterial infection.22-23 Bacterial infections are detrimental in their own right but they also lead to the increased liberation of cytotoxic metal ions at the implant-coating interface, thus resulting in a decrease of patient general health and wellbeing. The patient may in turn require future revision surgeries.24-25

The biocompatibility of polycrystalline diamond (PCD) makes it desirable for cardiovascular, orthopaedic and dental applications.14,

26-29

PCD coatings show great promise for biomedical


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applications due to their increased hardness, corrosion resistance, chemical inertness, excellent biocompatibility and antibacterial properties.30 The mechanical properties of diamond have already been shown to have promise for biomedical applications. As such, PCD coatings on SLM-Ti are potentially more advantageous when compared to the as-fabricated SLM-Ti implants. PCD is successfully used to coat knee and hip implants with the aim of increasing the biocompatibility and non-cytotoxic properties of the implant. PCD films can exhibit excellent wear resistance depending on their adhesion to the underlying substrate, roughness and coating thickness. These films have also been shown to exhibit the formation of an improved boneimplant interface.27, 31-32

Here, for the first time, we report the deposition of PCD thin films on additively manufactured SLM-Ti structures (planar plates and non-planar 3 mm3 open cubes) using chemical vapour deposition (CVD) to produce PCD coatings of 2 – 3 µm. We demonstrate that PCD coatings broaden the viability of SLM-Ti for the biomedical applications. The PCD films were evaluated for their adhesion onto complex non-planar scaffolds. The bacterial response and cellular biocompatibility of the PCD coated SLM-Ti samples were further examined by performing bacterial adhesion assessments and cell viability studies.


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2. EXPERIMENTAL SECTION

2.1 Synthesis of PCD scaffolds Sample preparation: Additively manufactured SLM-Ti substrates (10 x 10 x 0.5 mm3) and hollow SLM-Ti cubes (3 x 3 x 3 mm3, with strut diameter 1 mm) were fabricated using the method detailed by Xu et al.1 Standard cleaning of SLM-Ti substrates was performed using sequential sonication in acetone, methanol and isopropanol and drying under a steady flow of nitrogen gas.

Diamond coating: PCD samples were prepared by chemical vapor deposition CVD (iPlas) on nanodiamond-seeded (Nabond, 120 nm nanodiamond) SLM-Ti. The CVD gas (1:75, Hydrogen:Methane) mixture was 2% methane (CH4) and 98% hydrogen (H2) following the conditions outlined by Fox et al.26 and Garrett et al.33 Use of the line of-sight PCD deposition method meant that the sides of SLM-Ti substrates were also exposed to the plasma depending on position within the reactor, although such edge effects were considered to be unimportant for our study. SLM-Ti substrates were exposed to the plasma for 1 h at a stage temperature of 586.5 °C and sample temperature of 1100 °C. These temperatures were selected with respect to the optimal conditions for PCD growth identified by Fox et al.26 After the growth, the chamber was cooled down in an atmosphere of argon. The resulting PCD film thickness was approximately 12 µm.26 SLM-Ti cubes were exposed to the plasma for 1h and were not rotated. Heating of the


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cube within the plasma was controlled by placing six unseeded cubes within the plasma, with the seeded cube at the center.

2.2 Material characterization Surface characterization: Scanning electron micrographs were obtained using xT Nova Nanolab 200 and JEOL JSM-5910 scanning electron microscopes, using an acceleration voltage of 30 kV for faces and edge imaging was performed at 15 kV. Thickness of simulated body fluid (SBF) was obtained using FEI Scios DualBeam FIB by coating the desired milling site with Platinum (using Helium in chamber gas injection system GIS) to protect the area form ion beam damage. Cross-section trench was cut at a larger beam current (5 or 7 nA) then use a smaller beam current to clean the face (100 pA). Backscatter electron micrograph (z-contrast) was obtained using the T1 (Trinity 1) detector. Raman spectroscopy was used to examine the composition of the grown PCD films with a Renishaw inVia Raman Microscope with a laser wavelength of 532 nm and a laser spot size of 10 µm within the range 700 cm-1 to 2100 cm-1. The higher levels of surface roughness (Ra > 1 µm) was measured using a XP-2 Stylus Profiler (Ambios Technology, Inc) at a force 0.1 mg, scan rate of 0.01 mm/s over an area of 0.5 mm. Lower levels of roughness (Ra < 1 µm) were obtained using a NanoScope III Dimension 3100 AFM at a scan rate of 0.6 µm/s over an area of 10 x 10 µm and 30 x 30 µm. The Ra value was obtained by averaging the values obtained across each sample. Gwyddion software was used to analyze the surface roughness.


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The surface wettability of the sample was characterized using contact angle goniometry. Static water contact angle measurements were performed using Kruss, Drop Shape Analyzer of water sessile drops at room temperature. Images were acquired using the inbuilt software. The contact angles reported in this work were the average of five different spots on each sample. A PerkinElmer 2000 infrared spectrometer was used to analyze the chemical bonds in the simulated body fluid all samples. The functional group vibrations of the SLM-Ti, P-Ti and PCD coated with SBF were analysed at different wavenumbers of FTIR spectra. The FTIR spectra were obtained for 32 scans at a resolution of 4 cm-1 over a range of 0 - 4000 cm-1. Background spectra were obtained in the chamber for a sample holder without samples prior to the actual analysis. The inclination angle of the measured surface was then adjusted for high energy before scanning. The inbuilt software was used for processing the acquired spectra. At least three measurements were carried out for each sample.

2.3 Cell viability In vitro assays: The cell viability of Chinese Hamster Ovarian (CHO) cells was examined on the PCD coated SLM-Ti and bare SLM-Ti samples. CHO cells were chosen due to their reported effectiveness for the evaluation of cytotoxicity in biomaterials. In this study, the cells were genetically modified to express a green fluorescent protein (GFP) for easier detection using a fluorescent microscope. Samples were placed into a 24-well plate, then seeded with CHO cells in Alpha Modified Eagle Medium (α-MEM, supplemented with 10% fetal bovine serum (FBS) and 1% Penicillin Streptomycin (P/S) all obtained from Sigma) at the density of 40,000 cells per well. The samples were incubated for 24 hours at 37 °C under a 5% CO2 atmosphere. After


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incubation, the cells were rinsed with Phosphate Buffer Solution (PBS) and fixed using paraformaldehyde for 30 minutes. One batch was placed in PBS and held at 4°C for subsequent confocal microscopic analysis (N-STORM SuperResolution/Confocal microscope). Another batch was sequentially dehydrated using 50%, 70%, 90% 95% ethanol solution for 10 minutes each, then with 100 % ethanol for 15 minutes. This was then repeated using 100% ethanol for 30 minutes.34 After this was completed, the samples were allowed to dry overnight in a fume hood for subsequent SEM imaging (FEI Quanta 200 ESEM (2002)) imaging.

MTS assay: For MTS assay, approximately 4 × 104 cell/cm2 CHO cells in medium were seeded on flat PCD and SLM-Ti samples in a 24-well plate. The plate was incubated in a CO2 incubator for 24, 72, and 96 hrs. Then, 200 µL of [3-(4, 5-dimethylthiazol- 2-yl)-5-(3carboxymethoxyphenyl)-2- (4-sulfophenyl)- 2H-tetrazolium, inner salt; MTS(a)] solution was added to the wells. The optical density (OD) corresponds to the viable cell numbers. In this study, the level of cell proliferation for different time periods was quantified after finishing the successive incubation periods by adding the MTS solution to each of the wells. 100µl medium of the medium from each well was transferred to an ELISA 96>well plate and the absorbance was recorded at 490nm using a plate reader-spectramax paradism-molecular device.

2.4 Bacterial assessment Staphylococcus aureus (S. aureus) (CIP 65.8T) was purchased from the Culture Institute Pasteur, France. Bacterial stocks were prepared in 20% glycerol nutrient broth (Oxoid) and stored at –


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80°C until needed. Prior to experiment, the stock was refreshed upon nutrient agar (Oxoid) for 24 h at 37° C. S. aureus suspensions were prepared by suspending 1 loopful of bacteria into 5 mL of nutrient broth (Oxoid). The suspensions were then further diluted using nutrient broth to obtain an optical density OD600 = 0.1. The SLM-Ti and PCD samples were sterilised with repeated washes of ethanol then rinsed with MilliQ water and dried under a gentle N2 flow. The sterilised surfaces were then incubated, in 1 mL of bacterial suspension in a sterilised 12-well plate (In Vitro Technologies) for durations of 1 h and 18 h at 25° C.

Following incubation, confocal laser scanning microscopy (CSLM) was utilised to assess the surface attachment of live and dead bacteria. Prior to CSLM imaging, samples were removed from the bacterial suspensions and washed gently twice with MilliQ water for 3 second durations. This process removes unattached bacteria from the surface, allowing all imaging experiments to be performed under similar conditions. Surface bacteria were then stained for 30 min in the dark using LIVE/DEAD Baclight Bacterial Viability Kit, L7012 (Molecular Probes, Life Technologies) according to the manufacturer’s protocol. SYTO®9 permeates both intact and damaged membranes of the cells, and fluoresces green when bound to nucleic acids and excited by a 485 nm wavelength laser; however, propidium iodide (PI), only enters bacteria with significant membrane damage (non-viable cells) and binds with higher affinity to the intracellular nucleic acids than SYTO®9. The proportion of each type of bacteria was determined by pixel counting at their respective fluorescence emission wavelengths. CSLM images were obtained using a FV1000 Spectroscopic Confocal System (Olympus, Tokyo, Japan) with an inverted microscope at 60x magnification. Images were taken for 5 different fields of view to obtain the representative data for the entire surface. After image acquisition, the imaging


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software Fluoview FV 4.2 was utilised to assess the CSLM images on a frame-by-frame basis, which allowed the total surface-bacteria count to be calculated.

A second supporting bacterial study reported in the Supplementary Data, S. aureus (ATCC 25923) was streaked onto tryptic soy agar (TSA) plates and incubated at 37ºC overnight. A single colony was selected using a disposable loop and mixed into 10 mL of tryptic soy broth (TSB) in a centrifuge tube. The tube was placed in a shaking incubator at 200 rpm and 37ºC overnight. 2 mL of bacterial solution was collected and diluted with 8 mL of fresh TSB and was incubated (37ºC, 200 rpm) for 4 hours until the bacteria reached its logarithm growth phase. The bacterial solution was then diluted by a factor of 10 (to 8 x 107 cfu/mL) with TSB for subsequent seeding onto the substrates. PCD coated SLM-Ti substrates were placed in each well of a 24-well plate. Non-coated bare SLM-Ti substrates were used as the control. 1 mL of bacterial solution was added to each well and the plate was incubated at 37ºC and rotated at 100 rpm for 1 hour to allow the bacteria to adhere to the substrates. After 1 hour, the substrates were washed with PBS to remove any non-adhered bacteria. 500 µL of 4% formaldehyde solution was added to each well to fix the bacteria. The plate was left for 15 minutes at room temperature before the formaldehyde was removed and the samples were then twice rinsed with PBS. 500 µL of TritonX 100 solution was then added to each sample to enhance the permeability of the cell membrane for subsequent fluorescent staining. Triton-X 100 solution was removed after 15 minutes and the samples were twice rinsed with PBS before adding 500 µL of blocking solution containing 3% bovine serum albumin (BSA) to each sample to prevent non-specific binding of the dye. After 30 minutes, 100 µL of nucleus dye DAPI was added to each sample. The substrates were allowed to remain in the DAPI solution for 15 min before finally being twice rinsed with PBS to remove


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any excess dye. The samples were kept in PBS and at 4ºC until analysed using confocal fluorescent microscopy. At least three random fields on each sample were captured using DAPI filter (350 nm/470 nm). The number of bacteria present were counted using Fiji software (ImageJ, NIH).

2.5 Simulated body fluid assay A 1.0M simulated body fluid (SBF) solution having inorganic ion concentrations simulating human blood plasma (Table 1) was prepared following the methodology outlined by Kokubo et al.35 Briefly, reagent-grade chemicals (NaCl, NaHCO3, KCl, K2HPO4·3H2O, MgCl2·6H2O, CaCl2, Na2SO4) were sequentially dissolved in distilled water using a hot plate and a stirring rod. The solution was brought to physiological pH 7.35, using 1 M HCl. The SBF solution was refrigerated at 4 °C overnight to ensure that no precipitation occurred before experimental use. SLM-Ti, P-Ti and PCD samples were suspended into a polystyrene container vertically with a lid. The polystyrene container was filled with 6 mL of SBF solution with respect to the suggested volume defined by the formula Vs = Sa/10 (Vs volume of simulated body fluid (mL), Sa the surface area of the sample (mm2))

26, 35

. Samples were suspended to ensure apatite deposition

onto samples was due to heterogeneous nucleation. Samples were immersed in the SBF solution for time periods of 7 and 14 days and kept within an incubator set at 37 °C. SBF solution was not replenished across the 14 day period of experimentation. Post-removal, SBF immersed samples were washed in distilled water and dried under a flow of ultrapure nitrogen gas.


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Table 1. Concentration of ions in SBF as compared to that in human blood plasma Ion concentration

Human blood plasma (mM)

Simulated body fluid (mM)

Na+

142.0

142.0

K+

5.0

5.0

Mg2+

1.5

1.5

Ca2+

2.5

2.5

Cl¯

103.0

147.8

HCO3¯

27.0

4.2

HPO42¯

1.0

1.0

SO42¯

0.5

0.5

3. RESULTS AND DISCUSSION

3.1 PCD adherence and delamination characterization


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Figure 1 shows the surface of the SLM-Ti sample before and after being coated with the PCD. Through the use of SLM we can additively manufacture desired complex geometries (Fig. 1A). After we print the SLM-Ti substrates, we can then grow CVD diamond and deposit thin films to cover the entire substrate. Optimization is necessary to ensure that the coverage is uniform and well adherent (Fig. 1B). It is evident from Figs 1(C) and 1(D) that the underlying SLM-Ti samples have a large number of unmelted titanium particles in addition to the laser striations produced during the manufacturing process. The SLM-Ti has an average roughness between Ra = 7 – 12 µm, due to the partially melted particles on the surface (Fig. 1(C)).36 The PCD thin film provides a well-adhered and approximately uniform coverage of the SLM-Ti surface, as shown in Figs 1(D), 1(E) and 1(F). The PCD film coating the partially melted titanium particles has an average local roughness of ~100 nm in the area shown in Fig. 1(E). The roughness value of the PCD can be further reduced by the application of heat treatment and post-surface modification. Figure 1(F) shows the PCD grain distribution of diamond using a cross-sectional representation. By imaging the surface morphology of the SLM-Ti and PCD, we were able to determine both the crystal size of the diamond. We obtained similar roughness results for our PCD films as obtained by Tong et al.37 As reported by both Lechleitner et al.38 and Hulander et al.,39 nanometre scale roughness is beneficial for both the binding and proliferation of cells, however, due to the discrepancies in other reported studies there appears to be no agreement as to the most suitable surface roughness for orthopaedic tissue interfacing biomaterials.27, 38-39

One of the competing factors includes cell growth which varies in the amount of adherence to the material interface. The material interface is a significant factor for induced cell proliferation and attachment In addition, we can observe the PCD growth in SLM-Ti build striations. The AFM


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height image shows that the individual grains of PCD crystals are approximately 2-3 µm. Figure 1(G) shows that PCD coatings are coated in a well adherent manner all across the substrate. The coating shows no sign of delamination as the underlying SLM-Ti provides a rough texture which is able to fuse the plasma enhanced diamond on the substrate (Fig. 1(G).I, II, III). Importantly, show that PCD can be grown on SLM-Ti without an interlayer due to the surface structural integrity for potential implant coating. Figure 1(D) shows that the partially melted particles are also coated in addition to the flatter surface around it. In contrast, we show that P-Ti is a poor substrate for PCD coating as indicated by the yellow regions of delamination diamond films (Fig. 1(H). I, II, III). Non-uniform growth is clearly visible due to the lack of an interlayer and strength of Ti-C bonding. Diamond adherence is commonly induced in other circumstances using an additional exposure of amorphous carbon or a fabricated interlayer.28,

40

In another

instance, Weiser and Prawer et al. examine Fe substrates for adherence of CVD grown diamond, however the lack of structural integrity only produces average quality unaffected diamond.41


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Figure 1. Scanning Electron Micrograph (SEM), Atomic Force Micrograph (AFM) and Optical Micrograph of SLM-Ti compared to that of PCD coated SLM-Ti using CVD: (a) Selective laser melting (SLM) process to create additively manufactured SLM-Ti substrates. (b) Plasma enhanced chemical vapour deposition process for the coating of polycrystalline diamond on the top surface of substrate. (c) SEM micrograph showing as-fabricated surface morphology of SLM-Ti. Prominent ball-like partially melted titanium are visible with radii between 7 to 12 µm (d) SEM micrograph showing the PCD surface morphology which consists of striations with an average roughness of ~100 nm. The PCD has an average crystal size grain of 2 µm, determined using imageJ. (e) High resolution AFM micrograph showing the height of the PCD coated SLMTi (10 x 10 µm2). (f) Graphical representation of cross-sectional measurement along the line shown in panel. Individual PCD grains can be shown and profile measured. The blue arrows indicate the length and the red arrows indicate the width measurement. (g) Strongly adherent uniform coating of PCD on SLM-Ti on three substrates (I, II, III). (h) PCD coating on P-Ti showing regions of delamination indicated in yellow on three substrates (I, II, III).

For biomedical implants, to further assess the PCD coatings on the flat samples for delamination the ASTM standard ASTMD3359 (Adhesion Tape Test) was used. This Adhesion Tape Test provided a rigorous method to ensure that the PCD coating was robust enough to withstand handling and implantation. The adhesive strength was classified from 0 – 5, where 0 represents the least adhesive and 5 denotes the most adhesive film. We found that the PCD had some parts of the film disturbed or ruptured by the force applied by the blade before applying the pressuresensitive tape (classified as 4). The tape shows less than 5% removal from the underlying SLMTi substrate. As such, the observed regions of the cross-hatch are ruptured due to the shear force


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applied. The adhesion test is repeated three times to ensure results are satisfactory. Further details on the test procedure and guidelines are outlined by standard, ASTMD3359.42

3.2 Biocompatibility assessment

3.2.1 In vitro cell culture on PCD and SLM-Ti Figure 2 shows the response of Chinese Hamster Ovarian (CHO) cells when exposed to the SLM-Ti and PCD samples. The CHO cells adhered onto the PCD coated substrates (Figure 2B, D) to a greater extent than unmodified SLM-Ti (Fig. 2(A) and (C)). The crevices around the partially melted particles of the SLM-Ti make the cell growth discontinuous on the whole substrate. However, elongated CHO are visible on the surface of the PCD. More cell attachment on the PCD coating is noticeable due to the diamond having full coverage on the SLM-Ti surface. Although the PCD samples were shown to possess a more hydrophobic surface than the SLM-Ti, the cell density upon the PCD was significantly higher than the SLM-Ti (thirty six thousand five hundred and twenty four thousand two hundred and forty, p

Polycrystalline Diamond Coating of Additively Manufactured Titanium

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