Plasma-Activated Tropoelastin Functionalization of Zirconium for

Article pubs.acs.org/journal/abseba

Plasma-Activated Tropoelastin Functionalization of Zirconium for Improved Bone Cell Response Giselle C. Yeo,*,†,‡,§ Miguel Santos,*,§,∥ Alexey Kondyurin,§ Jana Liskova,⊥ Anthony S. Weiss,†,‡,#,|| and Marcela M. M. Bilek§,|| †

Charles Perkins Centre, ‡School of Molecular Bioscience, §School of Physics, and #Bosch Institute, The University of Sydney, Sydney, New South Wales 2006, Australia ∥ The Heart Research Institute, 7 Eliza Street, Newtown, New South Wales 2050, Australia ⊥ Institute of Physiology, Academy of Sciences of the Czech Republic, Národní 1009/3, Prague 14220, Czech Republic ABSTRACT: The mechanical strength, durability, corrosion resistance, and biocompatibility of metal alloys based on zirconium (Zr) and titanium (Ti) make them desirable materials for orthopedic implants. However, as bioinert metals, they do not actively promote bone formation and integration. Here we report a plasma coating process for improving integration of such metal implants with local bone tissue. The coating is a stable carbonbased plasma polymer layer that increased surface wettability by 28%, improved surface elasticity to the range exhibited by natural bone, and additionally covalently bound the extracellular matrix protein, tropoelastin, in an active conformation. The thus biofunctionalized material was significantly more resistant to medical-grade sterilization by steam, autoclaving or gamma-ray irradiation, retaining >60% of the adhered tropoelastin molecules and preserving full bioactivity. The interface of the coating and metal was robust so as to resist delamination during surgical insertion and in vivo deployment, and the plasma process employed was utilized to also coat the complex 3D geometries typical of orthopedic implants. Osteoblast-like osteosarcoma cells cultured on the biofunctionalized Zr surface exhibited a significant 30% increase in adhesion and up to 70% improvement in proliferation. Cells on these materials also showed significant early stage up-regulation of bone marker expression (alkaline phosphatase, 1.8 fold; osteocalcin, 1.4 fold), and sustained up-regulation of these genes (alkaline phosphatase, 1.3 fold; osteocalcin, 1.2 fold) in osteogenic conditions. In addition, alkaline phosphatase production significantly increased (2-fold) on the functionalized surfaces, whereas bone mineral deposition increased by 30% above background levels compared to bare Zr. These findings have the potential to be readily translated to the development of improved Zr and Ti-based implants for accelerated bone repair. KEYWORDS: bone, plasma-activated coating, titanium, tropoelastin, zirconium

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INTRODUCTION

Zirconium (Zr), as an alternative to Ti, possesses advantageous physical and biological characteristics for use in permanent orthopedic implants.7 Structurally, Zr displays similar or greater strength and hardness compared to Tibased materials.1b,8 In fact, the incorporation of a small percentage of Zr confers increased mechanical stability, fatigue strength, and hardness to pure Ti.2,9 Zr also has a lower Young’s modulus, which is beneficial in reducing stress shielding.10 Compared to Ti-based alloys, Zr is more electrochemically inert and resists corrosion.11 Zr naturally forms a thin film of amorphous metallic oxide, which acts as a barrier between the bulk metal and the host tissue. As a result, Zr has been used to coat Ti materials to prevent ion leaching.12

Metals are commonly used in bone prostheses and in bone fixation devices because of their mechanical stability and resistance to cyclic loading.1 Titanium (Ti) is widely utilized in orthopedic applications because of its biocompatibility, strength, and durability. However, the mechanical properties of Ti remain inadequate under inappropriate loading conditions such as in small diameter implants.2 In addition, the biocompatibility of Ti can be compromised by variable surface microtopography and hydrophobicity. 3 To improve its mechanical stability, Ti is often alloyed with other elements, particularly aluminum (Al) and vanadium(V). However, metal ions released from these additives can accumulate in the periimplant space and in local lymph nodes,4 resulting in toxicity and implant failure5 by inhibiting osteoblast differentiation, bone regeneration, and even inducing bone necrosis.6 © XXXX American Chemical Society

Received: January 28, 2016 Accepted: March 2, 2016

A

DOI: 10.1021/acsbiomaterials.6b00049 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

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ACS Biomaterials Science & Engineering

Figure 1. Biofunctionalization of metals such as zirconium or titanium for bone applications. The first step is plasma-activated coating of the metal, in which acetylene, nitrogen, and argon plasma is used to deposit a reactive carbonized layer on the material to improve surface biocompatibility and to enable covalent biomolecule linking. The second step is the immobilization of tropoelastin molecules, which retain functionality even after medical-grade sterilization, to promote bone cell activity.

generates a reservoir of radicals for the rapid covalent immobilization of compounds without the need for chemical linkers that may pose toxicity risks, present solvent disposal issues, and variable yields because of side reactions.20 Reaction with the radicals in the PAC layer allows further functionalization of the Zr surface with a protein that can mediate cell responses for bone formation and healing. One such candidate is tropoelastin, the monomer of elastin found in the extracellular matrix of elastic tissues. Tropoelastin classically has applications in soft tissue engineering, but its utility in orthopedic settings has never been explored. Because of the conformational flexibility and stability of this protein,22 it has the potential to withstand medical-grade sterilization in the same manner as its demonstrated resistance to ultraviolet light23 and proteolysis.24 Importantly, tropoelastin is known to modulate a number of processes including adhesion, spreading, proliferation, and migration in a number of cell types such as fibroblasts and endothelial cells.25 In this work, we aimed to develop a stable, bioactive Zr-based implant material that can modulate bone cell activity for improved bone healing. We modified commercially available Zr with the deposition of a plasma-activated coating followed by the immobilization of a tropoelastin monolayer (Figure 1). We demonstrated application of the biofunctionalization process to a three-dimensional spinal implant. We characterized the chemical and mechanical properties of the PAC layer on Zr, compared the bioactivity of the tropoelastin coating pre- and poststerilization, and evaluated the bone response to standard and functionalized Zr.

The high bending strength, toughness, and corrosion resistance of Zr makes it a desirable base material even for small-diameter bone implants which possess an increased risk of fracture.6 Similar to Ti, Zr is noncytotoxic13 and biocompatible.14 Zr and Zr alloys are well-tolerated in host tissues,15 and display equal or superior osteoconductivity 2 and osseointegration14,16 compared to pure Ti. Furthermore, the degree of bone-implant contact is reported to be higher with Zr than Ti,17 as supported by the consistently high shear strength and removal torque values associated with Zr implants.2 Zr has been shown to promote an osteoblastic response for bone healing.4 This initial anchorage of bone to implant surface and the maintenance of fixation over time largely determine the clinical success of orthopedic prostheses.18 Although Zr allows bone cell adhesion and growth, the metal as a largely bioinert material10 can be modified to actively promote bone formation. Zr implant surfaces can be coated with a compatible compound such as hydroxyapatite, or with a cell-interactive protein.1a,2,19 However, the adhesion of a chemical coating is often challenging due to differences in mechanical moduli, resulting in eventual failure at the bone-implant or implant-coating interface.1a Alternatively, surface functionalization with proteins may be impractical due to the requirement for chemical linkers,20 and/or loss of bioactivity after sterilization for clinical use. To address these issues, we can modify Zr by the deposition of a plasma-activated coating (PAC). This technique utilizes acetylene, nitrogen, and argon plasma to form a highly crosslinked carbonized layer on the metal surface.21 The PAC layer not only minimizes the risk of ion leaching from the underlying metal substrate, but also increases surface hydrophilicity and wettability for cell attachment.21 Additionally, the plasma ion implantation that is applied during the plasma polymerization

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MATERIALS AND METHODS

PAC Deposition. PAC was deposited onto Ti (FIRMETAL, China), Zr (FIRMETAL, China), 316L stainless steel (SS) (Brown B


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A wavelength-by-wavelength fit of the experimental SE data was performed until model convergence was achieved. Scanning Electron Microscopy. Samples were imaged with a Zeiss ULTRA PLUS scanning electron microscope at acceleration voltages ranging from 5 to 10 kV and working distances between 3 mm and 10 mm. Electron Paramagnetic Resonance (EPR). The volume density of unpaired electrons on PAC coated polystyrene foils (6 cm × 6 cm × 25 μm, Goodfellow) was measured using a Bruker EMXplus X-band spectrometer. Spectra were acquired by setting a center of field of 3510 G, modulation amplitude and frequency of 3 G and 105 Hz, sampling time of 90 ms and microwave frequency and power of 9.8 GHz and 25 mW, respectively. The spectra of both uncoated and plasma treated substrates in N2+Ar atmospheres using a bias voltage of −500 V were also acquired and subsequently subtracted from the spectrum of the PAC coated polystyrene. Such a subtraction method determines the intrinsic radical density of PAC by eliminating any background radicals and ion induced radical activation on the bare polystyrene, the latter of which is triggered by ion bombardment during the first stages of the deposition process. The resulting EPR spectra were processed with OriginPro 9.1 and compared with the measured spectrum of a standard Bruker weak pitch sample (∼1 × 1013 spins/cm) for quantitative analysis. Surface Coating with Tropoelastin. Recombinant mature-form human tropoelastin (corresponding to residues 27−724 of GenBank entry AAC98394) was produced in-lab from a large-scale Escherichia coli fermentation process.28 PAC Zr samples were aged for at least 7 days after PAC deposition to allow stabilization of surface properties prior to protein immobilization. Zr and PAC Zr were incubated in 20 μg/mL tropoelastin dissolved in PBS (10 mM phosphate, 150 mM NaCl) at pH 7.4 or 11.5 at 4 °C overnight. Samples were then washed with PBS at pH 7.4 to remove unbound protein. Enzyme-Linked Immunosorbent Assay. The binding of tropoelastin to Zr surfaces was determined with an enzyme-linked immunosorbent assay (ELISA). Where indicated, Zr and PAC Zr samples with and without tropoelastin were washed in 5% (w/v) sodium dodecyl sulfate at 80 °C for 10 min then rinsed with PBS. Samples were then blocked with 3% (w/v) bovine serum albumin (BSA) (Sigma-Aldrich) for 1 h at room temperature and washed with PBS. Bound tropoelastin was detected with 1:2000 dilution of mouse antielastin BA4 antibody (Sigma-Aldrich) for 1 h, 1:5000 dilution of goat antimouse IgG conjugated with horseradish peroxidase (SigmaAldrich) for 1 h, and visualized with substrate solution (1.04 mg/mL 2,2′-azino-bis(3-ethylbenzothiazoline-6-sulfonic acid)diammonium salt, 0.05% (v/v) H2O2, 10 mM CH3COONa, 5 mM Na2HPO4) at 37 °C for 1 h. Sample absorbances at 405 nm were read with a plate reader. Steam Autoclaving. Zr samples, with and without PAC and/or tropoelastin, were placed in 10 mL PBS and steam autoclaved at 120 °C and at 100 kPa above atmospheric pressure for 20 min. After autoclaving, samples were washed three times with PBS and stored in PBS at room temperature until use. Gamma-Ray Irradiation. Zr samples, with and without PAC and/ or tropoelastin, were air-dried overnight in polystyrene well plates and posted to the Australian Nuclear Science and Technology Organisation (ANSTO) for cobalt-60 gamma-ray irradiation at 25 or 40 kGy. Samples were irradiated within a 3% deviation of the required dose. Nonirradiated control samples were prepared as part of the same batch as the irradiated samples, but were kept on the lab bench until use. Cell Culture. MG63 human osteoblast-like osteosarcoma cells (obtained from William Lu, University of Sydney) were grown in Dulbecco’s Modified Eagle Medium (DMEM) (Sigma-Aldrich) with 10% (v/v) fetal bovine serum (FBS) (Life Technologies). SAOS-2 human osteoblast-like osteosarcoma cells (Sigma-Aldrich) were cultured in McCoy’s 5A Modified Medium (Sigma-Aldrich) supplemented with 10% (v/v) FBS and 2 mM L-glutamine (Life Technologies). Cell Attachment. Zr samples, with and without PAC and/or tropoelastin, were blocked for 1 h at room temperature with 10 mg/ mL BSA that has been heat-denatured at 80 °C for 10 min and then

Metals, USA) and silicon (Si) (Addison, USA) substrates by means of plasma enhanced chemical vapor deposition with enhanced ion bombardment. Plasma discharges were generated and sustained by coupling radiofrequency power supplied at 13.56 MHz and samples were electrically biased using a DC high voltage pulse generator to provide a bias voltage of −500 V. Pulse frequency and pulse duration were set at 3 kHz and 20 μs, respectively. Reactive gaseous mixtures of argon, nitrogen and acetylene were used to sustain the plasma discharge and the flow rates of the gases were individually controlled by mass flow controllers and set at 5 sccm, 10 and 20 sccm, respectively. The working pressure prior to plasma generation was maintained constant at 110 mTorr. All samples were cleaned prior to PAC deposition with toluene, acetone, and ethanol in an ultrasonic bath (samples were submerged for 10 min in each organic solvent), followed by 15 min of argon plasma cleaning. X-ray Photoelectron Spectroscopy. Chemical characterization of the coatings was studied by means of X-ray photoelectron spectroscopy (XPS) using a SPECS-XPS (Germany) equipped with a hemispherical analyzer and an Al Kα monochromatic X-ray source. The system was operated at a constant pressure of ∼1 × 10−9 Torr, with a takeoff angle of 90° at a power of 200 W. Survey spectra of the samples were scanned in the energy range of 50 eV − 1400 eV using an energy step of 1 eV and a pass energy of 30 eV. High resolution scans for C 1s, N 1s and O 1s were recorded with an energy step of 0.03 eV and a pass energy of 23 eV. Peak analysis was performed using CasaXPS software. The atomic fraction in the coatings was determined by calculating the integrated areas of the C 1s, N 1s, and O 1s peaks. Peak fitting was carried out by adopting a Shirley background and convoluted Lorentzian (30%) − Gaussian (70%) line shapes with identical full-width half-maximum for each component. Charge compensation was applied in all spectra by assigning the C−C/C− H component in the C 1s peak at a binding energy of ∼285 eV. Fourier Transform Spectroscopy in Attenuated Total Reflectance Mode. Infrared spectra were recorded by means of Fourier transform spectroscopy in attenuated total reflectance mode (FTIR-ATR) using a Bruker Hyperion FTIR microscope integrated in Vertex v70 system and equipped with a 20x ATR objective with a germanium crystal. Each spectrum is the result of averaging a total of 500 scans at a spectral resolution of 4 cm−1 in the wavenumber range of 3750−750 cm−1. Spectral subtraction and baselines were applied to eliminate the background signal from the underlying substrate as previously described.26 Water Contact Angle Measurements. Wettability measurements were performed using a Kruss DS10 analyzer equipped with a CCD camera and water contact angles were estimated employing the sessile-drop method. Reported values for each sample are averages over 10 measurements and were obtained using 1 μL drops of deionized water. Nanoindentation. Nanoindentation was performed using a G200 Nano Indenter (Agilent) with a displacement and loading resolution of 0.01 nm and 50 nN respectively. The nano indenter was equipped with calibrated Berkovich diamond tip with a Poisson ratio of 0.18. Thermal drift corrections were applied and the maximum allowable drift rate was 0.05 nm/s. The loading rate was set at 0.02 nm/s and applied loads were chosen to be between 0.03 and 19 mN. A total of 120 indentations were performed on each sample and the results presented here are averages of 12 indentations made in different regions of the sample. Contact stiffness, residual area, indentation depth and reduced Young’s Modulus were determined by analyzing the resulting unloading curves using the Oliver−Pharr method.27 Elastic recovery was estimated by integration of the loading and unloading curves to allow calculation of the ratio of the irreversible work to the total work during the indentation process. Spectroscopic Ellipsometry. PAC thickness was determined via spectroscopic ellipsometry (SE) using a Woollam M2000 V spectrometer and the WVASE32 software. SE data acquisition was performed at three angles of incidence (65, 70, and 75°) and covering a spectral region ranging from 200 nm up to 1000 nm. A Cauchy layer was implemented in the model to represent the coating layer on top of Si substrates of which optical parameters were previously determined. C


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Table 1. Sequence of RT-PCR Primers65 Used in the Analysis of Bone Marker Expression of Cells Cultured on Zr Samples marker

forward primer (5′-3′)

reverse primer (5′-3′)

ALP COL1 OCN GAPDH

GACCCTTGACCCCACAAT CAGCCGCTTCACCTACAGC GAAGCCCAGCGGTGCA TGCACCACCAACTGCTTAGC

GCTGCTACTGCATGTCCCCT TTTTGTATTCAATCACTGTCTTGCC CACTACCTCGCTGCCCTCC GGCATGGACTGTGGTCATGAG

Statistical Analyses. All data were reported as mean ± standard error (n = 3 unless otherwise indicated). Two-way analysis of variance was performed using GraphPad Prism (GraphPad Software). Statistical significance was set at p < 0.05 or higher, and indicated in the figures as ‘ns (not significant)’ (p ≥ 0.05), * (p < 0.05), ** (p < 0.01) or *** (p < 0.001). Where relevant, the p-values of statistical comparisons are included in figures.

cooled. MG63 cells were trypsinised with 0.25% trypsin-EDTA (Sigma-Aldrich) at 37 °C for 3 min, centrifuged at 800 g for 5 min and resuspended in serum-free DMEM. Samples were washed with PBS and seeded with 1.5 × 105 cells/cm2 at 37 °C in 5% CO2 for 1 h. Nonadherent cells were removed with PBS, while adherent cells were fixed with 3% (w/v) formaldehyde in PBS for 20 min. Samples were washed with PBS and the fixed cells were stained with 0.1% (w/v) crystal violet in 0.2 M MES buffer, pH 5.0 at room temperature for 1 h. Samples were rinsed thoroughly in reverse osmosis water prior to solubilizing the stain with 10% (w/v) acetic acid. Absorbance was measured at 570 nm using a plate reader. Absorbance readings from standards containing 20, 40, 60, 80, and 100% of the seeding density were fitted to a linear regression, which was used to convert sample absorbances into percentage of cell attachment. Cell Proliferation. Zr samples, with and without PAC and/or tropoelastin, were sterilized by steam autoclaving or gamma irradiation at 40 kGy. The surfaces were BSA-blocked similarly to a cell attachment assay and seeded with MG63 cells at a density of 8.0 × 103 cells/cm2 in normal media. The culture media was changed every 2 days. At 1, 3, 5, and 7 days postseeding, samples were washed with PBS, and the bound cells were fixed, stained with crystal violet, and quantified as described in the previous section. Real-Time Reverse Transcriptase Polymerase Chain Reaction. SAOS-2 cells were seeded at a density of 2.5 × 104 cells/cm2 in normal media on sterilized Zr samples with and without PAC and/or tropoelastin and cultured until confluence. At this point, the samples were either kept in normal media for 7 days, or transferred to osteogenic media [Alpha Modified Eagle Medium (Lonza) with 20% (v/v) FBS, 2 mM L-glutamine, 1% (v/v) penicillin/streptomycin (Life Technologies), 1 mM sodium glycerophosphate (Sigma-Aldrich), 50 μg/mL L-ascorbate (Sigma-Aldrich), and 1 × 10−8 M dexamethasone (Sigma-Aldrich)] for 14 days. Media was refreshed every 2 days. Total RNA was extracted from the cells at various time points with the High Pure RNA Isolation Kit (Roche). RNA was converted to cDNA using the Transcriptor First Strand cDNA Synthesis Kit (Roche). Expression of bone markers including alkaline phosphatase (ALP), collagen type I (COL1), and osteocalcin (OCN) were quantified with real-time reverse transcriptase polymerase chain reaction (RT-PCR) using specific primers (Table 1) in a LightCycler 480 System (Roche). Bone marker expression levels were referenced against GAPDH expression and normalized to those of cells cultured to confluence on standard Zr. Alkaline Phosphatase Quantification. SAOS-2 cells were seeded at a density of 1.5 × 104 cells/cm2 and cultured in osteogenic media on Zr with and without PAC and/or tropoelastin for 7 days, with media changes every 2 days. The amount of ALP enzyme in condition media was quantified using the Alkaline Phosphatase Assay Kit (Abnova) following the manufacturer’s instructions. Cells on the Zr surfaces were trypsinized and counted using a hemocytometer. Enzyme levels were normalized according to cell numbers in each sample. Bone Mineralization. MG63 cells were seeded at a density of 1.5 × 104 cells/cm2 and cultured in normal media on sterilized Zr samples with and without PAC and/or tropoelastin until confluence. Samples were then continually incubated in normal media or induced with osteogenic media. At 0, 10, 20, and 30 days postconfluence, samples were assayed for mineralization according to standard protocol.29 Briefly, cells were fixed in 10% (v/v) formaldehyde and stained for calcium deposits with 40 mM Alizarin Red S. Cells were resuspended in 10% (v/v) acetic acid, overlaid with mineral oil, and heated at 85 °C for 10 min. After centrifugation, the supernatant was alkali-neutralized and absorbance values were obtained at 405 nm.

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RESULTS PAC Deposition on Zr. XPS measurements on PAC surfaces revealed the presence of carbon, nitrogen and oxygen in relative atomic concentrations of 69, 24, and 7%, respectively. Figure 2A details the high resolution XPS C 1s

Figure 2. (A) C 1s and N 1s peaks of PAC by XPS. The carbon peak was deconvoluted into four contributions located at 284.9, 286.3, 287.8, and 289.3 eV. The nitrogen peak was also deconvoluted into four contributions at 398.0, 399.2, 400.4, and 402 eV. (B) FTIR spectrum of PAC measured in total reflectance mode in the wavenumber range between 3750 and 750 cm−1.

and N 1s peaks with their corresponding deconvolutions. Four components contribute to the C 1s peak with positions 284.94, 286.25, 287.50, and 288.78 eV and relative fractions of 42.9, 35.6, 17.4, and 4.1%, respectively. The first peak at the lowest binding energy corresponds to pure carbon configurations in an amorphous network as well as C−H bonds, whereas the fourth peak at the highest binding energy is attributed to different COO compounds.30 The interpretation and assignment of the CN components in the XPS C 1s peak in amorphous CN:H materials greatly depends on the local chemical environments in which the various CN configurations occur.31 The peaks at D


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known as “hydrophobic recovery”. Indeed, our PAC coatings modify the thermodynamical equilibrium of the Zr surface. Here, hydrophobic recovery in ambient air is driven by surface mechanisms that occur to minimize the PAC surface energy, while a new thermodynamical equilibrium is achieved over time after PAC deposition. Such surface mechanisms may include adsorption and chemical reactions, reorientation of surface polar groups, oxidation and changes to the surface electrical charge. The extent of the hydrophobic recovery and its kinetics greatly depends on the surface treatment and the storage conditions. Mechanical Properties of Zr and Ti. The Young’s modulus as a function of the indentation depth was measured by nanoindentation in both bare and PAC Zr and Ti. The PAC/metal composites were able to resist plastic deformation and displayed superior elasticity when compared to uncoated Zr and Ti (Figure 3A). The elastic recovery for both PAC/

∼286 eV and ∼287 eV have been assigned in the literature either to fixed sp2 or sp3 CN hybridizations32 or according to the number of nitrogen neighbors for each carbon atoms.33 Because of the wide range of local chemical environments typically occurring in plasma polymers, the assignment of the two peaks located at 286.25 and 287.50 eV cannot be fixed to specific CN hybridizations. Therefore, we will compare their relative fractions in the C 1s peak with the results obtained from the deconvolution of the N 1s peak. Although the assignment of different CN components in the XPS N 1s peak is also somewhat controversial, there seems to be general agreement with respect to their relative position in the spectrum. The XPS N 1s peak was deconvoluted into 4 distinct components. The first peak is located at 398.30 eV (17.9%) and has been assigned in the literature either to diagonal nitrogen atoms bonded to two carbon atoms, such as those found in pyridinic-N form compounds;34 or to sp3 nitrogen atoms bonded to sp3-hybridized carbon atoms.35 The second component with a peak centered at 399.37 eV (44.9%) is related to both nitrile and amine groups.36 The third peak centered at 400.39 eV (33.8%) is generally attributed to nitrogen trigonally bonded to three sp2 carbon atoms, like in graphitic-N structures;35 or nitrogen atoms bonded to two sp2 and one sp3 carbon atoms;37 or to sp2 nitrogen atoms bonded to sp2-hybridized carbon atoms.34 Finally, the fourth peak at 401.80 eV (3.4%) is commonly attributed to nitric oxide compounds.36 Interestingly, the ratio between the second and third components of the C 1s peak yields 2.045 which appears to be correlated with the ratio between the relative contributions of the third and first components of the N 1s peak, 1.888. This suggests that their assignments may contain equivalent CN hybridizations. There is also a contribution from the oxygen species C−O and CO, respectively, in the C 1s second and third component peaks. FTIR-ATR is consistent with the chemical composition of PAC obtained by XPS. Figure 2B shows a typical infrared spectrum of PAC deposited on metallic substrates in the 3750− 750 cm−1 wavenumber range. The spectrum features two broad absorption bands located in the 3600−2800 and 1750−1050 cm−1 ranges. The first band is attributed to O−H (3450−3400 cm−1), NH (3320 and 3200 cm−1) and CH (2880, 2935, and 2970 cm−1) stretching vibrations. The second and broader band is assigned to CO (1715−1680 cm−1), CC (1640− 1680 cm−1), CN (1640−1690 cm−1) stretching vibrations, as well as to N−H (1550−1500 cm−1) and C−H (∼1450 and 1380 cm−1) bending vibrations, stretching vibrations in aromatic rings (1500−1450 cm−1), and N−O (1340 cm−1), C−N (1250 cm−1), and C−O (1200−1000 cm−1) stretching vibrations. The peak located at ∼2200 cm−1 is attributed to stretching vibrations of nitrile groups, and the peak at ∼800 cm−1 to out-of-plane vibrations in unsaturated carbon assemblies. Wettability measurements revealed that PAC improves the hydrophilic character of bare Zr substrates. The water contact angle of PAC coated Zr was 42.0° ± 1.1° 30 min after deposition, increasing and stabilizing at 61.3° ± 2.3° by day 10. Conversely, the water contact angle measured on bare Zr substrates was 84.7° ± 5.8°. It was reduced to 66.6° ± 1.5° after employing a cleaning protocol with organic solvents but typically the contact angle reverts to precleaning values within 24 h of being cleaned.38 The increase in the water contact angle with aging time is typically observed in surfaces modified either by plasma coatings or cleaning processes and is commonly

Figure 3. (A) Young’s modulus as a function of the indentation depth measured by nanoindentation on both bare and PAC coated Zr and Ti substrates. The PAC/metal interface depth was estimated using ellipsometric spectroscopy by measuring the thickness of PAC deposited onto Si substrates. The elastic modulus of bone52 is indicated. (B) SEM images of PAC deposited on Zr, Ti and SS substrates after sample deformation. Samples were bent over a sharp edge with a ∼750 μm radius of curvature with the film on the outside of the bend.35 Although PAC strongly adhered to Zr and Ti substrates, coating delamination was observed on SS substrates. The arrows indicate crack patterns propagating in small patches along the deformed region and the dashed line indicates the position of the bend.

metal composites was estimated by calculating the amount of dissipated energy during the indentation process and was found to be 59.6% and 62.6% for PAC Zr and PAC Ti, respectively. At lower indentation depths (25 and 27 nm for PAC Zr and PAC Ti, respectively), a loading of 0.03 mN returned a Young’s modulus of 30 GPa, corresponding to an 80% reduction of the ∼140 GPa measured on the bare substrates. The plateau observed for indentation depths above 100 nm is consistent with the thickness of PAC deposited onto Si substrates and estimated with spectroscopic ellipsometry. As expected, the Young’s moduli of the PAC/metal composites recovered the bare metal values at higher indentation depths. Most importantly, we show here that the Young’s moduli of the surface layers of PAC/metal composites are compatible with the reported range of 10−40 GPa for natural bone.39 E


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Figure 4. (A) Typical electron paramagnetic resonance spectrum of PAC. (B) Antibody detection of tropoelastin retained on Zr and PAC Zr before and after SDS washing. Tropoelastin was associated with the Zr surfaces at either pH 7.4 or 11.5. (C) Superimposed FTIR ATR spectra of a functionalized titanium alloy intervertebral cage: a, initial raw spectrum of PAC and tropoelastin; b, subtracted PAC spectrum with the y-scale multiplied 5 times and offset, showing a broad band in the 1800−1400 cm−1 region; c, subtracted spectrum of tropoelastin layer with the y-scale multiplied 12 times and offset, showing the amide I and amide II peaks with characteristic positions, amplitudes and widths. (D) Amide I vibration line absorbance at 1650 cm−1 from tropoelastin immobilized on a titanium alloy intervertebral cage modified with PAC. The inset shows a photograph of the device with the arrows indicating the locations from which micro-FTIR ATR spectra were recorded. (E) Distribution map of covalently bound tropoelastin within a 90 × 90 μm2 area on the bottom of the functionalized titanium cage, calculated by the relative Amide I vibration line absorbance of the protein-coated samples before and after SDS detergent washing. On average, 71% of the total attached tropoelastin molecules are covalently linked to the material surface. (F) MG63 cell attachment to bare, BSA-blocked, or tropoelastin-coated and BSA-blocked Zr and PAC Zr surfaces. Tropoelastin binding to Zr was performed at either pH 7.4 or 11.5.

features a broad resonance peak centered at 3513 G, corresponding to a g-factor of 2.003. This resonance peak comprises different spin components attributed to the unpaired elections associated with radicals within the CN:H coating. Double integration of the EPR spectrum and further quantitative analysis with a standard sample (∼1 × 1013 spins/cm3) showed a radical density of 3 × 1017/cm3 15 days after PAC deposition. To determine if biomolecules can be covalently bound on the plasma-modified Zr surfaces, we washed tropoelastin-coated Zr and PAC Zr with SDS and assayed for residual protein with an antielastin antibody (Figure 4B). Prior to SDS washing, similar levels of tropoelastin were detected on the Zr and PAC Zr surfaces. However, on bare Zr, the amount of detected tropoelastin dropped to background levels after SDS washing, indicating that the protein molecules were simply adsorbed on the surface. In contrast, up to 79% of tropoelastin molecules were retained on PAC Zr after SDS washing, pointing to the covalent immobilization of the majority of the protein on the plasma-modified substrate. In addition, the pH at which tropoelastin associates with the surface did not significantly affect the extent of covalent protein retention. Comparable amounts of tropoelastin were detected on PAC Zr after SDS washing, regardless of whether a buffer pH of 7.4 or 11.5 was used during protein coating. Application of Functionalized PAC to Complex 3D Geometries of Orthopedic Implants. To assess the applicability of our process to three-dimensional complex

The elevated elastic moduli of both uncoated metals measured at low indentation depths indicates the presence of native oxides (ZrO2 and TiO2) on their surfaces. The presence of such oxides on the uncoated metals was confirmed by XPS. Ion bombardment in the initial phases of coating deposition appears to have substantially removed these oxide layers from the coated materials. The adhesion of PAC to the underlying material was tested by deforming the coated substrates and folding the samples inward and outward with small radii of curvature using a sharp edge.35 As a control, PAC was also deposited onto 316 L stainless steel substrates. While PAC strongly adhered to all metallic surfaces following substrate deformation, adhesion strength was markedly higher for PAC deposited on Zr and Ti (Figure 3B), suggestive of a stronger interface between these metals and the PAC coating. Atomic interfacial mixing (AIM)40 is expected to occur at the interface of the coating and the metal due to the ion bombardment induced by the applied bias used during deposition. The mechanical strength is likely to be significantly enhanced for the PAC-Zr/Ti interfaces by the formation of covalently bonded metal carbide structures, and extended compositional grading required to optimize adhesion to stainless steel41 does not appear to be required for the carbide forming metals. Covalent Protein Immobilization on PAC Zr. PAC contains a reservoir of radicals that migrate to the surface over time to allow covalent linkage of molecules. The electron paramagnetic resonance (EPR) spectrum of PAC (Figure 4A) F


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Figure 5. Antibody detection of tropoelastin on Zr samples before and after sterilization. Antibodies were targeted against (A,C) multiple sites within tropoelastin, or (B, D) its C-terminal region. Zr samples were sterilized by (A, B) steam autoclaving or (C, D) gamma ray irradiation at different doses.

gamma-ray irradiation for up to 40 kGy. The amount of substrate-bound tropoelastin after sterilization was quantified with an antielastin antibody that recognizes multiple sites within the tropoelastin molecule (Figure 5A and 5C). After autoclaving, intact protein levels on Zr were reduced to near background levels, whereas 63% of tropoelastin molecules remained detectable on the PAC modified surface (Figure 5A). Similarly, while minimal tropoelastin was measured on Zr postirradiation, a significant majority of the tropoelastin coating remained intact on PAC Zr. Protein preservation on PAC Zr was estimated at 80% after irradiation at 25 kGy, and 66% after irradiation at 40 kGy. The surfaces were also probed for the presence of a Cterminal cell-binding region within tropoelastin using a specific antibody targeted against this sequence (Figure 5B, D). Consistent with the relative overall amounts of detected tropoelastin on Zr and PAC Zr, the PAC modified surface showed significantly higher levels of the cell-interactive region poststerilization. Antibody recognition of this region was 9-fold higher on PAC Zr than on Zr after autoclaving, and 2-fold higher after gamma irradiation. Crucially, the tropoelastin-functionalized Zr samples remained cell-adhesive after sterilization. While MG63 cell attachment decreased by 59% on tropoelastin-coated Zr after autoclaving, it was fully retained on tropoelastin-coated PAC Zr (Figure 6A). In the same manner, tropoelastin adsorbed to Zr displayed 39 and 19% loss of cell binding activity after gamma irradiation at 25 and 40 kGy, respectively. In contrast, bone cell adhesion to tropoelastin-coated PAC Zr was completely preserved up to the highest dose of radiation (Figure 6B). Interestingly, gamma irradiation even appeared to improve cell attachment to Zr, which was particularly evident on the PAC modified surfaces even in the presence of BSA blocking. This, combined with the increased protein retention observed on PAC Zr, accounted for the consistently markedly higher bone

shaped medical devices, a threaded Ti alloy intervertebral cage was functionalized with PAC and tropoelastin. After drying the implant, FTIR ATR spectra were taken from selected locations on the implant. The results showed a reasonably uniform distribution of the protein across the implant as indicated by the absorbance of the Amide I line (Figure 4C, D). For comparison, protein coverage of a flat (2D) sample was also shown. Given that the flat surface has been shown to be covered by a dense protein monolayer,41a we infer that the 3D device is also completely coated with a tropoelastin monolayer. On average, 71% of the adhered tropoelastin molecules were retained on the implant after SDS washing (Figure 4E), comparable to the extent of covalent protein binding described for the PAC treated 2D surfaces. Increased Bone Cell Adhesion to Functionalized PAC Zr. The ability of Zr to support bone cell adhesion was determined by seeding osteoblast-like osteosarcoma cells on the substrates (Figure 4F). Bare Zr, as a biocompatible material, supported a high level of cell attachment, with 90 ± 2% of cells adhering to the surface after an hour. Modification of Zr by PAC alone did not significantly improve bone cell attachment. Coating of Zr and PAC Zr with a noncell-adhesive reagent such as denatured BSA abolished the cell binding properties of either surface. However, Zr and PAC Zr that had been tropoelastin coated prior to BSA blocking exhibited significantly increased bone cell attachment compared to bare (30% and 28% increase over Zr and PAC Zr, respectively) or BSA-blocked surfaces, demonstrating the benefit of the tropoelastin layer. Cell binding did not significantly differ between surfaces on which tropoelastin was physisorbed (i.e., Zr) or covalently bound (i.e., PAC Zr), or between surfaces on which the protein was associated at pH 7.4 or 11.5. Sustained Poststerilization Bioactivity of Functionalized PAC Zr. Tropoelastin-coated Zr and PAC Zr were sterilized to clinical standards either by steam autoclaving or by G


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Figure 6. MG63 cell adhesion to Zr samples before and after sterilization by (A) steam autoclaving or (B) gamma ray irradiation at different doses. Where present, significance markers directly on top of columns indicate comparisons against the corresponding nonirradiated tropoelastin-coated PAC Zr or Zr sample.

cell adhesion to tropoelastin-coated PAC Zr over tropoelastincoated Zr after sterilization. These results confirmed the ability of the functionalized Zr material to withstand medical-grade sterilization. Improved Bone Cell Proliferation on Functionalized Zr. The ability of medically sterilized, tropoelastin-coated, PAC modified Zr to support MG63 proliferation was determined (Figure 7). At 3, 5, and 7 days postseeding, significantly higher cell numbers were observed on autoclaved tropoelastinfunctionalized Zr surfaces over BSA-blocked controls (Figure 7A). Over the course of 7 days, tropoelastin-coated Zr and PAC Zr promoted a 55 and 29% increase in cell abundance, respectively, compared to samples without tropoelastin. Likewise, protein-functionalized Zr films that had been subjected to gamma irradiation at 40 kGy significantly induced proliferation of MG63 cells (Figure 7B). At every time point after the first day postseeding, consistently higher cell densities were observed on the tropoelastin-coated Zr (74% increase) and PAC Zr (56% increase) surfaces compared to BSA-blocked samples. There was no significant difference in bone cell proliferation on Zr surfaces where tropoelastin was adsorbed or covalently immobilized. However, an adsorbed protein layer would not be viable for in vivo use because of the protein exchange that occurs when adsorbed surface protein layers interact with a protein-rich environment such as blood.42 Tropoelastin-functionalized Zr surfaces proved advantageous for MG63 proliferation even when compared to non-BSAblocked Zr and PAC Zr under ideal culture conditions in which the bare surfaces were not compromised by nonspecific protein adhesion as might be expected in vivo. Cell numbers on tropoelastin-coated Zr and PAC Zr were 20% and 14% higher, respectively, relative to bare Zr and PAC Zr 7 days postseeding (Figure 7C). PAC deposition on Zr, by itself, did not stimulate

Figure 7. MG63 cell proliferation on Zr samples sterilized by (A) autoclaving or (B−C) gamma irradiation at 40 kGy. Substrate-bound cells were quantified at 1, 3, 5, and 7 days postseeding. Panels B and C correspond to one experiment in which bare, BSA-blocked, and tropoelastin-coated samples were analyzed together. Pairwise comparisons between bare and tropoelastin-coated samples, and between BSA-blocked and tropoelastin-coated samples, were separated into the two panels for clarity.

cell proliferation, as indicated by consistently comparable cell abundance on Zr and PAC Zr. These results validate the retained bioactivity of tropoelastin poststerilization, and reinforce the benefits of Zr functionalization with this protein. Bone Marker Expression on Functionalized Zr. To determine the effect of Zr functionalization on the cellular expression of bone markers, ALP, COL1, and OCN transcript levels of SAOS-2 cells grown on modified Zr surfaces in normal or osteogenic media were compared to those of cells on standard Zr at confluence (Figure 8A−D). In normal media, newly confluent cells on PAC modified Zr showed a slight increase in ALP expression (1.3 ± 0.1 fold), which persisted for 7 days postconfluence. Newly confluent cells on tropoelastincoated PAC Zr exhibited more up-regulated ALP (1.8 ± 0.2 fold) and OCN (1.4 ± 0.1 fold) expression levels, which were also significantly higher compared to cells on PAC Zr and tropoelastin-coated Zr. However, these transcriptional modulatory effects were not sustained long-term in normal media. Zr modification by PAC deposition and/or tropoelastin coating was associated with increased ALP expression (1.3 ± 0.1 fold) of bone cells for up to 10 days in osteogenic media. Up-regulated osteocalcin expression (1.2 ± 0.2 fold) was also observed with cells on tropoelastin-coated PAC Zr after 14 days in osteogenic media, further illustrating the modulatory activity H


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Figure 8. Expression of bone markers by SAOS-2 cells grown on Zr surfaces. Transcript levels of ALP, COL1 and OCN normalized against GAPDH were measured (A) at cell confluence in normal media, (B) at 7 days postconfluence in normal media, (C) at 10 days postconfluence in osteogenic media, and (D) at 14 days postconfluence in osteogenic media. Expression levels indicated are relative to those of newly confluent cells on standard Zr. Where present, significance markers directly on top of columns indicate comparisons against standard Zr. (E) Extracellular ALP secretion by SAOS-2 cells cultured on Zr surfaces. Enzyme levels were normalized to cell numbers. ALP production by human dermal fibroblast (HDF) controls is indicated.

of functionalized Zr materials. Additionally, this up-regulation of ALP and OCN was also observed when expression levels on functionalized Zr were adjusted to the base expression levels on standard Zr at the corresponding time and media condition, confirming the benefit of PAC and tropoelastin surface modification. On all Zr surfaces, the SAOS-2 cells exhibited a time-dependent increase in ALP, COL1, and OCN expression consistent with their osteoblast-like phenotype. The stimulation of bone marker expression by modified Zr was confirmed via the detection of secreted ALP by SAOS-2 cells cultured on these surfaces (Figure 8E). After 7 days in osteogenic media, cells on PAC-modified Zr substrates produced significantly more ALP compared to those on standard Zr (1.5-fold increase), whereas tropoelastin-coated substrates showed a 2-fold increase in ALP compared to unmodified Zr. Human dermal fibroblasts grown on tropoelastin-coated PAC Zr as a negative control did not produce detectable levels of the ALP enzyme. Bone Mineralization on Functionalized Zr. MG63 cells cultured on Zr surfaces in normal and osteogenic media were examined for bone nodule formation via staining with Alizarin Red S (Figure 9). Cells on tropoelastin-coated PAC Zr displayed maximum calcium deposition by 10 days postconfluence, which was significantly increased compared to that by cells on all other Zr surfaces. This early stage formation of mineralized deposits on functionalized PAC Zr was detected in

both normal and osteogenic culture conditions. The higher bone mineralization levels on functionalized PAC Zr were sustained until 20 days postconfluence in normal media, and until 30 days postconfluence in osteogenic media. In osteogenic conditions, bone nodule formation on tropoelastin-coated PAC Zr was significantly increased by 21 ± 2%, and by 30 ± 4% above background levels over bare Zr.

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DISCUSSION Zirconium, like titanium, is used in permanent orthopedic implants because of its mechanical strength, chemical stability and biocompatibility.1b Nevertheless, whereas Zr materials are widely shown to sustain osteoblast function43 and support bone apposition without the formation of intervening connective tissue,44 they do not actively enhance bone healing because of the bioinert nature of the Zr metal. To confer bioactivity to Zr implants while retaining the desirable physical properties of the bulk metal, we first coated the Zr surface with an organic material produced from plasma-induced deposition of acetylene, nitrogen and argon vapors. This PAC layer is composed of carbon, nitrogen, hydrogen, and oxygen similar to coatings previously deposited on stainless steel substrates and associated with favorable biological responses in the context of cardiovascular applications.45 PAC deposition was shown to improve surface hydrophilicity of metal substrates, similar in effect to the sand-blasting and acid-etching processes I


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expected to occur at SS substrate interfaces. The formation of extensive carbon−zirconium bonding is likely to be facilitated by the influx of high energy carbon ions into the surface of the metallic substrate, provided by the application of negative bias voltage pulses in plasma immersion ion implantation (PIII).21 Another important consequence of PAC deposition on Zr is the significant increase in the surface elasticity of the material to levels comparable to natural bone.39 The Young’s modulus of PAC Zr at ∼30 GPa is even lower than the minimum value (∼35 GPa) reported for modified Ti-based materials.43 The stiffness of PAC is also lower than that of plasma polymerized amorphous hydrogenated carbon coatings prepared in methane and argon mixtures,49 and its Young’s modulus is lower than typical diamond-like carbon coatings.50 The incorporation of carbon−nitrogen hybridizations in the PAC bonding configuration, confirmed by XPS and FTIR measurements, render the implant surface highly elastic and able to resist plastic deformation. Carbon nitride films are relatively hard and elastic when compared to the softer and stiffer hydrogenated carbon materials deposited by plasma discharges in hydrocarbon gases alone. The mechanical properties of orthopedic implants dictate the efficacy of bone contact.51 A clear discrepancy in stiffness between the implant and bone, such as that observed with traditional Ti or Zr, results in nonhomogeneous stress transfer and inevitably leads to bone resorption, prosthesis loosening, and bone refracture.52 In contrast, the equivalent stiffness of the PAC Zr surfaces and bone may help to reduce the stress shielding effect and potentially improve orthopedic implant viability. In addition to providing physicochemical and mechanical benefits to metals, the PAC layer serves as a reservoir of free radicals for rapid and direct covalent attachment of bioactive molecules to the implant surface.20 Extracellular matrix protein derivatives such as tropoelastin fragments53 and a fibronectinosteocalcin fusion protein26b have previously been immobilized on PAC. The majority of the tropoelastin molecules adhered to PAC Zr films and PAC Ti implants exhibited resistance to stringent SDS washing at high temperatures, indicating covalent immobilization on the surface as also demonstrated in previous work.41b,45,54,41a,55 The covalent coupling between PAC and tropoelastin is due to unpaired electrons of radicals embedded within the CN:H coating. These unpaired electrons diffuse to the surface over time through a series of radical reactions in the bulk. Unpaired electrons have been detected in PAC and in similar coatings for up to 12 months using electron spin resonance.41a,55 In previous reports,41a,55 we have shown increasing tropoelastin retention after SDS washing with increasing ion energies during film growth and for thicker coatings, i.e., conditions that impart higher radical densities within the coating bulk. Covalent binding is necessary for sufficient retention of the tropoelastin layer during sterilization and to eliminate protein loss via dynamic protein exchange that rapidly displaces physisorbed molecules, particularly in the in vivo setting of implant materials. Aside from the stability of the protein coating, equally important to effective bare metal functionalization is the functional presentation of the tropoelastin molecules. Tropoelastin orients on PAC surfaces such that covalent linkages primarily occur within its central region, leaving the C-terminal cell-binding region accessible for functional interactions.24 The activity of substrate-bound tropoelastin was validated by the high attachment of bone cells to tropoelastin-coated Zr that had also been blocked with noncell-adhesive denatured BSA. This

Figure 9. Bone nodule formation by MG63 cells grown on Zr surfaces in (A) normal or (B) osteogenic media (n = 6; samples are independent and normally distributed as validated by the Kolmogorov−Smirnov test). Calcium deposition was quantified using Alizarin Red S staining of cells at 1, 10, 20, and 30 days postconfluence. Where present, significance markers directly on top of columns refer to comparisons against standard Zr at the corresponding time point. Dotted lines indicate the background absorbance of control samples without cells.

performed on Zr,46 which are thought to facilitate early stage osseointegration7 through the adsorption of proteins and inorganic compounds for rapid bone healing.47 Surface roughness is also correlated with improved osteoblast viability, proliferation, differentiation and mineralization.43 Because the PAC coating conforms closely to the substrate, further enhancements in bone healing might be achievable by coating PAC onto preroughened substrates. PAC has the added benefit of uniform surface chemistry and topography,45 in contrast to bare Zr surfaces that have undergone abrasive techniques.3 The presence of carboxyl, hydroxyl, amine, and methyl functional groups has additionally been shown to modulate the adsorption of proteins which can influence cellular interactions.48 The strong adhesion of PAC to Zr is crucial for resisting mechanical delamination during in vivo implantation. The high adhesion strength of the PAC layer to Zr suggests that a pulsed bias of −500 V provides sufficient AIM to achieve strong adhesion to carbide forming materials without the use of intentionally graded interfaces,41a where the interface grading is a result of explicit gradual variation of the depositing flux composition. This represents a significant time and cost reduction in the coating fabrication process. The concept of AIM envisages the chemical alteration of the substrate surface during the plasma polymerization process, creating a newly modified interface that provides superior coating adhesion strength. While different mechanisms can account for AIM,40 our results are suggestive of strong and direct atomic bonding leading to the robust adhesion of PAC on Zr metal, most likely through the formation of carbide structures that would not be J


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tropoelastin also appears to be unaffected by sterilization, likely due to the inherent conformational flexibility of the protein,22 with the C-terminal region remaining accessible for solvent interactions. The preserved activity of the tropoelastin layer on PAC Zr is further confirmed by the maintenance of high cell attachment after sterilization, with the relative cell binding on PAC and standard Zr corresponding to the relative protein retention on each substrate. Interestingly, PAC Zr displayed complete poststerilization retention of tropoelastin bioactivity despite a slight decrease in intact protein levels. In the same manner, the poststerilization decrease in cell binding to tropoelastin-coated standard Zr did not reflect the more dramatic reduction in detected levels of intact protein. An explanation may be the cellular recognition of tropoelastin subregions that are not targeted by antibodies. Alternatively, these results may indicate the efficacy of tropoelastin in achieving bone cell adhesion, such that even partial protein coverage of the implant surface is sufficient for maximum cell response. One surprising observation is the apparent increase in cell attachment to PAC Zr after gamma irradiation. Although the underlying mechanisms are unclear, we surmise that gamma irradiation, as a source of ionizing energy similar to plasma, induces or enhances surface chemistry changes on plasma-modified materials that are favorable for cell interaction. The benefit of tropoelastin on bone implant materials extends beyond facilitating osteoblast adhesion to also promoting a high degree of proliferation. Tropoelastin-coated Zr surfaces consistently displayed increased bone cell numbers compared to bare or BSA-blocked controls. The amplified expansion of bone cells on the functionalized implant surfaces results in rapid cell coverage, which in turn is likely to expedite the bone remodeling events, including new bone formation, that rely on contact-dependent direct cell−cell signaling.59 The functional modifications on Zr also augment bone matrix maturation, characterized by osteoblast-specific marker expression of cells grown on these surfaces. Significant early, short-term up-regulation of ALP transcription was observed on tropoelastin-coated PAC Zr even in the absence of osteogenic supplements in media, whereas this effect was sustained longerterm in osteogenic media. ALP expression by cells on PAC Zr was also significantly increased compared to bare Zr, although early stage expression was lower than that on the tropoelastincoated sample. These results suggest that the regulation of ALP expression is likely attributed to tropoelastin activity, as well as the mechanical properties of the PAC layer consistent with the known role of substrate stiffness in modulating osteoblast differentiation and marker expression.59 The elevated ALP transcript levels on functionalized Zr are consistent with increased extracellular ALP enzyme production. ALP activity on tropoelastin-coated surfaces was significantly increased compared to that on PAC Zr, which in turn was higher than that on uncoated Zr. Increased ALP levels are associated with enhanced bone apposition on the implant surface, elevated local phosphorus concentrations, and removal of hydroxyapatite growth inhibitors for bone mineralization.60 In addition to ALP, transcript levels of OCN by cells on tropoelastin-coated PAC Zr also showed an early increase in nonosteogenic media, and a later-stage spike in osteogenic conditions. Osteocalcin, a marker of osteoblast maturation,12 binds calcium and phosphate to regulate the extent of hydroxyapatite crystal formation and mineral deposition.60 Expression of COL1, an osteoblast-

blocking step ensures that the observed cell binding can be attributed to tropoelastin rather than to the biocompatible bare or PAC Zr substrate. Furthermore, since the assay was performed in serum-free media, the cells could only bind directly to tropoelastin rather than through a bridging molecule such as a serum protein. The advantage conferred by tropoelastin in the context of cell attachment is evidenced by higher cell binding to protein-coated Zr compared to bare/PAC Zr. Also, whereas bare/PAC Zr supported cell adhesion under in vitro culture conditions, their activity cannot be guaranteed in vivo, where these surfaces can bind components that are bioinert or deleterious to bone healing. The activity of substrate-bound tropoelastin is not dependent on the pH of the surface association condition. Initially, we hypothesized that the negatively charged PAC layer may electrostatically attract the positively charged C-terminus of tropoelastin, a known cell-binding region within the protein, and therefore inhibit cell interaction similarly to previous observations.54 Accordingly, if tropoelastin were to associate with the surface at a pH above its isoelectric point, at which the C-terminus would be uncharged and unattracted to the PAC layer, the protein would be more likely to be covalently captured in an orientation that facilitates cell adhesion. However, the comparable cell attachment to PAC Zr surfaces, on which tropoelastin was coated at pHs above and below its pI, suggests that surface−protein charge interactions do not play a role in the functional presentation of tropoelastin in this case. A reason may be that the surface orientation of tropoelastin, which possesses a highly flexible structure,22 is not highly sensitive to electrostatic interactions between local protein regions and the underlying substrate. An alternative explanation may be that tropoelastin interacts with osteoblasts via other cell-binding region/s56 that are exposed and accessible even when the C-terminus is sequestered by charge interactions with the surface. The rapid, high-level bone cell adhesion to tropoelastincoated Zr indicates the potential of such functionalized implant materials to improve early stage osseointegration. Gaps at the bone-implant interface resulting from inadequate osseointegration increase the risk of implant failure.1b,18 Osseointegration is essential for the early recruitment of cells for debridement, and for the migration of mesenchymal cells to the implant surface where they differentiate into osteoblasts and deposit a collagen scaffold for bone regeneration.57 As osseointegration on Zr is already considered satisfactory, characterized by close contact of the material with bone without fibrous tissue or mononuclear cell accumulation at the interface,1b improvements due to the tropoelastin coating are more likely to reduce the time frame required to achieve the same extent of integration. The ability to sterilize tropoelastin-functionalized Zr is necessary for use in a clinical setting, and crucial for the commercial production of such an implant material. The sterilization methods tested in this study include steam autoclaving and gamma irradiation at a minimum of 25 kGy, which is the recommended and widely accepted dose for biological materials such as tissue allografts.58 Remarkably, the majority of the tropoelastin coating on PAC Zr remained intact and functional after either steam sterilization or gamma irradiation, in contrast to the significant drop in functional protein levels on standard Zr. This can be attributed to tropoelastin stabilization by covalent binding to PAC Zr, which again points to the necessity of PAC deposition on the substrate prior to protein coating. The surface orientation of K



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synthesized bone matrix protein that fills resorbed bone pits,60 remained unaltered after Zr surface modification. The early stage increase in osteoblast activity on functionalized Zr, combined with the accelerated cell proliferation on these surfaces, strongly points to the potential of such materials to stimulate bone mineralization. Accordingly, significantly more advanced calcification was observed on the functionalized Zr early postconfluence, which was sustained long-term in osteogenic conditions. Clearly, the functionalization of Zr provides a better biological and physical microenvironment for mineral deposition, particularly in an osteogenic environment such as in vivo where resident osteoblasts and bone marrowderived mesenchymal stem cells secrete osteogenic bone morphogenetic proteins and growth factors.61 Bone mineral content reflects the mechanical competence of the newly formed bone.10 Increased early mineralization therefore allows rapid mechanical reinforcement of the implant for improved shear strength, suggesting that materials manufactured from PAC modified, tropoelastin-coated Zr are likely to have higher removal torque values compared to those exhibited by standard Zr,2 thereby improving the chances of implant success. The two-step functionalization of zirconium with PAC and tropoelastin coating can be readily translated to commercial scale. PAC, as a nonline-of-sight surface modification technique, can be applied easily to complex geometries typical of bone fixation and augmenting devices. Plasma-based processes can be readily upscaled to allow high throughput batch coating of large total surface areas of multiple targets without excessive manipulation. Carbon based coating processes similar to that described in this manuscript have been successfully scaled up to create homogeneous coatings in industrial scale reactors.62 Iterative feedback control strategies combining plasma modeling, plasma diagnostics and coating characterization can be developed,63 enabling process optimization and reproducibility in upscaling scenarios. Recombinant tropoelastin is approved for therapeutic use in humans, and can be manufactured in commercial quantities at high purity without reliance on animal sources, in compliance with Good Manufacturing Practice guidelines.64

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected]. Tel: +61286271727. Address: Charles Perkins Centre D17, The University of Sydney, NSW 2006, Australia. *E-mail: [email protected]. Tel: +61293516079. Address: Applied & Plasma Physics Group, A28 School of Physics, The University of Sydney, NSW 2006, Australia. Author Contributions

M.S. designed, performed, and analyzed XPS, FTIR, nanoindentation, ellipsometry, and plasma deposition experiments. A.K. performed FTIR on 3D constructs. J.L. and G.C.Y. designed, performed, and analyzed real-time RT-PCR. G.C.Y. designed, performed, and analyzed all other experiments. G.C.Y. and M.S. wrote the manuscript. A.S.W. and M.M.B. provided feedback at all stages of experimentation and manuscript preparation. All authors have given approval to the final version of the manuscript. Author Contributions ||

A.S.W. and M.M.M.B. contributed equally as senior investigators in this work. Funding

We gratefully acknowledge LfC Sp z. o. o. of Zielona Góra, Poland, and the Australian Research Council for supporting this work under the ARC Linkage program (LP110200316). Gamma-ray irradiation services were provided by ANSTO as part of the Australian Institute of Nuclear Science and Engineering (AINSE) Research Award (ALNGRA13025 and ALNGRA14526). ASW acknowledges funding from the NIH (EB014283), the Australian Research Council, and the National Health and Medical Research Council (NHMRC) (APP1093307). JL acknowledges financial support by the Czech Health Research Council (15−32497A) and by the project “BIOCEV − Biotechnology and Biomedicine Centre of the Academy of Sciences and Charles University” (CZ.1.05/ 1.1.00/02.0109) from the European Regional Development Fund. Notes

The authors declare the following competing financial interest(s): ASW is the scientific founder of Elastagen Pty Ltd.

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ACKNOWLEDGMENTS The authors thank ANSTO for sample irradiation, and Connie Banos and Justin Davies for their assistance with the sterilization process. The authors are also thankful for use of the facilities as well as scientific and technical assistance at the Australian Centre for Microscopy and Microanalysis for SEM access and the School of Engineering, especially Dr. Yixiang Gan, for access to nano indentation equipment. The authors thank Dr Wojciech Chrzanowski for helpful discussions on the paper. The authors gratefully acknowledge project funding from LfC Sp z. o. o. of Zielona Góra, Poland, and the Australian Research Council under the ARC Linkage program and LfC for providing the spinal implant.

CONCLUSIONS

We have demonstrated a plasma coating process, readily applicable to 3D surfaces of orthopedic implants, that creates a mechanically robust carbon-nitride-based PAC onto metallic implant materials. When applied to carbide forming metals such as Ti and Zr, the process provides adhesion strong enough to resist severe stresses such as major plastic deformation of the underlying material. The mechanical stiffness of the PAC is well matched to that of bone and the radicals embedded in the coating enable the direct covalent immobilization of bioactive molecules which resist dynamic protein exchange processes that typically occur on surfaces in vivo. We describe the first use of tropoelastin, a protein classically utilized in elastic tissue engineering, to stably functionalize orthopedic implant materials for improved bone cell response. The thus-modified Zr material can withstand medical-grade sterilization and promotes increased bone adhesion, proliferation, matrix maturation, and mineralization for potentially accelerated bone repair.

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ABBREVIATIONS AIM, atomic interfacial mixing Aluminum, Al ALP, alkaline phosphatase BSA, bovine serum albumin COL1, collagen type I DOI: 10.1021/acsbiomaterials.6b00049 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

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ELISA, enzyme-linked immunosorbent assay FBS, fetal bovine serum OCN, osteocalcin PAC, plasma-activated coating RT-PCR, real-time polymerase chain reaction Si, silicon SS, stainless steel Ti, titanium V, vanadium XPS, X-ray photoelectron spectroscopy Zr, zirconium

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