pH-Controlled Release of Proteins from Polyelectrolyte-Modified

Sep 10, 2012 - Citation data is made available by participants in Crossref's Cited-by Linking service. For a more ... Roman A. Perez , Seog-Jin Seo , ...
2 downloads 13 Views 5MB Size
Download PDF
Article pubs.acs.org/Biomac

pH-Controlled Release of Proteins from Polyelectrolyte-Modified Anodized Titanium Surfaces for Implant Applications Amy M. Peterson,*,† Helmuth Möhwald,† and Dmitry G. Shchukin†,‡ †

Department of Interfaces, Max Planck Institute of Colloids and Interfaces, 14476 Potsdam-Golm, Germany Stephenson Institute for Renewable Energy, University of Liverpool, Liverpool L69 3BX, United Kingdom

‡

ABSTRACT: Titanium is a popular choice of implant material given its strength, durability, and biocompatibility; however, strong interfaces with the surrounding tissue are not achieved, resulting in stress shielding and implant loosening. One option for improving adhesion is modification of the surface chemistry and topography through anodization, while another option is coating the titanium surface with a protein-eluting polyelectrolyte complex. Morphogenetic proteins such as BMP-2 have been shown to cause cell migration, expression of different genes, and development of different tissues. Anodization was used to form a porous oxide structure across the surface. A polyelectrolyte coating of poly-l-histidine and poly(methacrylic acid) was prepared and was shown to be effective for sustained release of negatively charged species under physiological conditions. This complex demonstrated pH-dependent release, with maximum release at pH = 5−6, but low levels of sustained release at pH = 7−8. Smaller initial burst release and higher amounts of sustained release were observed when lower molecular weight poly(methacrylic acid) was used. Different methods of loading the polyelectrolyte with the model species were compared. Immersion of the coating for loading provided greater release, albeit a larger initial burst release.

■

biointerfaces.7,8 However, much is still unknown given the complexity of interactions at the interface and the many parameters (including topography, molecular design, biofouling, cell-material interactions, inter- and intramolecular forces) that control interactions at the interface.9 Morphogens, biomolecules that act as spatial regulators, determine cell behavior and tissue development through concentration gradients. Morphogen gradients can cause cell migration, expression of different genes, and development of different tissues.10−12 Many morphogens, in particular, transcription and growth factors, have been investigated for their ability to control osteoblast outcomes and bone formation. One of these, Cbfa1, is an osteoblast-specific transcription factor that is essential for osteoblast differentiation as well as bone formation. Growth factors such as bone morphogenic proteins (BMPs), transforming growth factor-β (TGF-β), members of the fibroblast growth factor (FGF) family and Indian hedgehog (Ihh) have been implicated in Cbfa1 expression, and therefore, osteoblast differentiation.13 Growth factors play key roles in regulating osteoblast behavior and osteoid and bone formation.14 As such, they were selected as the molecules of interest for encapsulation and pH-controlled release. Polyelectrolytes are polymers with positively (polycation) or negatively (polyanion) charged repeat units. When polycations and polyanions are deposited on a surface in an alternating

INTRODUCTION The interactions at the interface between the body and an implanted device dictate the success of the device. As synthetic materials are increasingly used as stents, bone scaffolds, and hip and other implants, it is essential to improve the understanding of biointerfaces and to tailor the interface for a given application. Titanium is a popular choice of implant material given its strength, durability, and biocompatibility; however, strong interfaces with the surrounding tissue are not achieved, resulting in stress shielding and implant loosening. In America alone, over 1.14 million arthroplasties were performed in 2010, with this number increasing every year.1 Despite titanium’s superior properties, 7% of these procedures were for replacement of hip and knee implants, at a cost of $5.54 billion. Implant failure is a misnomer; it is not the titanium device that fails, but rather the interface between the implant and bone that fails due to insufficient osseointegration of the implant and/or stress shielding due to the stiffness mismatch between implant and bone. Bacterial adsorption on the implant surface and the resulting biofilm formation are also significant causes of implant failure, as well as bacterial infection, which can lead to illness, amputation, and even mortality.2−5 If the implanted device surface is incompatible, the device will be rejected. Even if the surface is not incompatible, there can be limited integration of the implant within the body as a result of limited adhesion and growth of desired cells, either as a result of poor surface properties or adhesion of bacteria and other undesired cells and proteins.6 Over the past two decades significant strides have been made in developing improved © 2012 American Chemical Society

Received: June 18, 2012 Revised: August 31, 2012 Published: September 10, 2012 3120

dx.doi.org/10.1021/bm300928s | Biomacromolecules 2012, 13, 3120−3126

Biomacromolecules

Article

fashion, electrostatic interactions form a pH-sensitive film that exhibits a “closed” state when the opposing charges of the polyelectrolytes neutralize each other, but an “open” state when the environment has to provide additional charges to neutralize either the polyanion (low pH) or polycation (high pH). Polyanions and polycations in polyelectrolyte films produced in a layer-by-layer (LbL) fashion are highly entangled with each other, which allows for closing and opening of the films several times.15 Polyelectrolyte films and microcapsules have been used for a variety of applications, including drug delivery,16 microreactors for synthesis of difficult to achieve crystalline nanomaterials,17 and encapsulation of corrosion inhibitors for self-healing coatings.18 The properties of a polyelectrolyte film are dependent upon many processing parameters including polyelectrolyte pair, molecular weight,15 chain structure (linear vs branched), deposition pH,19 and number of layers.20 Therefore, the properties of the film can be tailored for a given application. Polyelectrolyte complexes have been used for decades for the sustained release of proteins and drugs.21 In addition to polyelectrolyte films, microcapsules based on polyelectrolytes have been used for controlled release for biological applications.16,22−24 In a recent example, basic fibroblast growth factor (FBF2) was encapsulated in dextran sulfate and poly-larginine. These capsules demonstrated fast in vitro release kinetics (release of FGF2 to an optimal level within one day) while also maintaining this optimal level for days.24 Recently, Macdonald et al. demonstrated the coating of a polymer scaffold with LbL-deposited poly(β-aminoester) and chondroitin sulfate, a complex capable of delivering microgram scale amounts of BMP-2.25 Poly(L-lysine) (PLL)/hyaluronic acid (HA) coatings on a porous ceramic also showed microgram level release of BMP-2 from porous ceramic scaffolds.26 However, in this case, over 60% of release was observed in the first day. Tryoen-Tóth et al. demonstrated that polyelectrolyte coatings terminating in poly(sodium 4-styenesulfonate) (PSS), poly(L-glutamic acid) (PGA), and PLL show good biocompatibility for osteoblast-like cells.27 However, Schultz et al. reported a more regular and less obstructed fibrobrast layer on PGA-terminated coatings as compared to PLL-terminated coatings. Additionally, by functionalizing the PGA layer with an anti-inflammatory peptide, in vivo production of an anti-inflammatory was detected.28 A polyelectrolyte coating of HA and chitosan was developed by Chua et al. to confer antibacterial properties.29 When arginineglycine-aspartic acid was immobilized on this coating, osteoblast adhesion was also significantly improved as compared to pristine titanium.30 Other polyelectrolyte coatings that have been shown to improve cell adhesion include PSS/ poly(allyamine hydrochloride) (PAH),31 chitosan/sulfated chitosan,32 chitosan/heparin,33 protamine sulfate/PSS,34 and chitosan/alginate.35 Much of this work focuses on developing coatings that are stable over long time scales, leading to cross-linking of some of these systems to reduce the amount of coating degradation over time. However, as porous titanium implants achieve greater popularity due to mechanical properties comparable to bone and increased potential for osseointegration, degradable coatings become more desirable.36−41 The focus of this work is the engineering of biocompatible LbL polyelectrolyte coatings for encapsulation/immobilization of growth factors, which play key roles in regulating osteoblast behavior and osteoid and bone formation. Additionally,

anodization was used to modify the structure of the titanium surface through the formation of a porous titanium dioxide structure. Anodization has previously been shown to be an effective strategy for improving the adhesion of osteoblasts to titanium surfaces, since topographical features and surface chemistry significantly influence cell shape and behavior.42−44 The following report contains the first example of a proteinloaded coating on a titanium surface with pH-controlled release and removal of the coating. By developing a biocompatible coating that dissolves and then degrades under physiological conditions, the titanium surface can be protected from biofouling prior to implantation, and morphogenetic proteins can be released under physiological conditions. For proof of concept studies, model proteins were sandwiched between the titanium surface and a polyelectrolyte complex. At physiological pH, the polyelectrolyte coating fell apart, releasing proteins and exposing the titanium surface. The focus of this research is the development of biocompatible polyelectrolyte coatings with the ability to release growth factors for application in implanted devices. The release profiles under different pHs were investigated, as was the role of polyelectrolyte molecular weight and procedure for loading of the model protein on release behavior. In order to expose the titanium surface for cell adhesion and osseointegration, the polyelectrolyte film must dissolve and release growth factors under physiological conditions.

■

EXPERIMENTAL SECTION

Materials. Titanium plates (99.5% Ti) were acquired from Alfa Aesar and polished prior to use. Sulfuric acid was used to clean titanium prior to anodization and also acted as the anodizing electrolyte. Poly(methacrylic acid) (PMA, Mn ≈ 100 kDa, PDI = 7.9) was used as received from Polysciences. Poly(methacrylic acid sodium salt) (PMA-2, Mn ≈ 5400, PDI = 1.8), poly-l-histidine (PH), and fluorescently labeled poly-l-lysine (PL-FITC) were used as received from Sigma-Aldrich. Increased polydispersity will cause slower decay of a complex as longer chains contain more charged groups, which bond with more oppositely charged chains throughout a greater volume of the coating.45 In this study, PL-FITC acts as a model compound, standing in for a protein. A further discussion of the use of PL-FITC is included in the Release Studies section. Titanium Preparation. Prior to anodization, the titanium plate was cleaned with 1.5 M H2SO4. Titanium was anodized in a 165 g L−1 solution of H2SO4 at a potential of 30 V for 5 min. Anodization under these conditions results in a porous oxide structure with pores ranging in size from 40 to 200 nm in diameter. Titration. Aqueous solutions containing 1 mg mL−1 of polyanion and/or polycation were titrated against strong acids or bases (HCl or NaOH) using a Metrohm Autotitrator. Titration allowed for determination of pKa's and approximate pH ranges in which polyelectrolyte complexes exhibit an open state. Polyelectrolyte Coating. Coating of the titanium with the polyelectrolyte was achieved by first immersing the titanium specimen in a 1 mg mL−1 PH solution for 15 min. After immersion, the specimen was removed, washed three times in water, and then dipped in a 1 mg mL−1 PMA solution, leaving the specimen immersed for 15 min. Following washing three times in water, this procedure was repeated until the titanium specimen had been dipped five times in each polyelectrolyte, for a total of 10 layers. This specimen is denoted as (PH/PMA)5. The method of PL-FITC loading was investigated. The base method was to adsorb PL-FITC to the titanium plate by immersing the plate in a 0.1 mg mL−1 solution of PL-FITC for 15 min. The polyelectrolyte coating was then formed on top of this adsorbed layer by immersing the plate in PMA, then alternating layers of PH and PMA until 10 layers (5 bilayers) are achieved. PL-FITC was also incorporated in the coating by first preparing the (PH/PMA)5 coating 3121

dx.doi.org/10.1021/bm300928s | Biomacromolecules 2012, 13, 3120−3126

Biomacromolecules

Article

and immersing the coated titanium plate in 0.1 mg mL−1 PL-FITC for 1 h. The polycation method was to adsorb PH to the titanium plate and adsorb PL-FITC to this layer by immersion in the PL-FITC solution for 15 min and then continue complex formation of the PLFITC layer. The polyanion method was to adsorb a layer each of PH and PMA to the titanium plate and then adsorb PL-FITC to this layer by immersion in the PL-FITC solution for 15 min, followed by continuing the formation of the polyelectrolyte coating. After the coating was completed, specimens were dried with compressed air and left overnight under ambient conditions. Scanning electron miscroscopy (SEM) was used to verify that specimens were coated evenly with the polyelectrolyte. Release from Coatings. To evaluate release of PL-FITC from PH/PMA and PH/PMA-2 coatings, specimens were immersed in individual buffered solutions over a range of pHs. At least three specimens were used per condition. Aliquots were taken from the solutions daily, with the aliquot volume replaced with fresh buffer. The collected solution was investigated using florescence spectroscopy with a FluoroMax fluorometer from Jobin Yvon. Given the pH-dependence of FITC fluorescence, calibration curves were constructed for each pH investigated using the same buffer solutions. These curves were used to correlate intensity to amount of PL-FITC released.

approximation of where the polyelectrolyte complex is in its open state. The pKa values for PMA and PMA-2 were found to be practically the same (6.7 and 6.8, respectively). Since pKa is related to the strength of an acid in solution, similarity in pKa values indicates that there are minimal differences between the acid strengths of PMA and PMA-2. This result is important because the acid groups of PMA are methacrylic acid, while the acid groups of PMA-2 are methacrylic acid and sodium salt. Therefore, any differences between properties of polyelectrolyte complexes containing PMA and PMA-2 are the result of other properties, mainly molecular weight, and are not related to different dissociation constants of the acid groups. PH/PMA is insoluble in water over the pH range 5.7−7.7, while PH/PMA-2 is insoluble in water over the range pH = 4.9−9.0. The difference in ranges is most likely due to differences in molecular weight and/or coiling of the polyanions. Release Studies. Figure 2 shows a SEM image of anodized titanium compared to SEM images of anodized titanium coated with 5 bilayers of PH/PMA and PH/PMA-2. Anodization results in a porous oxide structure with pores ranging in size from 40 to 200 nm in diameter. After three bilayers, there is a relatively even coating of the polyelectrolyte, and the porous anodized structure was covered. In this study, the polypeptide PL-FITC acts as a model compound, standing in for a protein. The molecular weight of PL-FITC (15−30 kDa) compares well with the apparent molecular weight of BMP-2 (26 kDa) and other proteins of interest to bone growth, although the isoelectric point values of PL-FITC and BMP-2 (6.8 and 8.5, respectively) differ.25 By using PL-FITC, the mechanism of release can also be determined. If release is diffusion-limited, release profiles will be the same at all pHs. In this case, pore structure of the coating determines release behavior. However, if release is controlled primarily by electrostatics, then release will be pHdependent. Release at different pHs from (PH/PMA)5 is shown in Figure 3. (PH/PMA)5 shows microgram-scale release of PLFITC at all pHs. The absolute amount of release increases and then decreases with increasing pH, with the maximum amount of release achieved at pH = 5. The amount of release per unit area after 25 days ranges from 3.0−4.3 ng mm−2. In all cases, there is a relatively large burst release during the first day, with the size of this burst increasing with increasing pH. Over the first 7−10 days, the rate of release is relatively constant for pH = 4−6, followed by a different, lower rate of release after 10 days. For pH = 7−8, the amount of release effectively plateaus after four days. However, the effect of pH on the absolute

■

RESULTS AND DISCUSSION Polyelectrolyte Selection. In order to consider a polyelectrolyte for use in a protein-loaded coating for implantation in bone, it must be biocompatible and the polyelectrolyte complex must be in its open state at physiological pHs. PMA and PH are both biocompatible.46 By titrating both individually and together, pKa values were determined. Figure 1 gives the titration curves for PMA, PMA-

Figure 1. Titration curves for PMA, PMA-2, PH, PMA/PH, and PMA2/PH.

2, PH, PMA/PH and PMA-2/PH. Additionally, the pH range in which the polyelectrolyte complexes are insoluble gives an

Figure 2. SEM images of (a) anodized titanium, (b) anodized titanium coated with 5 bilayers of PH/PMA, and (c) anodized titanium coated with 5 bilayers of PH/PMA-2. Scale bars represent 200 nm. 3122

dx.doi.org/10.1021/bm300928s | Biomacromolecules 2012, 13, 3120−3126

Biomacromolecules

Article

Figure 4. Release from (PH/PMA)5imm.

Figure 3. Release at different pHs from (PH/PMA)5.

amount of PL-FITC release from the coating is not statistically significant (p > 0.05). The release behavior compares well with solution experiments, which show insolubility over the pH range 5.7−7.7. Titration experiments were performed on a solution, while release experiments were performed using a polyelectrolyte complex dip-coated on a titanium surface. Slight differences in the acid−base equilibrium are expected for different complex formation techniques.11 Assuming electrostatics is the rate limiting release mechanism, the greatest release is related to a more open state of the polyelectrolyte complex; therefore dipcoated (PH/PMA)5 is most open around pH = 5−6. The pH of bone varies naturally to some extent. While the pH of arterial blood is approximately 7.4, the pH of the interstitial fluid in contact with bone tissue is lower. At pH ≥ 7.4, matrix mineralization occurs, while osteoclast behavior is limited. At pH ≤ 7.0, matrix mineralization is inhibited and osteoclasts are activated to resorb bone. Bone maintenance is achieved at a pH of approximately 7.2.47 Therefore, coatings should be tuned to optimal release at pH = 7.2. In this range, (PH/PMA)5 demonstrates a quick burst release of 1.5−2 μg PL-FITC, followed by almost complete release of 2−2.3 μg after 5 days. Sustained release of morphogenetic proteins is desired for bone implantation, so this combination of polyelectrolytes and loading technique is not ideal. However, alternate loading techniques may provide different release profiles. Effect of Loading Technique. Release from (PH/PMA)5 immersed in PL-FITC, described herein as (PH/PMA)5imm, was evaluated over the same range of pHs. Results are shown in Figure 4. Similar trends are observed in (PH/PMA)5imm as in (PH/PMA)5. The maximum amount of PL-FITC released is much greater at pH = 5, with 10 ± 3.2 μg PL-FITC released, or 16 ± 5.1 ng mm−2. However, at physiological pHs, release is not significantly different than from (PH/PMA)5 coatings. The difference between release at pH = 5 and at other pH values is statistically significant (p = 0.019), suggesting that electrostatics are important in dictating release kinetics. Since release was greatest at pH = 5 for (PH/PMA)5 and (PH/PMA)5imm, two other PL-FITC loading techniques were investigated with release evaluated at pH = 5. As described in the Experimental Section, PL-FITC was adsorbed to PH (herein described as (PH/PMA)5PHad) or PMA (herein denoted as (PH/PMA)5PMAad). Comparison of the release profiles for different loading techniques is shown in Figure 5.

Figure 5. Comparison of the release profiles for different loading techniques.

While the amount of PL-FITC release is dependent on the loading technique, release in all cases follows the same general trend. Over the first 7−10 days, the rate of release is relatively constant, followed by a different, lower rate of release. The smallest amount of PL-FITC is released when it is initially adsorbed to PH, a polycation. Since PH and PL-FITC have the same charge under the dipping conditions, only a monolayer of PL-FITC will be adsorbed to the PH. Since anodized titanium (titanium dioxide) and PMA have a negative charge under the dipping conditions, more PL-FITC is adsorbed in these cases.48 This indicates that PMA is more negatively charged than the anodized titanium surface. Alternative loading techniques can be used to change the burst release as well as the total amount of PL-FITC release; however, the profile is not significantly affected by the loading technique. With that in mind, a different polyanion, PMA-2, was evaluated. Effect of Different Polyanions. (PH/PMA-2)5 was prepared in the same manner as (PH/PMA)5, and the release at different pHs was observed. Results are given in Figure 6. Smaller burst release is observed in the first day as compared to that of (PH/PMA)5. The general release profiles are similar for pH = 4−6. However, for pH = 7−8, the range of interest for physiological release, there is a smaller (0.5−1 μg) burst release followed by a constant release of 0.02−0.03 μg day−1 PL-FITC from the fourth day on. After 50 days, this relatively small release rate was still observed. The differences between release at pH = 4−6 and at pH = 7−8 is statistically significant (p = 3123

dx.doi.org/10.1021/bm300928s | Biomacromolecules 2012, 13, 3120−3126

Biomacromolecules

Article

Figure 6. Release from (PH/PMA-2)5 at different pHs.

0.0061). Similar behavior was recorded in the (PH/PMA)5 system after the first week, although in the lower molecular weight system, the initial release is decreased by over half. The difference between release from (PH/PMA)5 and (PH/PMA2)5 results from the molecular weight difference between PMA and PMA-2. The lower molecular weight of PMA-2 may have resulted in a more compact coating that decreases diffusivity. Release results from all systems suggest that both diffusion and electrostatic effect play important roles in release from coatings consisting of poly-l-histidine and PMA. The burst release observed is most likely related to the release of electrostatically bound PL-FITC, whereas longer term sustained release comes from physically entrapped PLFITC. To explore this behavior further, an anodized piece of titanium plate was immersed in the PL-FITC solution for 15 min. Desorption of PL-FITC from the titanium plate in pH = 7 and pH = 8 buffered solutions was evaluated as a way to look at burst release. In this situation, there is no physical entrapment. As a result, half of release occurred within one hour of immersion in the buffered solution and over 95% of release was achieved within one day. A potential method for reducing the burst release is to rinse off the electrostatically bound material prior to release study/implantation. After 25 days at pH = 5, the percentage of PL-FITC released from different coatings is as follows: (PH/PMA)5 = 78%; (PH/ PMA)5imm =88%; (PH/PMA-2)5 = 65%. The differences in release over this time period correlate well to differences in initial burst release. Table 1 summarizes the percentage of total release after 25 days as a function of pH for different coatings.

Figure 7. Total loading of PL-FITC in different coatings.

release of morphogenetic proteins to induce bone growth. Figure 8 displays SEM images of anodized titanium and anodized titanium coated (PH/PMA-2)5 after 25 days of release at pH = 7. Since PMA-2 exhibits lower polydispersity and molecular weight than PMA-2, it is expected that this coating will dissolve faster than (PH/PMA)5. At pH = 7, the coating is completely removed from the anodized titanium surface. SEM images of uncoated anodized titanium and coated titanium after 25 days are similar to one another and do not have the same appearance as the polyelectrolyte-coated surface prior to immersion in the buffered solution. The presence of PMA in the buffered solution, confirmed by Raman spectroscopy, supports these results. Given that there is continued slow release after 25 days, it is reasonable that some coating remains on the surface shown in Figure 8b. Since no release is possible once dissolution of the coating is complete, dissolution of the coating should be tuned to the desirable release time scale (months to years in the case of growth factor release). With this in mind, dissolvable coatings may be better suited for microporous titanium surfaces, where slow dissolution of a polyelectrolyte coating can provide both low levels of sustained release and controlled diffusion of osteoblasts into micropores.

■

CONCLUSIONS A polyelectrolyte coating of PH and PMA was prepared and was shown to be effective for sustained release of negatively charged species under physiological conditions. This complex demonstrated pH-dependent release, with low levels of sustained release at pH = 7−8. Smaller initial burst release and higher amounts of sustained release were observed when a lower molecular weight PMA was used. Different methods of loading the polyelectrolyte with a model species were investigated. Immersion of the coating for loading provided greater overall release and a larger initial burst release than loading by adsorption on the anodized titanium. Importantly, controlled release on the microgram scale over 25 days was shown at physiological pH, which is advantageous and necessary for the desired in vivo effect, i.e., signaling of osteoblasts to the implant surface.

Table 1. Percentage of Total Release on Day 25 pH

(PH/PMA)5

(PH/PMA)5imm

(PH/PMA-2)5

4 5 6 7 8

71.3 75.8 82.9 68.7 49.9

81.9 87.9 82.5 86.6 87.6

55.7 62.1 68.9 58.1 64.2

The total average release from different coatings is given in Figure 7. Unsurprisingly, (PH/PMA)5imm coatings demonstrate the largest loading and the largest standard deviation in total loading. Dissolution of Coating. In this system, a balance must be struck between dissolution of the coating to allow for cell adhesion and bone growth on the titanium and the sustained 3124

dx.doi.org/10.1021/bm300928s | Biomacromolecules 2012, 13, 3120−3126

Biomacromolecules

Article

Figure 8. SEM images of (a) uncoated anodized titanium and (b) anodized titanium coated with (PH/PMA-2)5 after 25 days of release at pH = 7. Scale bars represent 1 μm. (10) Gordon, P. V.; Sample, C.; Berezhkovskii, A. M.; Muratov, C. B.; Shvartsman, S. Y. Proc. Natl. Acad. Sci. U.S.A. 2011, 108 (15), 6157. (11) Zeng, X.; Goetz, J. A.; Suber, L. M.; Scott, W. J., Jr.; Schreiner, C. M.; Robbins, D. J. Nature 2001, 411 (6838), 716−720. (12) Dessaud, E.; McMahon, A. P.; Briscoe, J. Development 2008, 135 (15), 2489−2503. (13) Ducy, P.; Schinke, T.; Karsenty, G. Science 2000, 289 (5484), 1501. (14) Burdick, J. A.; Mason, M. N.; Hinman, A. D.; Thorne, K.; Anseth, K. S. J. Controlled Release 2002, 83 (1), 53−63. (15) Mauser, T.; Déjugnat, C.; Möhwald, H.; Sukhorukov, G. Langmuir 2006, 22, 5888−5893. (16) Shchukina, E. M.; Shchukin, D. G. Adv. Drug Delivery Rev. 2011, 63 (9), 837−846. (17) Shchukin, D. G.; Sukhorukov, G. B. Adv. Mater. 2004, 16 (8), 671−682. (18) Andreeva, D. V.; Fix, D.; Möhwald, H.; Shchukin, D. G. Adv. Mater. 2008, 20, 2789−2794. (19) Antipov, A. A.; Sukhorukov, G. B.; Leporatti, S.; Radtchenko, I. L.; Donath, E.; Möhwald, H. Colloids Surf., A 2002, 198, 535−541. (20) Antipov, A. A.; Sukhorukov, G. B.; Donath, E.; Möhwald, H. J. Phys. Chem. B 2001, 105 (12), 2281−2284. (21) Tabata, Y.; Ikada, Y. Adv. Drug. Deliv. Rev. 1998, 31 (3), 287− 301. (22) Sato, K.; Yoshida, K.; Takahashi, S.; Anzai, J. Adv. Drug Delivery Rev. 2011, 63, 809−821. (23) Sukhorukov, G. B.; Volodkin, D. V.; Günther, A. M.; Petrov, A. I.; Shenoy, D. B.; Möhwald, H. J. Mater. Chem. 2004, 14, 2073. (24) She, Z.; Wang, C.; Li, J.; Sukhorukov, G. B.; Antipina, M. N. Biomacromolecules 2012, 13, 2174−2180. (25) Macdonald, M. L.; Samuel, R. E.; Shah, N. J.; Padera, R. F.; Beben, Y. M.; Hammond, P. T. Biomaterials 2011, 32 (5), 1446−1453. (26) Crouzier, T.; Sailhan, F.; Becquart, P.; Guillot, R.; LogeartAvramoglou, D.; Picart, C. Biomaterials 2011, 32 (30), 7543−7554. (27) Tryoen-Tóth, P.; Vautier, D.; Haikel, Y.; Voegel, J.-C.; Schaaf, P.; Chluba, J.; Ogier, J. J. Biomed. Mater. Res. 2002, 60 (4), 657−667. (28) Schultz, P.; Vautier, D.; Richert, L.; Jessel, N.; Haikel, Y.; Schaaf, P.; Voegel, J.-C.; Ogier, J.; Debry, C. Biomaterials 2005, 26, 2621− 2630. (29) Chua, P. H.; Neoh, K. G.; Shi, Z.; Kang, E. T. J. Biomed. Mater. Res. 2008, 87A (4), 1061−1074. (30) Chua, P.-H.; Neoh, K.-G.; Kang, E.-T.; Wang, W. Biomaterials 2008, 29, 1412−1421. (31) Brunot, C.; Grosgogeat, B.; Picart, C.; Lagneau, C.; JaffrezicRenault, N.; Ponsonnet, L. Dent. Mater. 2008, 24, 1025−1035. (32) Li, Q.-L.; Huang, N.; Chen, J.; Chen, C.; Chen, J.; Chen, H. J. Bioact. Compat. Polym. 2009, 24 (2), 129−150. (33) Schweizer, S.; Schuster, T.; Junginger, M.; Siekmeyer, G.; Taubert, A. Macromol. Mater. Eng. 2010, 295 (6), 535−543. (34) Samuel, R. E.; Shukla, A.; Paik, D. H.; Wang, M. X.; Fang, J. C.; Schmidt, D. J.; Hammond, P. T. Biomaterials 2011, 32 (30), 7491− 7502. (35) Miranda, E. S.; Silva, T. H.; Reis, R. L.; Mano, J. F. Tissue Eng., Part A 2011, 17 (21−22), 2663−2674.

The PH and PMA coating was mostly removed after 25 days, exposing the porous anodized titanium surface. Coatings capable of dissolution under physiological conditions are wellsuited to application on porous titanium surfaces because their removal from the surface exposes the porous structure to cells, potentially allowing for greater osseointegration of the titanium. Development of novel biointerfaces is an active field of research. Improved surfaces for titanium implants offer the possibility of longer implant life and better quality of life for the millions of people that receive titanium implants annually. Ongoing work includes in vitro evaluation, longer-term investigation of release, and modification of these coatings to further reduce initial burst release under physiological conditions.

■

AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Author Contributions

The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. Notes

The authors declare no competing financial interest.

■

ACKNOWLEDGMENTS The authors acknowledge the EU FP7 Project NANOMAR for funding. The authors also thank Dimitriya Borisova for assistance with collecting SEM images.

■

REFERENCES

(1) HCUPnet Hompage. http://hcupnet.ahrq.gov/HCUPnet.jsp (accessed 06/11/2012). (2) Ercan, B.; Kummer, K. M.; Tarquinio, K. M.; Webster, T. J. Acta Biomater. 2011, 7, 3003−3012. (3) Shirai, T.; Shimizu, T.; Ohtani, K.; Zen, Y.; Takaya, M.; Tsuchiya, H. Acta Biomater. 2011, 7, 1928−1933. (4) Lee, J.-H.; Wang, H.; Kaplan, J. B.; Lee, W. Y. Acta Biomater. 2010, 6, 4422−4429. (5) Werner, S.; Huck, O.; Frisch, B.; Vautier, D.; Elkaim, R.; Voegel, J.-C.; Brunel, Ge.; Tenenbaum, H. Biomaterials 2009, 30, 2291−2301. (6) Folch, A.; Toner, M. Annu. Rev. Biomed. Eng. 2000, 2 (1), 227− 256. (7) Somorjai, G.; Frei, H.; Park, J. J. Am. Chem. Soc. 2009, 131 (46), 16589−16605. (8) Stuart, M. A. C.; Huck, W. T. S.; Genzer, J.; Müller, M.; Ober, C.; Stamm, M.; Sukhorukov, G. B.; Szleifer, I.; Tsukruk, V. V.; Urban, M.; Winnik, F.; Zauscher, S.; Luzinov, I.; Minko, S. Nat. Mater. 2010, 9 (2), 101−113. (9) Tiefenauer, L.; Ros, R. Colloid Surface B 2002, 23, 95−114. 3125

dx.doi.org/10.1021/bm300928s | Biomacromolecules 2012, 13, 3120−3126

Biomacromolecules

Article

(36) Baril, E.; Lefebvre, L. P.; Hacking, S. A. J. Mater. Sci.: Mater. Med. 2011, 22 (5), 1321−1332. (37) Li, J. C.; Dunand, D. C. Acta Mater. 2011, 59 (1), 146−158. (38) Vrana, N. E.; Dupret, A.; Coraux, C.; Vautier, D.; Debry, C.; Lavalle, P. PLoS ONE 2011, 6 (5), e20480. (39) Siegkas, P.; Tagarielli, V. L.; Petrinic, N.; Lefebvre, L. P. J. Mater. Sci. 2011, 46 (8), 2741−2747. (40) Singh, R.; Lee, P. D.; Jones, J. R.; Poologasundarampillai, G.; Post, T.; Lindley, T. C.; Dashwood, R. J. Acta Biomater. 2010, 6 (12), 4596−4604. (41) Kashef, S.; Asgari, A.; Hilditch, T. B.; Yan, W.; Goel, V. K.; Hodgson, P. D. Mater. Sci. Eng., A 2010, 527 (29−30), 7689−7693. (42) Bernard, S. A.; Balla, V. K.; Davies, N. M.; Bose, S.; Bandyopadhyay, A. Acta Biomater. 2011, 7 (4), 1902−1912. (43) Zinger, O.; Anselme, K.; Denzer, A.; Habersetzer, P.; Wieland, M.; Jeanfils, J.; Hardouin, P.; Landolt, D. Biomaterials 2004, 25 (14), 2695−2711. (44) Kim, Y.; Jang, J.-H.; Ku, Y.; Koak, J.-Y.; Chang, I.-T.; Kim, H.-E.; Lee, J.-B.; Heo, S.-J. Biotechnol. Lett. 2004, 26 (5), 399−402. (45) Dunlop, I. E.; Briscoe, W. H.; Titmuss, S.; Jacobs, R. M. J.; Osborne, V. L.; Edmonson, S.; Huck, W. T. S.; Klein, J. J. Phys. Chem. B 2009, 113, 3947−3956. (46) Detzel, C. J.; Larkin, A. L.; Rajagopalan, P. Tissue Eng., Part B: Rev. 2011, 17 (2), 101−113. (47) Arnett, T. R. J. Nutr. 2008, 138 (2), 415S−418S. (48) Ladam, G.; Schaaf, P.; Cuisinier, F. J. G.; Decher, G.; Voegel, J.C. Langmuir 2001, 17 (3), 878−882.

3126

dx.doi.org/10.1021/bm300928s | Biomacromolecules 2012, 13, 3120−3126