Brushite Bilayer Coating on

Subscriber access provided by University of Newcastle, Australia

Article

A facile preparation of poly(lactic acid)/brushite bilayer coating on biodegradable magnesium alloys with multiple functionalities for orthopedic application Lei Zhang, Jia Pei, Haodong Wang, Yongjuan Shi, Jialin Niu, Feng Yuan, Hua Huang, Hua Zhang, and Guangyin Yuan ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b00209 • Publication Date (Web): 28 Feb 2017 Downloaded from http://pubs.acs.org on March 4, 2017

Just Accepted “Just Accepted” manuscripts have been peer-reviewed and accepted for publication. They are posted online prior to technical editing, formatting for publication and author proofing. The American Chemical Society provides “Just Accepted” as a free service to the research community to expedite the dissemination of scientific material as soon as possible after acceptance. “Just Accepted” manuscripts appear in full in PDF format accompanied by an HTML abstract. “Just Accepted” manuscripts have been fully peer reviewed, but should not be considered the official version of record. They are accessible to all readers and citable by the Digital Object Identifier (DOI®). “Just Accepted” is an optional service offered to authors. Therefore, the “Just Accepted” Web site may not include all articles that will be published in the journal. After a manuscript is technically edited and formatted, it will be removed from the “Just Accepted” Web site and published as an ASAP article. Note that technical editing may introduce minor changes to the manuscript text and/or graphics which could affect content, and all legal disclaimers and ethical guidelines that apply to the journal pertain. ACS cannot be held responsible for errors or consequences arising from the use of information contained in these “Just Accepted” manuscripts.

ACS Applied Materials & Interfaces is published by the American Chemical Society. 1155 Sixteenth Street N.W., Washington, DC 20036 Published by American Chemical Society. Copyright © American Chemical Society. However, no copyright claim is made to original U.S. Government works, or works produced by employees of any Commonwealth realm Crown government in the course of their duties.

Page 1 of 34

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

ACS Applied Materials & Interfaces

A facile preparation of poly(lactic acid)/brushite bi-layer coating on biodegradable magnesium alloys with multiple functionalities for orthopedic application Lei Zhang†, Jia Pei*, †, Haodong Wang, Yongjuan Shi, Jialin Niu, Feng Yuan, Hua Huang, Hua Zhang, Guangyin Yuan* National Engineering Research Center of Light Alloy Net Forming & State Key Laboratory of Metal Matrix Composite, Shanghai Jiao Tong University, Shanghai, 200240, China *Corresponding author. E-mail: [email protected]; Tel.: +86 21 34203051; fax: +86 21 3420 2794. * Corresponding author. E-mail: [email protected]; Tel.: +86 21 34203051; fax: +86 21 3420 2794. † Lei Zhang and Jia Pei contribute equally to this work.

ABSTRACT Recently magnesium and its alloys has been proposed as a promising next generation orthopedic implant material, whereas the poor corrosion behavior, potential cytotoxicity and the lack of efficient drug delivery system have limited its further clinical application, especially for the local treatment of infections or musculoskeletal disorders and diseases. In this study, we designed and developed a multifunctional bi-layer composite coating of poly(lactic acid)/brushite with high interfacial bonding strength on a Mg-Nd-Zn-Zr alloy, aiming to improve the biocorrosion resistance and biocompatibility of magnesium-based substrate, as well as to further incorporate the biofunctionality of localized drug delivery. The composite coating consisted of an inner layer of poly(lactic acid) served as drug carrier, and an outer layer composed of brushite generated through chemical solution deposition, where a facile pretreatment of UV irradiation was applied to the poly(lactic acid) coating to facilitate the heterogeneous nucleation of brushite. The in vitro degradation results of electrochemical measurements and immersion tests indicated a considerable reduction of magnesium degradation provided the composite coating. A systematic investigation of cellular response with cell viability, adhesion and ALP assays confirmed the coated Mg alloy induced no toxicity to MC3T3-E1 osteoblastic cells but rather fostered cell attachment 1

ACS Paragon Plus Environment


1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

and proliferation, and promoted osteogenic differentiation, revealing excellent biosafety and biocompatibility, and enhanced osteoinductive potential. In vitro drug release profile of paclitaxel from the composite coating was monitored with UV-Vis spectroscopy, showing alleviated initial burst release and a sustained and controlled release feature of the drug-loaded composite coating. These findings suggested that the bi-layer poly(lactic acid)/brushite coating provided effective protection for Mg alloy, greatly enhanced cytocompatibility and bioactivity, and moreover, possessed local drug delivery capability; hence magnesium alloy with poly(lactic acid)/brushite coating presents great potential in orthopedic clinical applications, especially for localized bone therapy. KEYWORDS: Composite coating; biodegradable magnesium alloy; orthopedic implants; biocorrosion resistance; biocompatibility; local drug delivery

INTRODUCTION Orthopedic implants have been routinely and widely applied for the fixation and correction of bone fractures as well as the replacement of bones and joints, to help recover impaired functions. In addition to providing mechanical stabilization, it is also desired for orthopedic devices to function in an appropriate manner biologically, of which osseointegration is essential to determining the success post implantation1. Apart from these fundamental requirements, more clinical demands, especially those related to bone regeneration and infection issues are emerging. For instance, for treating patients suffering from osteoporosis, bone cancer and other musculoskeletal diseases, as well as bacterial-associated infections and inflammations, it often requires the intervention of therapeutic agents. To address these challenges, the strategy of local drug delivery at the time of implantation has been developed2. Compared to the systemic administration, local delivery of drugs provides various possibilities not only for minimizing side effects and risk of overdose, but also to improve the bioavailability of drugs with appropriate concentration effectively reaching the target site, maintaining drug stability for a longer period and hence enhancing localized efficacy and efficiency3. The major challenge of the next generation orthopedic implants thus lies in presenting appropriate integration of mechanical support and (bio)chemical and morphological cues for bone cell attachment and proliferation to accelerate new bone formation, 2


Page 2 of 34

Page 3 of 34

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


meanwhile, serving as matrices for sustained delivery of therapeutic agents to elicit desired cellular responses, or/and treat diseases or infections locally4. Development of implant devices with local drug delivery capability, thereby, has aroused increasing attention from the research community during the past decade. In recent years, magnesium and its alloys has been proposed as a class of revolutionary implant material in orthopedic field to alleviate the current drawbacks of permanent metallic implants (e.g. stainless steels, titanium, and cobalt-chromium based alloys). Taking advantages of their unique biodegradability and appropriate mechanical properties, it is expected that patients could avoid a secondary surgery for implant removal upon the completion of the remodeling process, and the stress shielding effect to be mitigated or even eliminated as well5. Ideally, the degradation rate of an implant device into nontoxic products should match the time scale required for bone tissue regeneration to prevent from the early loss of mechanical stabilization ability. However, the excessive degradation of magnesium, especially associated with rapid evolution of H2 and overhigh pH increase, which could largely undermine the implant lifetime and osseointegration properties, has posed the foremost obstacle on its way towards clinical acceptance. In addition, currently there still exists a lack of effective drug delivery systems developed on Mg-based implant materials, which limits its use for treating infections or beating bone diseases6. In spite of a lot of recent research which provided various surface coating solutions focusing on addressing the former challenge while ignoring the latter, to meet the clinical requirements for magnesium-based orthopedic implants, all of these aspects should be evaluated in parallel during the development of novel coating strategies1. Recently, the state-of-art concept of multiple functionalities for implant coatings has been explored for improving implant performance, yet is in its incipient stage of development. In order to improve the corrosion resistance to enhance durability and biocompatibility of magnesium-based implant, and simultaneously incorporate the function of local drug release, appropriate multifunctional coating is necessary through the use of a variety of surface modification methods7. Specifically of interest coating materials fall into two specific categories-degradable polymers and bioceramics, especially calcium phosphate (CaP). Amongst degradable polymers, the thermoplastic aliphatic poly(esters), such as 3



1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

poly(lactic acid), poly(glycolic acid), and poly(lactic-co-glycolic acid) often serve as vectors to deliver drugs, genes and growth factors8 due to their excellent biocompatibility and biodegradability. Nevertheless, the intrinsic drawbacks of polymeric coatings, such as the acidic degradation product, corroding polymer particles, poor osteointegration and frequently causing non-infection inflammation6, 9 may cause potential risks of serious clinical side-effects and implications. Alternatively, calcium phosphate coatings appears to be an efficacious means to improve biocompatibility and osteointegration properties10, among which dicalcium phosphate dihydrate (Brushite, DCPD) reveals greater solubility and have been demonstrated by previous in vitro and in vivo studies11-12 to own excellent osteoconductivity and osteoinductivity. Additionally, CaP coatings were also reported to be employed as drug carrier through physisorption13, which, however, were often associated with the issue of loosely bonded drugs leading to undesirable drug delivery, such as excessively high initial burst release and short release duration. Consequently, expectations of implants to favor bone tissue regeneration in addition to controlled and sustained drug delivery could hardly be achieved with either class of the coating materials alone. On the other hand, inspired by the distinct advantages of polymers and calcium phosphates, integrated composite may present a superior approach14-15. Nevertheless, so far, to our knowledge, there has been no research exploring multifunctional composite coating with drug delivery capability on biodegradable magnesium and its alloys, especially for orthopedic applications. Herein, we reported a biodegradable bi-layer composite coating of poly(lactic acid)/brushite developed on a Mg-Nd-Zn-Zr alloy based on a rational design of multifunctional surface modifications. PLA coating, mainly served as the drug carrier, was prepared as the intermediate layer above the Mg substrate, while brushite coating, the outermost layer prepared by chemical solution deposition, was incorporated to ameliorate the biocorrosion resistance, biocompatibility and osteocompatibility of Mg substrate, and in addition, to assist in controlling the drug release as the top barrier layer. The critical issue regarding composite materials often lies in the interfacial bonding strength, where in this study a facile method of UV irradiation was employed to enhance the surface wettability of the polymer film and thus induce brushite deposition and improve the bonding strength. The in vitro degradation, cytocompatibility and osteogenic differentiation of the resulting coated magnesium substrate were 4


Page 4 of 34

Page 5 of 34

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


subjected to a systematic investigation. Furthermore, the local drug delivery capability was evaluated by monitoring the drug release kinetics with a model drug paclitaxel (PTX), known for its significant antineoplastic activity against various solid tumors. MATERIALS AND METHODS Materials Mg-Nd-Zn-Zr alloy (JDBM) was used as the substrate material, of which the chemical composition and the processing could be found in Reference16. Racemic poly(D,L-lactic acid) (hereafter denoted as PLA) with a weight-average molecular weight of ~100,000 was supplied by Ji’nan Dai gang Biomaterial Co., Ltd (China). Hydrofluoric acid (40 % wt), calcium phosphate monobasic monohydrate reagent, sodium hydroxide, ethanol and acetone of analytical grade were all purchased from Sinopharm Chemical Reagent Co., Ltd (China). Methanol and ethyl acetate of analytical grade were from Shanghai Ling Feng Chemical Reagent Co., Ltd (China). Paclitaxel (99.5%) was obtained from Xi’an Hao Xuan Biological Technology Co., Ltd. (China). Alpha-modified Eagle's medium (α-MEM, Gibco), trypsin/EDTA (Gbico), fetal bovine serum (FBS, Gibco) and Dulbecco phosphate buffered saline (D-PBS, HyClone), LIVE/DEAD Viability/Cytotoxicity Assay Kit (Calcein AM and ethidium homodimer (EthD-1), Molecular Probes) were all purchased from Thermo Fisher Scientific Inc. (USA). Cell Counting Kit-8 was obtained from Shanghai Beyotime Biological Technology Co., Ltd (China). The p-nitrophenyl phosphate (pNPP) and BCA Protein Assay Kit were all purchased from Sigma-Aldrich Co. LLC (USA). Sample preparation JDBM samples were cut from the extruded rod to Ф19mm×3mm for extract preparation, Ф14mm×3mm for cell adhesion assays and electrochemical tests, and Ф12mm×3mm for immersion tests. Samples were ground with SiC paper progressively up to 3000 grid, then ultrasonically cleaned sequentially in acetone and alcohol for 10 min, and finally dried with a stream of nitrogen. Subsequently, cleaned JDBM samples were pre-treated with hydrofluoric acid (40 % wt) for 12 h at room temperature (~25 ℃) to form a dense conversion MgF2 layer (~1.5 µm)17 on JDBM surfaces (denoted as HF-Mg). The HF-Mg samples were stored in a desiccator until further treatments. 5



1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

The preparation process of DCPD/PLA bi-layer coating was carried out as follows, with schematic diagram illustrated in Fig. 1. A thin polymer layer was first prepared on the disk samples (denoted as PLA-Mg) from PLA solution (in ethyl acetate at a ratio of 6.0 % w/v) with the use of a dip-coater (SYDC-100, Shanghai Yanyan materials technology co., ltd, China) followed by drying in a vacuum desiccator for 2 days to remove the residual solvent. Prior to the brushite deposition, PLA-coated samples were pre-treated with a UV irradiator (PSD Series Digital UV Ozone System, Novascan, USA) for an optimized time of 3-5 min, and then immediately vertically placed in a supersaturated solution of Ca(H2PO4)2·H2O with an initial pH of 4.0 for up to 12 h at room temperature to deposit a DCPD coating on top of PLA (denoted as DCPD/PLA-Mg). Samples were then rinsed with deionized water, and dried with a stream of nitrogen.

Fig. 1 Schematic diagram illustrating the preparation process of PLA/DCPD bi-layer coating on Mg-based substrate. (A) Mg pre-treated with hydrofluoric acid for 12 h to form MgF2 conversion layer; (B) A thin PLA layer prepared on Mg-based substrate with dip-coating method; (C) PLA-coated sample pre-treated by quick UV irradiation; (D) The nucleation of DCPD on PLA-coated sample after immersion in a supersaturated solution of Ca(H2PO4)2•H2O; (E) a dense DCPD coating generated on top of PLA-coated sample after 12 h immersion.

6


Page 6 of 34

Page 7 of 34

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Physicochemical characterizations of the coating Samples were characterized by X-ray diffraction (XRD, Smart Lab, Japan) to identify the outermost inorganic layer of the coating. The surface morphology and elemental composition of the samples were examined by scanning electron microscopy (SEM, JSM 7600F, Japan) coupled with energy-dispersive X-ray spectroscopy (EDS, OXFORD, UK). Prior to SEM observation, samples were coated with a layer of gold with thickness of ~20 nm by a sputter coater (SHINKKU VD MSP-1S, Japan). Two different liquids were respectively used to carry out the static contact angle measurements: ultrapure water and diiodomethane. The values of the polar and dispersive components of their surface tension are shown in Supplementary Table S1. Measurements of static sessile drop contact angle (CA) were carried out with a drop volume of approximately 4 µL by an optical contact angle system (JC2000D1, Shanghai Zhongchen Digital Technic Apparatus Co., Ltd, China) at room temperature. Results are presented as an average of six measurements on at least three different sample surfaces. From the contact angle results obtained with the two liquids, the surface free energy (SFE) could be calculated using the Owens-Wendt theoretical model. This model gives the long-range dispersion γ and the short-range polar components of surface free energy (SFE) according to the following equation18: 1 + cos = 2 / / + 2 / /

(1)

where γ is the SFE of the surface and γ the SFE of the liquid. Samples were subjected to the “Tape Test” according to the standard test method (ASTM D3359-97) to semiquantify the interfacial bonding strength of the composite coating. 3M scotch transparent tape (44 N/100 mm) was applied to firmly attach onto the cross-hatch pattern area on the test sample followed by a quick removal of the tape. The adhesion strength of the coating was thus determined by the area ratio of the defect/detached region to the whole cross-hatch pattern area, ranging from 0B to 5B. In vitro degradation The electrochemical corrosion behavior was measured using a Princeton EG&G PARSTAT 2273 with a three-electrode corrosion cell system. A saturated calomel electrode (SCE) was used as the reference electrode, a Pt electrode as the counter electrode and a sample as the working electrode. The sample was fixed at a specially designed device with an exposed area of 1 cm2. Electrochemical impedance 7



1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

Page 8 of 34

spectroscopy (EIS) was measured with an AC amplitude of 10 mV and a frequency range from 100 kHz to 100 mHz. The potentiodynamic polarization was performed with a scanning rate of 1 mV/s and a potential from −0.25 mV vs. OCP to −1.2 mV. The overall process was carried out in m-SBF buffer solution at 37 °C. All the impedance data was fitted and analyzed with ZSimpwin software (version 3.10). Samples were immersed in cell culture medium (α-MEM) supplemented with 10% fetal bovine serum and 1% penicillin& streptomycin at 37°C with 5% CO2 for 3 days to obtain the extracts, with the sample area to solution volume ratio of 1.25 cm2/mL according to ISO 10993─5. The concentrations of Mg2+ in the extracts were analyzed with an inductively coupled plasma atomic emission spectrometer (ICP-AES, iCAP6300, USA), and the pH was measured with a pH meter (FE20, Mettler Toledo, Switzerland). The m-SBF buffer solution was used for the hydrogen evolution test. Samples were immersed in m-SBF solution at 37 °C for a period of up to 21 days. The ratio of sample area to solution volume was 60 mL/cm2, according to ASTM G31-72. A test sample was placed below a funnel in a beaker filled with m-SBF solution, and a burette was mounted over the funnel to collect the hydrogen evolved from the sample. Hydrogen evolution represents magnesium corrosion rate. The volume of the hydrogen evolution can be converted to the weight loss (W) of the sample according to the reaction formulation as below: Mg + 2H2O→Mg2+ +2OH-+H2↑

(2)

The corrosion rate is calculated by the following equation according to ASTM G31-72: Corrosion rate= (K×W)/ (A×T×D)

(3)

Where the coefficient K=8.76×104, W is the weight loss (g), A represents the sample area exposed to solution (cm2), T the exposure time (h) and D the density of the material (g·cm-3). After the hydrogen evolution test, the samples post immersion for 21d were removed from the m-SBF solutions, gently rinsed with deionized water, dried with a stream of nitrogen and then subjected to SEM observation. In vitro biocompatibility and bioactivity Cell culture Osteoblastic cells MC3T3-E1 (Cell Bank, Chinese Academy of Sciences) were cultured in growth 8


Page 9 of 34

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


medium consisting of α-MEM supplemented with 10% FBS and 1% penicillin & streptomycin in a cell incubator at humidified atmosphere with 5% CO2 at 37 °C, with fresh medium replaced every 2 days. When reaching 80-90% confluence, cells were detached with 0.25% trypsin/EDTA, and then subcultured at a density of 1×105 cells/cm2 in 100 mm petri dishes. MC3T3-E1 cells of 3rd-6th passage were used for the experiments. The following cell assays were all conducted in triplicate for each group of samples at each time and repeated at least once. Cell viability and proliferation assay Sample extract was prepared with cell culture medium as previously according to ISO 10993-5. The extract was then collected for cell viability assays. The cell viability assay was carried out using indirect contact method. MC3T3-E1 cells were seeded in a 96-well plate at a density of 2×104 cells/mL with 100 µL of suspension for each well, and then incubated for 24 h to allow for attachment. Afterwards the culture medium was replaced with 100 µL of the sample extracts. After incubation for 1, 3 and 5 days, respectively, 10 µL of CCK8 solution was added to each well and incubated for 2 h in a cell incubator, and then the optical density (OD) was measured at 450 nm with a microplate reader (iMARK, Bio-Rad, USA). Direct cell adhesion assay Sterile test samples were placed in 24-well plates with MC3T3-E1 cells seeded on the samples at a density of 1×104 cells/well, and then incubated for 6 h, 1 day and 3 days, respectively. Subsequently, the cells were gently rinsed with D-PBS, and then stained with Calcein-AM and Ethidium homodimer-1 reagents (LIVE/DEAD Viability/Cytotoxicity Assay Kit) for 15 min at 37 °C. After gently rinsed twice with D-PBS, samples were mounted onto glass slides for fluorescence microscopy observation (IX71, Olympus, Japan). Alkaline phosphatase (ALP) expression MC3T3-E1 cells were seeded in a 24-well plate at a density of 1.4×104 cells/cm2. After 24 h of culture, the medium were replaced with the sample extracts, while the cells cultured in α-MEM medium as the blank control group. After 7, 14 and 21 days of incubation, ALP activity of MC3T3-E1 cells was 9



1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

Page 10 of 34

measured. Cells were gently rinsed three times with D-PBS and then lysed in 0.1% Triton X-100 by freezing and thawing for two cycles. The lysate was centrifuged, and the resultant supernatant was assayed for ALP activity using p-nitrophenyl phosphate (pNPP) as the substrate. The reaction lasted for 15 min at 37 °C and then was terminated with the use of NaOH solution (3 M). The amount of pNPP produced was measured at 415 nm with a microplate reader and the total protein content was examined using a BCA Protein Assay Kit. The measured OD values were normalized to the total protein content accordingly. In vitro drug release profile Samples loaded with a model drug of paclitaxel were immersed into 15 mL of phosphate-buffer saline containing 1.0 % v/v tween 20 (PBST, pH=7.4) which was added to increase the solubility of PTX. After immersed in a shaking incubator (120 rpm, 37 °C) for predetermined time intervals, 2 ml of medium was taken out for analysis and replaced with the same amount of fresh medium. The concentration of PTX in the release medium was measured using an UV/Vis spectrophotometer (INESA instrument UV759) at the wavelength of 228 nm. Statistical analysis All experimental data were expressed as the means ± standard deviations. Student’s t-test was used to determine the level of significance and p

Brushite Bilayer Coating on

Recommend Documents