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Multifunctional PLGA/Parylene C Coating for Implant Materials – an Integral Approach for Biointerface Optimization Monika Golda-Cepa, Aleksandra Chorylek, Paulina Chytrosz, Monika Brzychczy–Wloch, Joanna Jaworska, Janusz Kasperczyk, Minna Hakkarainen, Klas Engvall, and Andrzej Kotarba ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.6b08025 • Publication Date (Web): 08 Aug 2016 Downloaded from http://pubs.acs.org on August 9, 2016

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ACS Applied Materials & Interfaces

Multifunctional PLGA/Parylene C Coating for Implant Materials – an Integral Approach for Biointerface Optimization M. Golda–Cepa1, A. Chorylek1, P. Chytrosz1, M. Brzychczy–Wloch2, J. Jaworska3, J. Kasperczyk3, M. Hakkarainen4, K. Engvall5, A. Kotarba1* 1

Faculty of Chemistry, Jagiellonian University, Ingardena 3, 30-060 Krakow, Poland

2

Department of Bacteriology, Microbial Ecology and Parasitology, Jagiellonian University

Medical College, Czysta 18, 31-121 Krakow, Poland 3

Centre of Polymer and Carbon Materials, Polish Academy of Science, Curie Skłodowskiej 34,

41-819 Zabrze, Poland 4

Department of Fiber and Polymer Technology, KTH Royal Institute of Technology, SE-100 44

Stockholm, Sweden 5

Department of Chemical Engineering and Technology, KTH Royal Institute of Technology,

SE-100 44 Stockholm, Sweden Keywords: parylene C, coating, implant, biointerface, oxygen plasma, PLGA, MG–63, controlled drug elution

ACS Paragon Plus Environment

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Abstract Functionalizing implant surfaces is critical for improving their performance. An integrated approach was employed to develop a multifunctional implant coating based on oxygen plasmamodified parylene C and drug-loaded, biodegradable poly(L–lactide-co-glycolide) (PLGA). The key functional attributes of the coating (i.e., anti–corrosion, biocompatible, anti–infection and therapeutic) were thoroughly characterized at each fabrication step by spectroscopic, microscopic and biologic methods and at different scales, ranging from molecular, through the nano– and microscales to the macroscopic scale. The chemistry of each layer was demonstrated separately, and their mutual affinity was shown to be indispensable for the development of versatile coatings for implant applications.

1. Introduction Dynamic growth in the demand for advanced implant materials is occurring globally. This is caused by a number of factors: an aging population, a desire to maintain a certain quality of life, an increased number of traffic accidents, and civilizational progress. The demand particularly concerns metal implants, which effectively restore the role of damaged bones and enable patients to properly function in terms of everyday life. Available data from recent years indicate that in Europe, 186/100 000 inhabitants hip joint replacement surgeries and 126/100 000 inhabitants arthroscopic/knee repair surgeries are performed each year using metal implants1. Hip and knee replacements are considered the most effective interventions for severe osteoarthritis, as they reduce pain and disability and restore patients to near-normal function. Osteoarthritis is one of the ten most-disabling diseases in developed countries. Worldwide estimates are that 10 % of men and 18 % of women aged over 60 years have symptomatic osteoarthritis, including moderate and severe forms (WHO, 2014).


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The most popular alloys used in orthopedics are stainless steel (e.g. SS 316) and titanium alloys (e.g. Ti–6Al–4V)2, which contain heavy metals, even toxic metals. Because the conditions inside the body promote corrosion, implant surfaces undergo slow destruction, and heavy metal ions migrate into the body at levels of up to 10 µl/ml/week.3 Thus, corrosion resistance is a fundamental property of metal implants, as it determines the interaction between the metal surface and tissues. Therefore, the negative effects of metal ions released in the human body continue to be the subject of many studies.4,5,6 Implantation procedures are complicated and associated with risks of several side effects, such as severe inflammatory reactions, bone loss (induced by corrosion and wear), lack of tissue–implant integration (lack of biocompatibility) and infection, which can eventually cause systemic toxicity and/or implant rejection, in a worstcase scenario. One of the most promising solutions to prevent implant corrosion is the use of polymer-based, surface-protective coatings; for practical purposes, these coatings should be easy to deposit, strongly adhere to an implant surface, be chemically inert in physiological environments, be biocompatible, and have suitable mechanical strength. One polymeric material meeting these criteria is parylene C (poly(chloro–para–xylene)), which has been applied in numerous medical applications, e.g., as a neural prosthesis coating and in cardiac devices.7,8 Parylene C is a crystalline–amorphous composite, has a non–porous structure, and has a low permeability to small molecules.9,10 While the chemical passivity of a parylene C coating determines its applicability as a protective, anti–corrosive layer, it is also problematic because hydrophobic surfaces do not support the growth of adhesive cells, e.g., osteoblasts. A parylene C surface can be successfully modified using oxygen plasma modification methods by introducing oxygen– containing functional groups and surface nanotopography.11 Rough surfaces with groups such as


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–OH and –COOH have been described to be important for initial osteogenesis, which induces wound healing and consequently osseointegration.12,13,14 It is important to emphasize that immediately after implantation, a so–called ‘race for the surface’ starts, i.e., cell integration competes with bacterial adhesion to a surface of a medical device.15 This is of particular concern because microbial colonization of an implant surface causes serious biomaterial–associated infections. Therefore, there is a strong need for designing biomaterial surfaces that enable osteoblasts to win the race for the surface and at the same time, remain inert to bacteria, delaying the process of biofilm formation. Aside from better biocompatibility, oxygen–containing functional groups on a parylene C surface provide additional binding sites for bioactive substances or for the anchorage of a biodegradable, drug–loaded polymer layer, which can improve the coating with an additional therapeutic function. In cases of postoperative complications, patients receive a variety of oral and intramuscular anti–inflammatory and antimicrobial drugs. For this reason, solutions relying on controlled, local drug release from the surface of an inserted implant have been investigated; these solutions are very beneficial for patients.16 The most important benefits include the application of lower drug doses and limiting their activity to the target tissue, which reduces the risk of side effects associated with the oral administration of high doses of medication.17,18 Major problems related to drug release from an implant surface is engineering its surface and controlling the amount and rate of drug release. Therefore, there is an ongoing need to improve the surface of implants to achieve controlled drug release at appropriate therapeutic levels for sufficiently long durations. Biodegradable polymers, e.g. poly(L–lactide-co-glycolide) (PLGA), are successfully used for the entrapment and controlled release of drugs.19 PLGA absorption time can be controlled by using various molar ratios of individual comonomers. Depositing a drug–


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loaded PLGA film on a non–degradable material ensures a steady drug release due to the gradual process of hydrolytic degradation.20 Moreover, drug release depends on the chain microstructure of the polymer used as the matrix of the drug delivery system.21 One challenge of engineering metal implants introducing specific properties to optimize the metal implant–tissue interface. A short list includes four key functional attributes: an ideal implant surface would be anti–corrosive, biocompatible, anti–infection microbial and therapeutic. The introduction of each function requires addressing specific practical problems. Fig. 1 shows the functions and main challenges together with strategic solutions. The clear goal is to retain the key physical properties of a coating while modifying only the outermost surface to influence the biointeraction and obtain a therapeutic function. If such surface modification is properly performed, the mechanical properties and functionalities of the protective layer will be unaffected. At the same time, the bioresponse related to the tissue–device interface will be improved or modulated, augmenting the implant with therapeutic function. Such a task requires a multidisciplinary approach spanning over several research fields, such as surface chemistry, materials science, and biology, as well as the application of a wide variety of bulk and surfacesensitive experimental techniques, such as spectroscopy, microscopy, microbiological tests, and in vitro tests. The proposed integral approach for biointerface optimization via better control of implant surface properties and drug release kinetics aimed at developing novel, more stable, biocompatible implant coating with therapeutic function. We take the advantage of the synergistic effect of physical (nanotopography) and chemical (surface groups) modification of parylene C by oxygen plasma to design and prepare a multilayer coating for novel, functional implants. Here, we aim to integrate these fundamental research results and the gained insights


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into a general concept of multifunctional polymer coatings based on PLGA+drug/parylene C. To our best knowledge such comprehensive investigations are proposed for the first time.

Figure 1. The scheme illustrating the main required functional attributes of a polymeric coating for a metal implant; these determine the main challenges and strategies for coating design and development (see also Fig. S1 in Supporting Information).

2. Materials and Methods 2.1. Parylene C Coating Parylene C (poly(chloro–para–xylylene)) films and coatings were prepared by a CVD technique (ParaTech Coating Scandinavia AB). Samples (2x2 cm) of medical-grade, rolled, polished SS 316L with bright annealed (BA) surface finishing were used as substrates.3 The thickness of the parylene C layer used in the experiments was 8 µm; before each experiment, parylene C foil was washed in 2–propanol (Avantor) and air-dried. 2.2. Oxygen Plasma Treatment To generate oxygen–containing functional groups and nanotopography, parylene C samples were modified using oxygen plasma (FEMTO system, Diener Electronics) at a controlled oxygen partial pressure of 0.2 mbar and with 50 W of plasma generator power and a sample exposure time of 8 min.


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2.3. Therapeutic Layer Composition For the therapeutic layer preparation, the biodegradable copolymer PLGA (85/15) was used (Centre of Polymer and Carbon Materials PAS). Copolymers were obtained by bulk ringopening polymerization using Zr(acac)4 as a nontoxic initiator, according to a previously described procedure.22 The ratio of D,L-lactidyl units and glycolidyl units in PGLA was 85:15. PLGA was precipitated in methanol and dried. Then, PLGA was dissolved in CH2Cl2 and one of the following active substances was added: ibuprofen ((±)–2–(4–isobutylphenyl)propanoic acid) (Sigma–Aldrich), diclofenac sodium salt (2–[(2,6–dichlorophenyl)amino]benzeneacetic acid sodium salt) (Sigma–Aldrich) or gentamicin sulfate salt (Biological Industries). For each polymeric disc (2 cm2), 300 µl of the drug+PLGA solution (0.13 %) was deposited. The concentration for each drug was adjusted with reference to the specific drug-loading efficiencies that were determined. The solutions were prepared at concentrations to achieve the theoretical drug loads of 1 mg/cm2 and 0.5 mg/cm2 for ibuprofen and diclofenac, respectively. In the case of gentamicin, the theoretical drug load was 0.04 mg/cm2. The drug-loading efficiencies (LEs) were calculated using the following formula:

= ∙ 100 %

(Eq. 1)

The actual drug loading was determined after the elution tests by summing the total amount of the eluted drug. To confirm that the degradation and elution from the systems were complete, FTIR measurements were performed afterward, revealing the extent of drug release and total PLGA degradation. 2.4. Airbrush Deposition of Therapeutic Layer


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Drug–loaded PLGA was deposited on oxygen plasma-treated parylene C using the airbrush method (Fengda, BD–320), with a fixed distance of 8 cm from the samples, a pressure of 4 bars and a volume 300 µm. Drug elution experiments for the PLGA+drug/parylene C layers were performed for 35 days in phosphate–buffered saline (Lonza). 2.5. Material Characterization Part of the parylene C characterization results have already been published, including the XRD, TG–DTA, XPS, AFM, contact angle and EIS results; therefore, the details of these experiments are not repeated here.9,11,23 However, for illustrating a complete picture of the developed system, they will be referred to in the discussion section. 2.5.1. Scanning Electron Microscopy (SEM) Images of biological moieties (MG–63 osteoblast–like cells and S. aureus DSM 24167, P. aeruginosa ATCC 27853 bacteria strains) on parylene C, as well as images of the parylene C and PLGA+drug/parylene C layer surfaces, were taken using a Hitachi S–4700 scanning electron microscope. All of the samples were coated with Au prior the observations. Biological samples were additionally fixed before SEM observation according to the following procedure: the investigated samples were fixed with 3 % glutaraldehyde (Sigma–Aldrich) in PBS (Lonza) for 24 h, washed in PBS three times for 10 min and incubated for 4 h with 1 % OsO4 (Avantor) in PBS. Subsequently, the samples were washed three times for 10 min in PBS and dehydrated in a series of ethanol (Sigma–Aldrich) concentrations, 50, 60, 70, 80, 90, 96, and 100 %, for 10 min each. Then, the samples were incubated twice for 15 min in absolute ethanol and dried twice for 30 s in hexamethyldisilazane (Sigma–Aldrich). 2.5.2. Fourier Transform Infrared (FTIR) Microscopy


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FTIR imaging analyses of the polymeric films were performed in reflectance mode on a Spectrum Spotlight 400 FTIR microscope connected to a Spectrum 100 FTIR spectrometer (Perkin-Elmer, Inc.). The images were taken at a resolution of 8 cm−1 between 4000 and 700 cm−1 with 16 scans per pixel. 2.5.3. Electrochemical Impedance Spectroscopy (EIS) EIS was performed using a typical three–electrode set–up consisting of a sample work electrode, a Pt mesh counter electrode and a KCl–saturated Ag/AgCl reference electrode according to the procedure reported in our previous study.24 The measurements were repeated after 1, 3, 7 and 9 days to monitor changes of the coated sample, i.e., at the stainless steel – coating interface. The impedance spectra of the samples were analyzed with ZView software. 2.5.4. Atomic Absorption Spectrometry (AAS) To assess the total number of metal ions released from the SS 316L, AAS analyses were performed according to the procedure reported in our previous study.3 The concentrations of Fe, Ni and Cr ions were determined with the use of Perkin–Elmer Model 3110 (with graphite furnace HGA–600) and Perkin–Elmer 4100 ZL systems. 2.5.5. Laser Desorption Ionization–Mass Spectrometry (LDI–MS) The surface analyses of the investigated drug molecules and PLGA+drug/parylene C coatings were performed using a TOF–MS (Bruker UltraFlex) with a SCOUT–MTP ion source (Bruker Daltonics, Bremen, Germany) and a 337 nm nitrogen laser. The spectrum obtained for each sample is an accumulation of 1000 shots. To determine the drug molecules in the PLGA+drug/parylene C systems, 10 µl of a 1 mg/ml solution of 2,5–dihydroxybenzoic acid (Fluka) was dropped on the surfaces and left to dry in the air. 2.5.6. Nuclear Magnetic Resonance Spectroscopy (NMR)


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Changes of the comonomeric unit composition in the PGLA layers during degradation were monitored with 1H NMR using a Bruker-Avance II Ultrashield Plus spectrometer operating at 600 MHz (Karlsruhe, Germany). Spectra were obtained in d-CDCl3 with 32 scans, a 11 ls pulse width, and a 2.65 s acquisition time. 2.5.7. In vitro Drug Release Studies Drug release studies of the therapeutic layers were performed in PBS (Lonza). The prepared samples of PLGA+drug/parylene C were placed in 4 ml of PBS and transferred into an orbital shaker–incubator (Biosan, ES–20/60) set at 130 rpm and 37 °C. All of the dissolution medium was collected at predetermined time intervals for analysis, and an equal volume of fresh buffer was added immediately. 2.5.8. Ultraviolet–Visible Spectroscopy (UV–Vis) Ibuprofen and diclofenac absorb UV light when dissolved in PBS, which enables analysis without pre–treatment. The collected samples were examined using a Cary 60 UV–Vis Spectrophotometer (Aglient) to quantify the amount of eluted drug. Absorbance was measured at 272 nm and 276 nm for ibuprofen and diclofenac, respectively. The corresponding drug concentrations were then found from calibration curves. For the UV–Vis detection of gentamicin sulfate, additional pre–treatment with the use of AgNPs was necessary. The method of gentamicin–AgNP detection is based on decreasing absorbance of the AgNP spectrum at a characteristic wavelength with increasing drug concentration.25 The collected samples were incubated for 2 h with freshly synthesized AgNPs. Absorbance was measured at 389 nm, and the corresponding drug concentration was then found from the calibration curve. 2.5.9. Drug Elution Kinetics


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The data obtained from the in vitro release studies were fitted to two semi–empirical kinetic models: first–order (Eq. 2) and Korsmeyer–Peppas26 models (Eq. 3): = ∙ 1– – ∙

where,

(Eq. 2)

ft – fraction of drug released over time, ftmax – maximum fraction of drug released during the process, k1 – the first-order kinetic constant [h–1] and = ∙

where,

(Eq. 3)

ft – fraction of drug released over time, a – constant incorporating structural and dosage form, and n – the release exponent indicating the drug release mechanism. For the Korsmeyer–Peppas model, the first 65 % of the drug release data were fitted to determine the mechanism of drug release. 2.6. Biological Tests 2.6.1. Fluorescent Staining of MG–63 Cells Cell culture tests of the PLGA+drug/parylene C system were performed using the MG–63 cell line (human osteosarcoma, ATCC 86051601) grown in DMEM (Lonza) supplemented with glutamine and FBS (Biowest). A live/dead assay was used to evaluate the viability of cells on the tested surfaces. To evaluate cell viability on the parylene C with therapeutic coatings, cells were incubated for 24 h and 72 h with the samples and stained using fluorescein diacetate (0.008 mg/ml in PBS) and propidium iodide (0.02 mg/ml in PBS) dyes (Sigma–Aldrich). The adhesion protein vinculin was used as a marker of focal contact. The cytoskeletal and focal adhesion analysis were performed after 24 h of culture, following a previously reported procedure27, with some modifications. Cells were fixed with 4 % paraformaldehyde in PBS


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(Lonza) (20 min) followed by washing with PBS. Then, the cells were permeabilized with 0.1 % Triton X–100 (5 min). Nonspecific binding sites were blocked by incubating the samples in 1 % bovine serum albumin in DPBS (30 min). Then, double staining was performed on each sample. Cells were first stained with monoclonal antibody (anti–vinculin) (FAK100, Chemicon) (1 h). Next, the cells were labeled with fluorescein (FITC)–anti–mouse IgG antibody (AP124F, Chemicon) in combination with TRITC–conjugated phalloidin (FAK100, Chemicon) (30 min). All steps were carried out at room temperature. The samples were mounted on microscope slides under glass coverslips using ProLong® Diamond Antifade Mountant with DAPI (Molecular Probes) for photo–bleaching reduction. The stained cells were imaged using a fluorescent microscope IX51 (Olympus) and then processed using ImageJ 1.48v software28 to establish the ratio of live/dead cells. The focal contact area (µm2/cell) was determined from the images of fluorescently stained cells, according to a recently described procedure.29 To ensure robust representative statistics for each evaluated surface, at least 50 cells were counted. The MTT tests were performed on parylene C samples with and without a drug+PLGA therapeutic layer to establish whether the oxygen plasma treatment or introduced active substances exhibited any toxic effects on MG–63 cells. Cells were seeded in 24–well plates at a density of 1×105 cells/well and incubated for 24 h, 48 h (data not shown) and 72 h with the samples. Then, the standard MTT assay protocol was applied. Absorbance was measured at 565 nm using an Infinite 200M PRO NanoQuant spectrophotometer (Tecan). 2.6.2. Microbiological Tests The microbial strains were incubated in 50 ml of Bacto™ Tryptic Soy Broth (TSB) (Becton Dickinson) for 16 h at 37 °C. The static adhesion assay method was applied to assess bacterial adhesion to the parylene C films. The experiments were performed in 24–well tissue culture


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plates; each film was placed in a well and incubated for 4 h in 1 ml of a specific bacterial strain suspension in TSB (~5 × 108 CFU/ml) (CFU: colony forming unit). Each experimental incubation was performed three times. After the incubation, the parylene films were carefully washed several times with PBS to remove non-attached bacteria from the surface. The samples were observed via SEM (Section 2.5.1). 2.7. Statistical Analysis All of the experiments were performed in at least three independent series, including three parallel experiments for each type of sample. One–way ANOVA with Tukey’s post hoc multiple comparison test were performed to determine any statistically significant differences in the biological experiments. All of the statistical calculations were carried out using the StatSoft tool Statistica 10.0. 3. Results and Discussion 3.1. Coating Morphology The key functions of the designed coatings are the anti–corrosive, biocompatible, anti– infection and therapeutic attributes (Fig. 1). The coating morphology evolved as these functions were subsequently added to the designed system. A microscopic view of the general changes at three pivotal stages of the coating fabrication is illustrated by the SEM observations (Fig. 2A–C). Fig. 2A shows the top view of the smooth surface of a CVD–fabricated parylene C coating (1st stage); also shown are images of an oxygen plasma-roughened surface (2nd stage, Fig. 2B) and the final system, consisting of PLGA+ibuprofen/parylene C (3rd stage, Fig. 2C). Because the changes in surface topography affect interactions with osteoblasts and bacteria, they are discussed in the subsections devoted to the biocompatible and anti–infection functions. The last


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image (Fig. 2C) presents the surface used for the drug release studies, which is discussed in the therapeutic subsection.


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Figure 2. Scanning electron microscopy images of the coating surface topography at different stages of development: unmodified parylene C (A), oxygen plasma-modified parylene C (B) and the PLGA+ibuprofen/parylene C system (C).

3.2. Anti–corrosion Functionality Upon prolonged contact of a metal implant surface with body fluids (i.e., conditions with elevated temperatures and saline), surface corrosion phenomena take place, resulting in a high rate of locally released corrosion products. Many metal ions in implant alloys (e.g., Ni, Cr, Fe, V, Mo, Co, Al) are highly hazardous to human health. Most of them are toxic and are causative factors in several diseases, e.g., cancer, dermatitis (Ni, Cr),30 liver failure, oxidative damage to lipid membranes, (Fe)31, neurotoxicity (V, Al),32,33 cardiomyopathy (Co),34 growth retardation and infertility (Mo).35 Most investigations on the coatings used for medical devices focus on increasing corrosion resistance. Electrochemical methods, such as EIS, are used for monitoring the processes taking place on metal implant surfaces in artificial body fluids. The representative parameter, which quantifies the resistivity of an implant to corrosion, is total impedance. As shown in Fig. 3A–B, for uncoated metal implants made of typical alloys, such as SS 316L and Ti–6Al–4V, the impedance values in Hank’s solution are in the range of 105–106 Ω cm2, while it reaches 109 Ω cm2 for the parylene C-coated samples.36,37 The thickness of the parylene C coating was optimized and the optimum thickness was found to be 8 µm which is a compromise


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between anti–corrosion protection and excellent wear resistance properties. Additionally, it was demonstrated that such coating thickness provides sufficient elastomer properties, essential to sustain the implantation surgery strains and micromotions during long-term usage in the body. Because the oxygen plasma modification of the parylene C coating is limited to the outermost surface and does not exceed the nanometer range,11 the impedance of the system remains the same within the experimental limit (Fig. 3A). Additionally, the anti–corrosion properties of parylene C have also been well documented in long–term in vitro experiments in Simulated Body Fluid with an H2O2 presence, simulating the inflammatory response.36,37 The anti– corrosion properties of the coating are reflected in the total level of metal ion release (Fig. 3B). For the uncoated and passivated samples of SS 316L and Ti–6Al–4V, the total metal release reached the dangerous levels of 92 µg/ml/28 days and 62 µg/ml/28 days, respectively. In contrast, for the parylene C-coated samples, a dramatically lower release of 0.1 µg/ml/28 days was observed. Such a strong effect clearly illustrates the potential for the application of this polymeric coating as a long–term protective layer on metal implants.

Figure 3. The total value of corrosion resistivity gauged by the total values of impedance (A) and metal ion release over 28 days (B) for parylene C-coated SS 316L in comparison with uncoated surfaces. All of the data were collected during at least 3 independent experiments.


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3.3. Biocompatibility Function The biological response to biomaterials is largely controlled by their surface chemistry and structure. Parylene C is described as a bioinert, non–toxic polymer,38,39 with advantages such as high impedance and low permeability for water, resulting in good anti–corrosion properties, as described in section 3.1. Therefore, it is a suitable biomaterial for the insulation of cardiac and neural interface devices and metal implants.40,41,37 To be biocompatible, which is indicated in this study by cell adhesion, the originally hydrophobic parylene C surface has to become hydrophilic. Changes in the parylene C surface chemical composition (e.g., oxygen–containing functional groups) and morphology (at the micro- and nanoscale) were achieved by an oxygen plasma treatment. The optimal oxygen plasma modification for best coating biocompatibility were established in our previous studies.9 The correlation between biocompatibility and oxygen plasma parameters such as oxygen partial pressure, time and generator power was established, revealing that the preferred conditions of oxygen plasma treatment of parylene C, stimulating adhesion of osteoblast cells, are t = 5 – 8 min, P = 50 W, pO2 = 0.2 mbar. The differences in cells adhesion to unmodified and 8 min plasma modified samples were quantified by crystal violet staining and optical density measurements. Here, for better insight into MG–63 cell–parylene C interactions, additional results of the fluorescent staining of focal contacts (Fig. 4) and SEM observations of cellular morphology (Fig. 5) are presented.


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Figure 4. Immunofluorescence labeling of vinculin (green), actin (red), and nuclei (blue) in osteoblast–like MG–63 cells on unmodified (A) and oxygen plasma-modified parylene C (B) together with the quantified focal contact area (C) based on the measurements of at least 150 cells (3 independent series × 50 cells). The red dotted line represents the average results for the reference well of a tissue culture plate (TCP).

Focal contacts (FC) are closed junctions where the distance between the substrate surface and the cell membrane is between 10–20 nm.42 In general, cells with a low motility form strong focal adhesions, while motile cells form less adhesive structures. Osteoblastic cells are especially sensitive to nanocues through filopodia production, which enhances the necessary adhesion required for successful osseointegration.43 The focal adhesion staining results are presented in Fig. 4A–B, where images of fluorescently stained MG–63 cells are presented with a plot of quantified focal contact area (Fig. 4C). The osteoblast–like cells were not spread and displayed limited or no focal contacts on the unmodified parylene C. On the oxygen plasma-modified samples, the cells were well spread with bipolar or tripolar morphology, as indicated by the parallel orientation of actin filaments with the long cell axis. The observed focal contacts were numerous and distributed around the cell peripheries (i.e., extremities of F–actin filaments). These results indicate that after 24 h of incubation, the adhesion phase occurred only on the oxygen plasma-modified parylene C surfaces and in the control wells (TCP).


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Figure 5. SEM images of MG–63 cells on unmodified (A–C) and oxygen plasma-modified parylene C (D–F). White rectangles indicate the magnified area of the cell peripheries.

The area of focal contacts was quantified for the oxygen plasma-treated samples and compared with that of the unmodified parylene C samples, where the FC level was minor or below the detection limit. These results were also compared to the control well (tissue culture plate, TCP), which did not have any polymeric films (the red, dotted line in Fig. 4C). As summarized in the plot, the average area of focal contacts on the oxygen plasma-modified parylene C and the TCP control sample was 6.21 ± 1.2 µm2 and 6.53 ± 1.6 µm2, respectively. No statistically significant differences were detected. Therefore, the results demonstrate that the surface is suitable for cell adhesion, as the wells of TCPs are designed for optimal cell adhesion using a similar surface modification method to create a hydrophilic and negatively charged surface. More detailed information on cell adhesion to the investigated surfaces can be obtained from the SEM observations. In Fig. 5A–F, SEM images of the MG–63 osteoblast–like cells on parylene C coating represent cell morphology on the investigated surfaces. The representative


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images of cell on unmodified parylene C (Fig. 5A–C) and on oxygen plasma-modified parylene C (Fig. 5D–F) are presented with corresponding magnifications (white rectangles) of some regions of the image. Cells cultured on unmodified parylene C exhibited reduced spreading and suppressed filopodia formation (Fig. 5C). In contrast, cells cultured on oxygen plasmamodified parylene C exhibited increased proliferation,44 a largely elongated morphology and surface adaptation via wide lamellipodia and dendritic filopodia (Fig. 5E–F). The effect of the surface modification is reflected in the filopodial length of the MG–63 cells. For pristine parylene C, the average length was 1.1±0.3 µm; however, it dramatically increased to 5.4±0.8 µm for the oxygen plasma-modified surfaces. The differences in cell morphology on the unmodified and oxygen plasma-modified parylene C arise from the ability of nanotopography and surface oxygen–containing functional groups to enhance osteoblast adhesion. As previously reported, surface chemistry and roughness strongly modulate cell attachment.45,46,47 The atomic force microscopy (AFM) measurements revealed the formation of nanocues in the range of 60– 200 nm, resulting from amorphous domains being etched in the parylene C. Such modifications have been found to be beneficial for osteoblasts9,23 while hindering bacterial attachment. Moreover, functional groups such as –OH and –COOH attract intracellular proteins responsible for focal adhesion (e.g., α–actinin, paxillin, talin, and vinculin).48 To investigate the biomimetic processes, which in the context of biomaterials for bone applications, are defined as a bone–like apatite layer formed on a substrate after immersion in simulated body fluids, calcium phosphate (CaP) formation on oxygen plasma-treated parylene C was monitored using XRD (see Fig. S2 in Supporting Information). The results showed that CaP formed after 7 days on oxygen plasma-treated parylene C, with a characteristic maximum at 2Θ=31.4° assigned to the monoclinic (space group P21/c) (511) crystallographic plane and


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crystallites size of ~30 nm. After 14 days of immersion, in addition to the maximum at 2Θ=31.4°, another maximum was observed at 2Θ=28.5°, and this was attributed to (122) or (221) crystallographic planes (ICSD# 060117).49 The size of the CaP crystallites after 14 days was ~290 nm. The bone–like apatite effectively formed on the oxygen plasma-treated parylene C surfaces where oxygen–containing groups provided nucleation sites for CaP crystallization. The maxima at 2Θ=13.7°, which was observed for all of the investigated samples, was due to the crystalline part of parylene C and can be attributed to the (020) crystallographic plane (space group P1).50 The results presented herein enforce the role of the rough substrate and oxygen–containing functional groups on the surface in promoting osteoblastic cell adhesion. They also provide valuable guidelines for designing material surfaces that are required for the development of an appropriate osteogenic surface for osteoblastic anchorage and maturation. The anti–corrosion and biocompatible coating obtained at this stage can be additionally equipped with anti–infection and therapeutic functions, which will be described in the following subsections. 3.4. Anti–infection Function Surfaces can be called anti–bacterial when the attachment of bacteria is hampered due to their physical or chemical morphologies or compositions.51 Despite exterior biomaterial properties, bacterial adhesion to surfaces is strongly affected by the surface properties and the bacterial cell wall. Due to the cell wall, bacteria are significantly less deformable than are eukaryotic cells, such as osteoblasts. Therefore, bacterial sensitivity to surface topography is limited to irregularities smaller than their dimensions, i.e., typically less than 1 µm. Indeed, various surface topographies (e.g., unmodified and oxygen plasma-treated parylene C) result in dramatically different bacterial surface colonization (Fig. 6).


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Figure 6. SEM images of S. aureus (A, C) and P. aeruginosa (B, D) bacteria on unmodified, smooth (A–B) and oxygen plasma-modified, nanorough parylene C (C–D). These images show the different natures of bacterial adhesion (i.e., biofilm vs single-cell attachment).

Oxygen plasma treatment generates nanoscale topography, effectively limiting the surface area available to bacteria, as presented in Fig. 6C–D. In this study, to achieve a comprehensive picture, gram–positive (thicker cell wall) and gram–negative (thinner cell wall) bacterial strains were used. As shown in Fig. 6A–B, the surface of unmodified parylene C is covered with biofilms of S. aureus (Fig. 6A) as well as P. aeruginosa (Fig. 6B), while for the oxygen plasmamodified parylene C, single cells that are not agglomerated are dominant. Our previous study revealed that the number of bacteria adhering to the surfaces of unmodified and oxygen plasmamodified parylene C were similar, within experimental error; approximately 9×104 CFU/cm2 S. aureus adhered to unmodified parylene C, which was not significantly different the 8×105 CFU/cm2 that adhered to modified parylene C. While P. aeruginosa adhered in higher numbers than S. aureus, no significant differences were found between the unmodified and oxygen plasma-modified parylene C values, which were 6×106 CFU/cm2 and 4×106 CFU/cm2, respectively.23


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The SEM observations revealed that although the number of adherent cells can be similar between investigated samples, the nature of bacterial adhesion can vary. Generally, bacteria with hydrophobic properties preferably adhere to hydrophobic material surfaces, while those with hydrophilic characteristics prefer hydrophilic surfaces.52,53 In scientific reports, S. aureus and P. aeruginosa are described as hydrophobic.54,55 In Fig. 6A–B, bacteria in the form of a biofilm on can be observed on unmodified parylene C. Because the incubation time with the bacterial strains was the same for the investigated samples, it can be concluded that early–stage biofilm formation is significantly faster on unmodified parylene C than on oxygen plasma-treated samples (Fig. 6C–D). The results of these investigations are especially important from a clinical perspective, as delaying the onset of biofilm formation has significant benefits, particularly for implantable biomaterials; most antibiotics cannot penetrate through biofilms to effectively kill bacteria and therefore become ineffective. The last function to be introduced to the polymeric coating consists of a controlled drug delivery system. Currently, this strategy is one of the most rapidly advancing areas in biointerface development, and it offers numerous advantages over conventional dosage forms, including local treatment, effectiveness, reduced toxicity, and improved patient convenience. 3.5. Therapeutic Function At this stage, introducing the therapeutic function, we focus on controlling the drug release kinetics. This area of research has grown and diversified rapidly in recent years, with an emphasis on understanding the controlled release mechanism together with a selection of biodegradable polymers, which can be applied as a barrier for the prolonged release of bioactive molecules.56 All controlled release systems aim to improve the effectiveness of drug therapy. This improvement can take the form of increasing therapeutic activity compared to the intensity


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of side effects and reducing the number of doses required during treatment. PLGA is one of the best–defined biomaterials concerning the design and performance of a drug release system.56,57 To achieve the desired functionality in this co-polyester, the formation of a polymer network architecture is required, including netpoints and elastic segments. The resulting microstructure provides suitable options for adjusting the hydrolytic degradation rate, which depends on the glycolide content and overall molecular weight. The success of a drug–loaded PLGA biodegradable layer depends on several factors, such as interactions between drug molecules and the PLGA matrix, drug dispersion homogeneity within the layer, and control of the drug release kinetics in both short- and long–term elution within the drug–specific therapeutic window. Each of these important issues will be addressed below. The first task was to investigate whether the drug loading was efficacious and whether the molecules remained intact upon interaction with the PLGA matrix. For this purpose, the direct laser desorption/ionization mass spectrometry analytic method58 was utilized to detect the drug molecules used in this study (i.e., ibuprofen, diclofenac, and gentamicin). The representative MS spectra, obtained from the clean PLGA/parylene C surface, the pure drugs molecules and the drug molecules loaded in the PLGA/parylene C surface, are presented in Fig. 7A–G, together with corresponding molecular structures of the drugs. The results revealed that in all of the investigated therapeutic layers, the drug molecules remained intact, preserving the molecular structure and thus, the therapeutic function. In Fig. 7A–C, the LDI–MS spectra of the clean PLGA/parylene C control surface, standard ibuprofen and PLGA+ibuprofen/parylene C surface, are illustrated. For the PLGA/parylene C reference samples (Fig. 7A), fragmentation of the matrix can be observed at the lower m/z region (200–300 u.). Nevertheless, the ibuprofen molecules are still detectable (Fig. 7B–C). In the spectra, the diagnostic m/z signal of ibuprofen


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can be distinguished and identified as ibuprofen molecules with sodium adduct ions [M+Na+] at 229.402 m/z (ibuprofen, exact mass= 206.130 u). In Fig. 7D–E, the spectra of diclofenac are presented, first for the pure standard drug molecule (Fig. 7D) and then for the drug present in PLGA+diclofenac/parylene C (Fig. 7E). In both spectra, the characteristic peak at 339.811 m/z for diclofenac (exact mass= 316.998 u) with sodium adduct ions [M+Na+] was detected, indicating the presence of intact drug molecules in the PLGA layer. Additional confirmation of diclofenac stability is provided by the characteristic splitting of the main line (339.811 m/z) of 2 m/z (341.810 m/z) due to the presence of the natural abundance of 35Cl and 37Cl isotopes. In Fig. 7F–G, the mass spectra of gentamicin samples are shown. The gentamicin B1 molecules (exact mass= 496.274 u) for the drug alone (Fig. 7F), as well as PLGA+gentamicin/parylene C systems, are identified at 496.436 m/z [M–H+] with the characteristic fragmentation profile. The loss of –CH3 end groups is due to fragmentation, caused by direct desorption/ionization, which can promote ion fragmentation, when higher laser powers are required to desorb the analyte.59


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Figure 7. LDI-MS spectra of the PLGA/parylene C layer (A), ibuprofen (B), diclofenac (D) and gentamicin (F) molecules and corresponding PLGA+drug/parylene C surfaces (C, E, G),


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showing the physical nature of the drug adsorption (intact molecules) in the therapeutic layer. Additional maxima for ibuprofen (B–C) at 228.318 and 227.402 m/z were identified as fragmented ibuprofen molecules without –H and 2×–H groups with sodium adduct ions, respectively. Additional maxima for gentamicin (C) at 482.512 and 468.589 m/z were identified as gentamicin without –CH3 and 2×–CH3 end groups, respectively.

After the confirmation of successful drug introduction into the multi–layered coating, the key aspect was to evaluate the uniformity of the drug dispersion in the biodegradable layer. To assess the homogeneity of the drug distribution, µFTIR mapping was performed. The results are presented in Fig. 8A–E, which shows representative average FTIR spectra in the range of 700– 1800 cm–1 for the investigated samples. First, the characteristic bands were assigned to each drug in a different region, at 1713 cm–1 for ibuprofen (group C=O60,61), 745 cm–1 for diclofenac (group C–Cl62) and 1037 cm–1 for gentamicin (group C–N and C–O63). Based on these assignments, the corresponding absorbance maps (20 µm × 20 µm) were collected at the characteristic selected wavelength of each drug (left–hand side of Fig. 8). As a reference, the results for parylene C and PLGA/parylene C without drugs were used. Because the characteristic wavelengths of the drug molecules do not overlap with that of the parylene C support and PLGA matrix, in each case, the absorbance level for the reference samples was lower than 0.2. Here, representative results are presented for the ibuprofen characteristic wavelength at 1713 cm–1 (Fig. 8A–B). An apparent increase in the absorbance can be observed for drug–containing samples when they are compared to the references. The minor differences in absorbance are ≈ 0.1 a.u. (FTIR images Fig. 8C–E), indicating similar homogeneous distributions for all of the investigated drug–loaded therapeutic layers at the micrometer scale.


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Figure 8. FTIR spectra (right) and corresponding maps (left) of oxygen plasma-modified parylene

C

(A),

PLGA/parylene

C

(B),

PLGA+ibuprofen/parylene

C

(C),

PLGA+diclofenac/parylene C (D), PLGA+gentamicin/parylene C (E). The FTIR maps were collected at the characteristic absorption bands of the drugs (C–E, selected wavelength for each drug is marked with a dotted line).


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Table 1. Characteristics of PLGA+drug/parylene C systems (drug-loading efficiency, sample drug load) and their performance (total time of drug elution).

drug ibuprofen diclofenac gentamicin

loading efficiency [%] 14 25 15

average sample drug load [µg/cm2] 180 ± 2 140 ± 3 1.50 ± 0.02

time of total drug elution [h] 192 h 12 h 504 h

The final test for the drug–loaded polymer coating included in vitro drug elution experiments. The samples used in the study are listed in Table 1, where drug-loading efficiency, sample drug load, and the total time of drug release are summarized. The drug elution results are summarized in Fig. 9A–C, showing remarkably different release profiles for diclofenac (Fig. 9A), ibuprofen (Fig. 9B) and gentamicin (Fig. 9C) for several time intervals (0–21 days). Each data point in the elution profile represents the mean of three independent concentration determinations. To assess whether the drug release and degradation of PLGA+drug layers were completed, FTIR imaging was performed after the release experiments. All the spectra were similar to the one presented in Fig. 8A for oxygen plasma modified parylene C. thus, it can be concluded that the prepared coating of parylene C remain in the unaffected form after total release of the loaded drug and degradation of PLGA.


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Figure 9. Kinetic curves of the diclofenac (A), ibuprofen (B) and gentamicin (C) fractions eluted over time from the PLGA+drug/parylene C systems.

When designing drug delivery systems, it is important to remember that drug release also depends on the chain microstructure of the polymer used as the matrix for a drug. We chose poly(lactide-co-glycolide) with a specific chain microstructure, which resulted from the synthesis conditions (i.e., appropriate time, temperature, and initiator of the ring-opening polymerization). The ratio of D,L-lactidyl units and glycolidyl units in PGLA was 85:15. An NMR analysis showed a gradual loss of units, suggesting a uniform degradation of the layers (Fig. S3 in Supporting Information). The fastest kinetics were observed for diclofenac release (Fig. 9A), for which 96 % of the drug was released during the first 3 h. The burst release of diclofenac was caused by the dissolution of the drug at or close to the surface of the PLGA coating and was facilitated by its swelling.20 The probable reason for the observed release profile is the relatively poor solubility of diclofenac in the CH2Cl2 needed to dissolve the PLGA matrix. In Fig. 9B, an intermediate profile of ibuprofen release is shown. In this case, during the first 3 h and at the end of 12 h, 46 % and 72 % of the drug was released, respectively. This system provided a slower and more sustained release for up to 8 days. The slowest release profile was observed for gentamicin (Fig. 9C), where two


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processes can be distinguished. During the first period, a minor burst release of gentamicin was observed (12 h; 79 %), followed by a slow, continuous release up to the 21 days. The obtained release profiles revealed that ibuprofen and gentamicin formulations are more suitable for prolonged elution. None of the investigated systems failed to release the drug and, as clearly indicated in Fig. 9A–C, none of the formulations followed zero–order release kinetics. As the drug dispersions, revealed by µFTIR, were quite similar, the differences in release kinetics can be associated mainly with the properties of the drug molecules, such as chemical composition, interaction with PLGA and water solubility.

Table 2. Comparison of first-order and Korsmeyer–Peppas kinetic model fittings of the ibuprofen and gentamicin release data for the PLGA+drug/parylene C systems. drug ibuprofen gentamicin

first order R2 k 0.9546 0.1481 0.4026 1.2748

Korsmeyer-Peppas R2 k n 0.9685 18.705 0.6289 0.9974 64.8423 0.0831

The drug release data of ibuprofen and gentamicin can be described by first-order and Korsmeyer–Peppas kinetic models.64 Because a burst release of diclofenac was observed, resulting in 70 % elution during the first hour of the experiments, none of the mentioned models are applicable to these systems. The model-fitting parameters (depending on the model used: R2, K and/or n values) are summarized in Table 2. Analyzing the experimental data in the context of first-order and Korsmeyer–Peppas equations leads to a better understanding of the mechanism involved in the degradation kinetics of PLGA. The adequacy of the model for the description of experimental data can be assessed by the correlation coefficient (R2). First-order kinetics represent pharmaceutical dosage forms that contain water–soluble drugs in porous matrices. The


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drug is released in such a way that it is proportional to the amount of the drug remaining in the matrix interior. As a consequence, the amount of drug released by a unit of time diminished.65 The system containing ibuprofen fit remarkably well with the first-order model (R2>0.95) and even better with the Korsmeyer–Peppas model (R2>0.96), with n>0.5 (0.6289). These results indicate that the drug elution is governed by both dispersion and diffusion with a non–Fickian transport mechanism but that the latter dominates (better fit). The gentamicin–loaded system fit well only with the Korsmeyer–Peppas model (R2>0.96), with n

Parylene C Coating for ... - ACS Publications

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