Since late 1990s, polyetheretherketone (PEEK) has presented a promising polymeric alternative to metal implant components, particularly in orthopedic and traumatic applications. However, PEEK is biologically inert, which has constrained its potential applications. In this manner, enhancing the bioactivity of PEEK is a huge challenge that must be comprehended to completely understand the potential advantages. Up to now, two noteworthy methodologies are discussed to enhance the bioactivity of PEEK, including bulk and surface modification. Although the latter is extremely challenging due to the very high physical and chemical stability of the high performance polymer, there are some stated modification reactions in the literature, which will be collocated with in the literature-reported PEEK composites in the present article. We will furthermore add information on polymer-based drug delivery systems and the biofunctionalization of polymers in general and discuss their applicability for PEEK, as we estimate that these strategies will gain greater attention in the future. At the end of the article, our own research on the development of a PEEK-associated biodegradable drug-delivery system with potential application in dentistry or orthopedics will be highlighted.
Polyetheretherketone (PEEK) is a semi-crystalline polycyclic aromatic thermoplastic that was initially created by a group of English researchers in 1978 . In the 1980s, PEEK was popularized for modern applications, for example, airplanes and turbine edges . By the late 1990s, PEEK turned into a promising polymeric alternative to metal implant components, particularly in orthopedic and for traumatic applications . The development of carbon fiber reinforced PEEK (CFR-PEEK) opened up new perspectives for the application of this novel composite material for more mechanically stressed implant components such as femoral prostheses in manufactured hip joints . In the course of recent years, PEEK and its composites have furthermore garnered much enthusiasm from dental technologists and dentists. Besides aesthetics, the fundamental main thrust is given by PEEK’s incredible biomechanical properties. In its natural form, the Young’s modulus of PEEK is around 3.6 GPa, while the Young’s modulus of CFR-PEEK is around 18 GPas  which is near that of cortical bone , . Subsequently, it has been proposed that PEEK could display less stress-shielding effects when compared to conventionally applied dental and orthopedic implant materials such as titanium, demonstrating a much higher Young’s modulus of 116 GPa . Moreover, PEEK shows great biocompatibility in vitro and in vivo, causing neither toxic or mutagenic effects nor clinically significant inflammation , , , , , . However, PEEK is biologically inert , , which has constrained its potential applications. In this manner, enhancing the bioactivity of PEEK is a huge challenge that must be comprehended to completely understand the potential advantages. Up to now, two noteworthy methodologies have been discussed to enhance the bioactivity of PEEK, including bulk and surface modification. Although the latter is extremely challenging due to the very high physical and chemical stability of the high performance polymer, there are some stated modification reactions in the literature, which will be collocated with in the literature-reported PEEK composites in the present article. We will furthermore add information on polymer-based drug delivery systems and the biofunctionalization of polymers in general and discuss their applicability for PEEK, as we estimate that these strategies will gain greater attention in the future. At the end of the article, our own research on the development of a PEEK-associated biodegradable drug-delivery system with potential application in dentistry or orthopedics will be highlighted.
With a focus on biomedical application, one can classify PEEK composites into three main categories: (i) conventional composites, (ii) bioactive composites and (iii) drug-release systems. The formation of conventional composites aims at the alteration of biomechanical properties adapted to the application site, while the generation of bioactive PEEK composites has evolved as an attractive strategy to improve the bioactivity in particular the osseointegrative property of PEEK-based implant components. Table 1 summarizes reported conventional and bioactive PEEK composites, their processing conditions and main properties for biomedical application. The third category of drug delivery is rather novel for PEEK but well established for softer polymers as, for example, poly(L-lactide) (PLLA) and can be understood as a subcategory of bioactive composites, where the bioactive substance is released from the bulk material in a controlled manner.
|Filler||Processing conditions||Interesting properties for biomedical application|
|PTW||Compression molding||Increased lifetime, tooth-like color |
|SiO2||Hot pressing and injection molding||Improved anti-inflammatory |
|AIN||Hot pressing||Higher storage modulus |
|HA||Selective laser sinteringIn situ synthetic process||Enhanced cell proliferation and osteogenic differentiation  Bioactivity and improved mechanical stability , |
|β-TCP||Injection molding||No confirmed bioactivity|
Most convenient composites of PEEK are fiber-reinforced materials. The choice of the fiber (material, geometry, orientation) depends on the later application prospect. Typical materials are carbon- and glassfibers. Those fibers are used to increase the stiffness of PEEK. It is also possible to combine carbon nano fibers (CNF) with PEEK to simultaneously increase the strength and generate conductivity, especially interesting for electrical devices. Another possibility to generate conductivity is to add carbon nano tubes (CNT). CNTs also improve mechanical properties and can be used with PEEK for aeronautic applications. Besides these additives there are some other filler, which are commercially as potassium titanate whiskers (PTW). This composite possesses much better properties than neat PEEK. The strength and stiffness are increased and the crystallization-temperature is decreased. Under water-lubricated conditions the friction coefficient is lower compared to neat PEEK and the wear resistance is increased. These improvements lead to composites that are more appropriate under water-lubricated conditions compared to neat PEEK . This could be a huge advantage for dental applications, because the lifetime of an implant will be increased. Also it is possible to lighten up the neat PEEK to get a more natural look of an implant, like PEEK combined with TiO2. Another example are composites with crystalline-silica (SiO2). These are used as electronic packaging substrates. SiO2 gives some important advantages for the composite. The thermal stability and the char yield are increased with SiO2. Furthermore, the heterogenous crystallization of the matrix is increased . Such a composite could be used for dental applications. The positive effect is, that SiO2 reacts as an anti-inflammatory. It is also possible to produce PEEK composites with ceramics, for example, aluminum nitride (AIN). The adhesion between these two components is very good and the strength is increased. This composite could be also used for applications in electronic packaging. For those applications it is a good property to have a high storage modulus. With the PEEK composite with AIN this modulus is significant increased and can be a possibility for electronic packaging . This composite can be used for orthopedic applications if the storage modulus of neat PEEK is not high enough. The good adhesion between both materials is necessary to build a useable implant.
Besides conventional PEEK composites there is the group of bioactive PEEK-composites. Those composites are used to get a better integration of implants and to improve the interface between bones and PEEK-composites .
One example for such a composite is the combination of PEEK with hydroxyapatite (HA) particles. The reason why HA is used is that the chemistry is similar to the chemistry of bones. HA is a bioceramic and has a good binding capacity to bones. Such an improvement of bioactivity is a huge advantage for medical applications, but the challenge is the compatible mixing with PEEK. Compared to pure PEEK the mechanical properties (tensile strength and strain to fracture) are lower and implants would break earlier than bones . To maintain the mechanical properties of PEEK and improve the bioactivity it is possible to use HA nanoparticles . The most promising way for particle embedding with regard to maintenance of mechanical stability of PEEK seems to be an in situ synthetic process, where the polymer building reaction is performed in presence of the particles at the site of application , . Another example of a bioactive PEEK-composite is the combination with tricalcium phosphate (TCP). The current problem of using β-TCP is, that the improvement of bioactivity is not confirmed. At the moment the proliferation of osteoblasts is not improved compared to pure PEEK, and there is no concentration-dependent decrease of proliferation. But there are some reports, stating that the shift of the pH value of the cell culture medium, that emerges from the degradation of PEEK-β-TCP-composites , might be an influencing factor on bioactivity. This requires further fundamental research.
The application of soft synthetic biodegradable polymers, e.g. poly(lactic acid) (PLA), poly(glycolic acid) PGA, poly(lactide-co-glycolide) (PLGA) or poly(ε-caprolactone) (PCL) , whose main benefit can be found in the complete metabolisation by the human organism prevails in the research and development of drug-release systems.
In general, drug-delivery systems can be divided into several mechanisms of release, mostly as diffusion- and chemically controlled. Diffusion-controlled drug release is an example of a physical mechanism that is based on the diffusion of a drug molecule through a certain coating or matrix, mostly of a polymeric nature . It can be formulated in two ways: On the one hand, by using a matrix system, where the drug is directly integrated within the carrier, or on the other hand, by utilizing a reservoir system – also called as core-shell system – that consists of an inert membrane protecting the active substances from the surroundings . The characteristics of the polymer matrix material significantly defines the diffusion and permeation rate of contained active components. The release coefficient of an active agent is, for example, lowered by the raise of molecular weight, degree of crosslinking and stiffness of the polymeric carrier as well as the interaction between the polymer and active agent molecules. An increase of the release coefficient is manageable by the addition of flexibilizers or filling materials.
The second main category of chemically controlled drug-release systems operate by polymer chain degradation, the rupture of chemically bound drugs or biomolecules or swelling of hydrophilic polymer carriers. Among these hydrogels, some have garnered special attention, because of their stimulus-responsive properties, where swelling characteristics and thereby drug release kinetics are defined by variation in pH, temperature, or the electrical field of their surroundings . For PEEK, such systems have not been described so far, probably due to the rather harsh processing conditions, being high melting temperature and solubility only in strong acids as HF and H2SO4, which can be hardly withstood by bioactive reagents. However, the establishment of a drug-delivery system on PEEK via the deposition of a polymeric drug-release coating is conceivable.
Except for an association with PEEK, there is a wide range of local drug delivery systems already used in orthopedics and surgery. An interesting example is the Septocoll E-fleece (Biomet, Berlin, Germany) . It is characterized by a metabolizing a collagen matrix which contains gentamicin as a preventive local antibiotic protection. The fleece is impregnated with two different derivatives of gentamicin in order to achieve short- and long-term drug release: a high initial drug delivery is caused by gentamicin sulfate which is easily soluble in water, while gentamicin crobefat is less soluble in water and therefore continues to be released over a longer period of time of about 10 days. In dentistry, local drug delivery systems are less established. One known example is the PerioChip®(Dexcel® Pharma GmbH, Alzenau, Germany), which is inserted directly into the periodontal pocket. The system persists of a gelatine matrix, which biodegrades completely and is equipped with the antibacterial chlorhexidine gluconate. The system is characterized by a high initial drug release of 40% in the first 24 h, while drug release is completed in the ensuing 7–10 days .
Different types of surface modifications have been reported for the improvement of PEEK regarding improved bioactivity. In contrast to the previously discussed bulk modification, surface modification aims to alter the surface of PEEK with little or no effect on the core. To date, three principal processes are discussed to modify the surface of PEEK for application in dentistry and orthopedics: (i) plasma treatment, (ii) chemical surface modification and (iii) surface coating. Biofunctionalization, so far rarely discussed for PEEK, will be added as a fourth subchapter in this main chapter of surface modifications, as the high impact for polymers in general is estimated to spill over the alteration for the bioactivity of PEEK.
Plasma describes an ionized gas mixture in which highly reactive radicals are formed. These can cause different reactions with the substrate surface. There are different methods for plasma generation, which differ in energy supply and pressure conditions. For example: there are low temperature and high temperature plasmas, as well as low-, high- and atmospheric-pressure plasmas . As the cold low-pressure plasma offers many advantages for the treatment of polymers, especially PEEK, such as the low demand for the process gas, the lack of thermal stress on the polymer and the ability to handle also complex geometries , this technology is the most described in the literature , . In this technique, the process gas is introduced into an evacuated chamber and from that a plasma is generated. The power supply in this case may take place via high voltage, microwaves or by laser. Furthermore, it is possible to generate a pulsating plasma to prevent instabilities. The main objective of the plasma treatment is usually an activation or functionalization of the polymer surface which is expressed in a better wettability and a higher surface energy. For that free bonds and/or functional groups must be added to the surface, which can interact witch potential coatings or liquids. If a polymer is directly exposed to a plasma two mechanisms occur: The high-energy UV radiation, generated in the plasma, breaks the chains of the macromolecules so that free bonds are formed , and further impurities and contamination are removed . At the same time the charged radicals attack the surface of the polymer. Due to their high energy they are capable to break both, C-H and C-C bonds, which results in an additional large number of free bonds on the polymer surface . If the process gas used to generate the plasma is an inert gas like argon, there is no reaction between the process gas and free binding site. These are thus available for further reactions, one speaks of the activation of the polymer surface. Of course, these bonds are occupied as soon as possible, for example, by a reaction with atmospheric components or other media where the polymer is exposed directly after the plasma treatment. Furthermore, polymer chains start to rotate, so that functional groups are no longer pointing out from the surface, so that the activation is not stable over time ,  When using a reactive gas in the plasma, the resulting free bonds are immediately occupied by functional groups from the process gas, as summarized in Table 2.
|Plasma process gas||Introduced functional group|
|CF4, SF6, XeF2, NF3, BF3, SOF2, SiF4||C-F|
|NH3, N2H4, N2/H2||C-NH2|
|H2O, H2/O2, O2||C-OH|
|H2O, H2/O2, CO2/H2||COOH|
|H2S, H2/S8, CS2||C-SH|
Plasma treatment of PEEK has been demonstrated with various modifications  in order to obtain significant changes in the surface properties. On the one hand, the hydrophilic character of the material is changed by the addition of polar groups which is expressed in a much better wettability and can be quantified by an increase in surface energy , . This effect has been, for example, described for O2-plasma treated PEEK by Inagaki et al. . The authors were able to reduce the water-contact angle from nearly 90° to 67° in a short direct treatment of about 30 s. By this treatment the first settlement of the implant can be facilitated. On the other hand, the inserted functional groups can react with coatings or further substances . Furthermore, the cell growth is directly influenced by applied functional groups such as Schröder et al. reported for NH3-plasma treated PEEK . This can significantly help to ensure colonization of the body’s cells on the implant and thus the acceptance of the body to the implant.
Besides breakage of bonds inside macromolecules, the plasma generated UV radiation leads to oxidative degradation of the polymer, contributing to a highly accelerated physical aging . As these effects occur uniformly and permanently, extremely thin layers with a very constant etch rate will be removed. Amorphous regions are etched faster compared to cross-linked structures. The result is a nanoroughness on the polymer surface, which increases wettability and possibly the adhesive strength of coatings , , which might simplify the application of bioactive layers. However, undesired modification of bulk properties could be the consequence. A further plasma-based treatment is the deposition of ultra-thin layers on surfaces by plasma polymerization. One can distinguish between two different approaches: First, monomers are directly applied to the surface and the plasma provides the initiation energy for free radical polymerization of the monomer, while the resulting layer thickness depends on the previously applied amount of monomer. Second, monomers are added to or completely replace the process gas. During plasma application, monomer radicals are formed which react with the surface of the polymer and with each other . Thus, a densely crosslinked, regular layer forms, whose thickness can be well controlled by concentration of the monomer and process time. This mechanism is not only applicable to polymers, but also for the deposition of metallic or non-metallic layers . The combination of different layers can lead to improved adhesive strength of bioactive coatings, which could increase the durability of the implant.
Although the chemical surface modification of PEEK is extremely challenging due to the very high physical and chemical stability of the high performance polymer, there are some stated modifications in the literature. The most convenient for biomedical application are probably two purely wet-chemical amination procedures (Figure 1). The first approach by Noiset et al. is a three-step process , , starting with the reduction of the keto groups of the benzophenone repeating units of the PEEK polymer backbone by sodium borohydride (NaBH4) in order to generate hydroxyl functions. The formed PEEK-OH is a key intermediate for the covalent anchorage of reactive molecules with various kinds of chemical functions. Noiset et al. describe the attachment of hexamethylene diisocyanate (HMDI) under formation of PEEK-NCO, which was then further hydrolyzed by immersion into 0.5 M aqueous sodium hydroxide (NaOH) with dioxane (1:1) for 5 h in order to yield PEEK-NH2. The water contact angle on PEEK-NH2 is only slightly lower than on PEEK-OH, but it serves as a base for the covalent immobilization of the highly cell-adhesive protein fibronectin, thus a prerequisite for covalent biofunctionalization. A further amination was recently published by Becker et al. . They present a facile technique for the wet-chemical amination of PEEK achieved by the formation of a Schiff base using ethylene diamine (EDA). The purpose was the subsequent conjugation of the cell-adhesive RGD-peptide. Results show significantly enhanced cell attachment on RGD-modified surfaces compared to untreated PEEK. One further often stated chemical treatment of PEEK is the sulfonation by immersion into concentrated sulfuric acid. However, this is not a pure surface treatment. Much more PEEK is sulfonated and thereby dissolved in sulfuric acid. The resulting SPEEK was reported to induce pre-osteoblast functions including initial cell adhesion, proliferation and osteogenic differentiation in vitro as well as substantially enhanced osseointegration and bone-implant bonding strength in vivo and apatite-forming ability . Thus this modification, if bulk properties could be sufficiently maintained, might be nervertheless promising for application in dentistry and orthopedics.
The deposition of a thin layer of a bioactive material applied as a surface coating on implants presents a further modification process to improve the bioactivity of surfaces. Often this modification is combined with a previous plasma or chemical treatment to augment the bonding properties. In particular, titanium and hydroxyapatite coatings have been shown to be significantly useful for the enhancement of osseointegration. Various techniques are available for the application of bioactive coatings on PEEK: (i) spray-coating, (ii) dip-coating, (iii) spin-coating, (iv) aerosol-coating and (v) physical vapor deposition. (i) Spray-coating can be divided into cold (CS) and plasma sprays (PS). During CS powder particles are accelerated towards a substrate. Thereby, the particles’ kinetic energy is used for the coating process and the thermal influence onto the substrate is minimized. This technique is used for different metallic coatings, like copper, nickel or titanium . Research reports show that it is also possible to use CS to develop a hydroxyapatite coating on PEEK, which is able to rise the bone-to-implant contact ratio in vitro and in vivo . PS is a type of thermal spraying, which uses plasma energy to deposit metallic and non-metallic materials on a substrate. Therefore, the coating material is heated by the plasma until a molten or semi-molten state is reached. While cooling the material, a dense layer on the substrate is formed. A further modification of PS, the vacuum plasma spraying, uses low pressure and low temperatures to ionize a small region on the substrate . Using this method, PEEK was modified with a titanium layer which directly increases the bioactivity or the strength of a subsequently applied HA coating. These implants showed better bone-implant contact in vivo compared to untreated PEEK. (ii) Dip-coating is probably the simplest methods for the establishment if surface coatings. Samples are simply dipped into the liquid coating material (dissolved or molten state) until the desired coating thickness is achieved. It is mostly used for simple geometries . Good results on PEEK were described with titanium , nickel-phosphorus  and silanes . (iii) During spin-coating, the coating material is applied in liquid form on the substrate and this is then made to rotate. Thus, the liquid coating material is evenly spread on the substrate, while excess material is spun down, so that the film thickness can be controlled by speed and process time. This method has been used for coating of PEEK with HA and yielded promising results . (iv) Aerosol-coating is a powder spray method, where the substrate to be coated is exposed to an aerosol containing fine particles of the coating material. These particles are swirled by the process gas in an aerosol chamber and hit against the substrate surface with high energy. Depending on the coating material a further thermal treatment is needed. This process can be also used for HA-coatings, but the results show a low crystallinity which leads to lower biological cellular response than HA with high crystallinity. At the same time, however, the adhesion strength compared to plasma sprayed HA-coatings could be more than doubled (from 7.5 MPa to 15.5 MPa) . (v) Physical vapor deposition combines numerous methods, such as plasma (PD), electron beam (EBD) and pulsed laser deposition (PLD). In these methods, the coating materials are either ionized (PD) or directly evaporated (PLD, EBD). All three coating methods yield ultrathin layers of the coating material on the substrate with a homogeneous thickness , . The e-beam deposition of titanium on PEEK, for example, improved in vitro cellular responses, cell attachment, proliferation and osteoblastic differentiation , . The PLD of 1D-alumina nanostructures also seems to be able to improve the biocompatibility of PEEK . The different presented coating methods, in particular spray-coating and dip-coating, could be furthermore used to deposit a polymer carrier containing a bioactive agent in order to establish substrate-associated drug delivery systems. This has been described for several polymers onto substrates of all material classes. To the best of our knowledge, no PEEK-associated drug-delivery systems have been reported so far.
Among the numerous reported methods for immobilizing biomolecules as proteins, peptides, polysaccharides, etc. to surfaces physical adsorption via van der Waals or electrostatic interactions, physical entrapment within hydrogels, ligand-receptor pairing (as in biotin-avidin) and covalent attachment via cleavable or non-cleavable bonds are most commonly used. Dependent on the appliedimmobilization protocol a short-term and long-term localization of the biomolecule on the implant surface can be achieved . Thus, biofunctionalization can be either used to provide the implant with a drug delivery or a stable novel surface functionality.
Due to the low abundance of terminal functional groups on most polymers as PEEK, most biofunctionalization methods require a previous surface modification by plasma or chemical treatment, which might be combined with the grafting of an intermediary spacer between the surface and the biomolecule for multiplication of available functional groups on the surface. Here, multifunctional spacers such as poly(ethylenimine) or dendrimers with a wide range of terminal functional groups in a defined quantity offer a way to increase this surface functionality . In addition, maintenance of biological activity is very important for hydrophobic polymer surfaces as these are often associated with non specific adsorption and denaturation. Besides the integration of the hydrophilic PEG spacer , which effectively shields the compound from these issues, we investigated in one of our own studies, the inclusion of the peptide spacer (GGAP)4 on PLLA surfaces. In this way we reached an efficient prevention of non-specific protein adsorption and enhanced bioactivity of subsequently covalently attached anti-CD34 antibodies . Moreover, a prerequisite for successful biomolecule immobilization is of course that the specific functionality imparted to the inert surface must be compatible with the reactive sites on the compound to be covalently attached to that surface. While for adsorption processes the introduction of charged groups or simple modification of hydrophilicity should be enough, covalent immobilization affords defined functional groups, including thiols, aldehydes, carboxylic acids, hydroxyls and primary amines. These can be directly used for biomolecule attachment. However, most cases apply additional cross-linking agents, which can link the bioactive compound directly to the functionalized substrate (zero-length cross-linkers), or themselves introduce a spacer of several angstroms. Moreover, they can be stable or introduce hydrolyzable bonds in order to allow for biomolecule delivery from the surface. A detailed description of these agents and the associated chemistries can be found in Bioconjugate Techniques, by Hermanson . Biofunctionalization for PEEK has been described to aminate surfaces by covalent attachment of RGD or fibronectin in order to improve cellular attachment , . For application in dentistry and orthopedics specific addressing of osteoblasts in terms of proliferation is conceivable by biofunctionalization. While current PEEK modification for the promotion of osseointegration foresee the surface coating or bulk modification with calcium phosphate or hydroxylapatit, a further possibility could be the long- or short-term immobilization of bone morphogenetic protein 2 (BMP-2). BMP-2 has strong osteoinductive properties and is therefore often used for bone regeneration and repair. However, biofunctionalization methods need to be adapted due to a short short half-life and rapid degradation of BMP-2 .
Our vision: PEEK-associated biodegradable drug-release system
The functionalization of PEEK for biomedical application is also one of our research foci. One vision is the establishment of a PEEK-associated biodegradable drug delivery system with possible applications as a material for antibacterial dental implant components. As model substrate PEEK films (VESTAKEEP 4000G, thickness approx. 0.5 µm) were kindly provided by Evonik Industries AG. Furthermore, we used PLLA (Resomer 210, Evonik, Marl) and chlorhexidine acetate (CHX, Sigma Aldrich) as a model biodegradable polymeric carrier and as a model drug, which is well established as an antiseptic in dentistry, respectively. In the following, we detail the required steps for the establishment of such a PEEK-associated biodegradable drug delivery system including chemical PEEK surface treatment and coating with PLLA.
With the purpose of improving the wettability of PEEK for subsequent deposition of a PLLA coating, we investigated the two above described amination procedures using HMDI  and EDA . Briefly, PEEK film (8×2.5 cm) amination according to Noiset et al. involved three reaction steps: (i) reduction with 0.5 g NaBH4 in 250 mL dimethyl sulfoxide (DMSO) for 3 h at 120 °C, (ii) modification with 12.5 mL HMDI in 250 mL dried toluene and 0.025 g triethylenediamine at room temperature for 3 days, (iii) hydrolysis of terminal cyanate groups with 125 mL NaOH (0.5 M) in 125 mL dioxane for 5 h at room-temperature. In contrast, PEEK amination, according to Becker et al., only involved one step being 3 h reaction with ethylenediamine (EDA) at 120 °C. After both amination procedures, PEEK films were dried in vacuum at 60 °C overnight.
The wettability of modified PEEK films was characterized by means of water contact angle measurements using the sessile drop technique (Dataphysics Contact Angle System OCA, Filderstadt). Figure 2 illustrates a considerable decreased contact angle on aminated PEEK surfaces according to Becker et al. compared to unmodified and aminated PEEK films according to Noiset et al., which show comparable water contact angles. This can be probably dedicated to the presence of the longer alkyl chain of the coupled HMDI compared to the shorter alkyl chain of EDA. On both films amine surface loads of 0.3 pmol*cm−2 could be detected via the colorimetric sulfo-SDTB assay according to the manufacturers’ instructions .
Subsequently, both aminated and untreated PEEK surfaces were coated with PLLA via a dip-coating procedure. Therefore, PEEK samples were dipped 5 times into a polymer solution (0.5% PLLA in chloroform). Between those dips the drying time is 10 min.
The adhesive strength of the resulting biodegradable coatings was analyzed by means of a grid scratch test. The impact of the modifications is illustrated in Figure 3 at the example of untreated and EDA-treated PEEK. During the scratch test, a defined grid is cut into the coating and an adhesive tape is pressed to the surface and pulled off. Results showed better adhesive strength of the PLLA coating to aminated PEEK compared to unmodified PEEK. While most parts of the coating were pulled off from unmodified PEEK surfaces, the coating on aminated surfaces hardly shows any defects after elimination of the adhesive tape. PLLA-coated PEEK surfaces treated with HMDI have a similar appearance as EDA-treated PEEK after the grid scratch test. Thus adhesive strength could be enhanced by the presence of surface amine groups independently of the hydrophilicity of the surfaces.
The appearance of the coating was furthermore investigated via scanning electron microscopy (ZEISS Auriga, Jena). The surface is homogenous. The coating thickness measures about 5 µm. As you can see in Figure 4A the overlap was intentionally generated to measure the thickness of the coating, which is shown in Figure 4B.
Summarizing, results show successful deposition of a homogenous and adhering PLLA coating with promising properties for application as a carrier for a drug delivery system. The next steps are the in vitro exploration of in vitro CHX release and biological activity. However, one can already state, that established methods build the basis for further PEEK-associated drug delivery systems including different polymer carriers as, for example, biopolyesters and various bioactive agents having a broad range of application prospects.
The authors thank Mrs. Schmidt, Mrs. Lümkemann and Mrs. Haberland for expert technical assistance in contact angle measurements, Mrs. Mey for expert technical assistance in scanning electron microscopy and Mrs. Terveen for expert technical assistance in laboratory work and amination.
Conflict of interest: Authors state no conflict of interest.
Material and Methods:
Informed consent: Informed consent has been obtained from all individuals included in this study.
Ethical approval: The research related to human use has been complied with all the relevant national regulations, institutional policies and in accordance the tenets of the Helsinki Declaration, and has been approved by the authors’ inistitutional review board or equivalent committee.
2. Margolis. Engineering thermoplastics: properties and applications. Miami: CRC Press, 1985.Search in Google Scholar
4. Williams D. Polyetheretherketone for long-term implantable devices. Med Device Technol. 2008;19:10–1.Search in Google Scholar
6. Godara A, Raabe D, Green S. The influence of sterilization processes on the micromechanical properties of carbon fiber-reinforced PEEK composites for bone implant applications. Acta Biomater. 2007;3:209–20.10.1016/j.actbio.2006.11.005Search in Google Scholar PubMed
7. Wenz LM, Merritt K, Brown SA, Moet A, Steffee AD. In vitro biocompatibility of polyetheretherketone and polysulfone composites. J Biomed Mater Res. 1990;24:207–15.10.1002/jbm.820240207Search in Google Scholar PubMed
8. Nieminen T, Kallela I, Wuolijoki E, Kainulainen H, Hiidenheimo I, Rantala I. Amorphous and crystalline polyetheretherketone: Mechanical properties and tissue reactions during a 3-year follow-up. J Biomed Mater Res A. 2008;84:377–83.10.1002/jbm.a.31310Search in Google Scholar PubMed
9. von Wilmowsky C, Vairaktaris E, Pohle D, Rechtenwald T, Lutz R, Münstedt H, et al. Effects of bioactive glass and beta-TCP containing three-dimensional laser sintered polyetheretherketone composites on osteoblasts in vitro. J Biomed Mater Res A. 2008;87:896–902.10.1002/jbm.a.31822Search in Google Scholar PubMed
10. Rossi F, Perale G, Masi M. Controlled drug delivery systems, Springer Briefs in applied sciences and technology. Cham, Switzerland: Springer International Publishing, 2016.10.1007/978-3-319-02288-8Search in Google Scholar
11. Sternberg K, Petersen S, Grabow N, Senz V, Meyer zu Schwabedissen H, Kroemer HK, et al. Implant-associated local drug delivery systems based on biodegradable polymers: customized designs for different medical applications. Biomed Tech (Berl). 2013;58:417–27.10.1515/bmt-2012-0049Search in Google Scholar PubMed
13. Holzer B, Grüssner U, Brückner B, Houf M, Kiffner E, Schildberg FW, et al. Efficacy and tolerance of a new gentamicin collagen fleece (Septocoll) after surgical treatment of a pilonidal sinus. Colorectal Dis. 2003;5:222–7.10.1046/j.1463-1318.2003.00471.xSearch in Google Scholar PubMed
14. Nair SC, Anoop KR. Intraperiodontal pocket: an ideal route for local antimicrobial drug delivery. J Adv Pharm Technol Res. 2012;3:9–15.Search in Google Scholar
16. Besch W, Schröder K, Ohl A. Plasmagestützte Oberflächenfunktionalisierung komplex strukturierter, miniaturisierter Kunststoff-Formteile; Plasma assisted surface functionalization of miniaturized plastic devices. Vakuum in Forschung und Praxis. 2005;17:126–30.10.1002/vipr.200500255Search in Google Scholar
17. Briem D, Strametz S, Schröder K, Meenen NM, Lehmann W, Linhart W, et al. Response of primary fibroblasts and osteoblasts to plasma treated polyetheretherketone (PEEK) surfaces. J Mater Sci Mater Med. 2005;16:671–7.10.1007/s10856-005-2539-zSearch in Google Scholar PubMed
18. Zhang Y, Hao L, Savalani MM, Harris RA, Di Silvio L, Tanner KE. In vitro biocompatibility of hydroxyapatite-reinforced polymeric composites manufactured by selective laser sintering. J Biomed Mater Res. 2009; 91: 1018–27.10.1002/jbm.a.32298Search in Google Scholar PubMed
19. Goyal RK, Rokade KA, Kapadia AS, Selukar BS, Garnaik B. PEEK/SiO2 composites with high thermal stability for electronic applications. Electron Mater Lett. 2013;9:95–100.10.1007/s13391-012-2107-xSearch in Google Scholar
20. Goyal RK, Tiwari AN, Negi YS. Role of interface on dynamic modulus of high-performance poly(etheretherketone)/ceramic composites. J Appl Polym Sci. 2011;121:436–44.10.1002/app.33684Search in Google Scholar
21. Wang H, Xu M, Zhang W, Kwok DT, Jiang J, Wu Z, et al. Mechanical and biological characteristics of diamond-like carbon coated poly aryl-ether-ether-ketone. Biomaterials. 2010;31:8181–7.10.1016/j.biomaterials.2010.07.054Search in Google Scholar PubMed
22. Xie G-Y, Zhong Y-J, Sui G-X, Yang R. Mechanical properties and sliding wear behavior of potassium titanate whiskers-reinforced poly(ether ether ketone) composites under water-lubricated condition. J Appl Polym Sci. 2010;117:186–93.10.1002/app.31946Search in Google Scholar
23. Ma R, Weng L, Bao X, Ni Z, Song S, Cai W. Characterization of in situ synthesized hydroxyapatite/polyetheretherketone composite materials. Mater Lett. 2012;71:117–119.10.1016/j.matlet.2011.12.007Search in Google Scholar
24. Ma R, Weng L, Bao X, Song S, Zhang Y. In vivo biocompatibility and bioactivity of in situ synthesized hydroxyapatite/polyetheretherketone composite materials. J Appl Polym Sci. 2013;127:2581–7.10.1002/app.37926Search in Google Scholar
26. Najeeb S, Khurshid Z, Matinlinna JP, Siddiqui F, Nassani MZ, Baroudi K. Nanomodified peek dental implants: bioactive composites and surface modification – A review. International Journal of Dentistry. 2015;2015:1–7.10.1155/2015/381759Search in Google Scholar
27. Stöhr U. Oberflächenaktivierung von Kunststoff mittels Plasma zur Haftvermittlung; Surface activation from plastics through plasma for bonding assistents. Vakuum in Forschung und Praxis. 2015;27:16–21.10.1002/vipr.201500579Search in Google Scholar
28. Friedrich J. The plasma chemistry of polymer surfaces: advanced techniques for surface design. Weinheim, Germany: Wiley-VCH Verlag GmbH & Co. KGaA, 2012.10.1002/9783527648009Search in Google Scholar
29. Awakowicz P, Keil G. Sterilisation von Packstoffen, Hohlkörpern und thermolabilen medizinischen Materialien mit Niederdruckplasmen. Sterilization of packaging materials, hollow bodies and thermolabile medical materials in low pressure plasma. Vakuum in Forschung und Praxis. 2001;13:294–9.10.1002/1522-2454(200110)13:5<294::AID-VIPR294>3.0.CO;2-5Search in Google Scholar
30. Rymuszka D, Terpiłowski K, Borowski P, Holysz L. Time-dependent changes of surface properties of polyether ether ketone caused by air plasma treatment: air plasma treatment of polyether ether ketone. Polym Int. 2016;65:827–34.10.1002/pi.5141Search in Google Scholar
31. Krüger P. Plasmamodifizierung von Kunststoffen und Aspekte der industriellen Umsetzung; Plasmamodification of plastics and aspects of industrial usage. Vakuum in Forschung und Praxis. 2000;12:231–4.10.1002/1522-2454(200008)12:4<231::AID-VIPR231>3.0.CO;2-NSearch in Google Scholar
32. Henneuse-Boxus C, Poleunis C, De Ro A, Adriaensen Y, Bertrand P, Marchand-Brynaert J. Surface functionalization of PEEK films studied by time-of-flight secondary ion mass spectrometry and x-ray photoelectron spectroscopy. Surf Interface Anal. 1999;27:142–52.10.1002/(SICI)1096-9918(199903)27:3<142::AID-SIA493>3.0.CO;2-6Search in Google Scholar
33. Inagaki N, Tasaka S, Horiuchi T, Suyama R. Surface modification of poly(aryl ether ether ketone) film by remote oxygen plasma. J Appl Polym Sci. 1998;68:271–9.10.1002/(SICI)1097-4628(19980411)68:2<271::AID-APP9>3.0.CO;2-NSearch in Google Scholar
34. Schröder K, Meyer-Plath A, Keller D, Ohl A. On the applicability if plasma-assisted chemical micropatterning to different polymeric biomaterials. Plasma Polym. 2002;7:103–25.10.1023/A:1016239302194Search in Google Scholar
35. Awaja F, Zhang S, James N, McKenzie DR. Free radicals generated by ion bombardment of a semi-crystalline PEEK surface. Plasma Processes Polym. 2012;9:174–9.10.1002/ppap.201100056Search in Google Scholar
36. Voit B, Gramm S, Steinert V, Werner C, Zschoche S. Biokompatible und bioaktive polymere Beschichtungen: Modifizierungmethoden für polymere Grenzflächen; Biocompatible and bioactice polymer coatings: modificationmethods for polymer interfaces. Vakuum in Forschung und Praxis. 2011;23:29–33.10.1002/vipr.201100448Search in Google Scholar
37. Noiset O, Schneider Y-J, Marchand-Brynaert J. Surface modification of poly(aryl ether ether ketone) (PEEK) film by covalent coupling of amines and amino acids through a spacer arm. J Polym Sci A Polym Chem. 1997;35:3779–90.10.1002/(SICI)1099-0518(199712)35:17<3779::AID-POLA17>3.0.CO;2-ASearch in Google Scholar
38. Noiset O, Schneider YJ, Marchand-Brynaert J. Fibronectin adsorption or/and covalent grafting on chemically modified PEEK film surfaces. J Biomater Sci Polym Ed. 1999;10:657–77.10.1163/156856299X00865Search in Google Scholar
39. Becker M, Lorenz S, Strand D, Vahl C-F, Gabriel M. Covalent grafting of the RGD-peptide onto polyetheretherketone surfaces via Schiff base formation. ScientificWorldJ. 2013;2013:1–5.10.1155/2013/616535Search in Google Scholar
40. Zhao Y, Wong HM, Wang W, Li P, Xu Z, Chong EY, et al. Cytocompatibility, osseointegration, and bioactivity of three-dimensional porous and nanostructured network on polyetheretherketone. Biomaterials. 2013;34:9264–77.10.1016/j.biomaterials.2013.08.071Search in Google Scholar
41. Du Y-W, Zhang L-N, Hou Z-T, Ye X, Gu H-S, Yan G-P, et al. Physical modification of polyetheretherketone for orthopedic implants. Front Mater Sci. 2014;8:313–24.10.1007/s11706-014-0266-4Search in Google Scholar
42. Vilardell AM, Cinca N, Concustell A, Dosta S, Cano IG, Guilemany JM. Cold spray as an emerging technology for biocompatible and antibacterial coatings: state of art. J Mater Sci. 2015;50:4441–62.10.1007/s10853-015-9013-1Search in Google Scholar
43. Wu W, Geng P, Li G, Zhao D, Zhang H, Zhao J. Influence of layer thickness and raster angle on the mechanical properties of 3D-printed PEEK and a comparative mechanical study between PEEK and ABS. Materials. 2015;8:5834–46.10.3390/ma8095271Search in Google Scholar
45. Zhai T, Di L, Yang D. Electroless nickel-phosphorus coating on poly (ether ether ketone)/carbon nanotubes composite. Electron Mater Lett. 2014;10:631–6.10.1007/s13391-013-3151-xSearch in Google Scholar
46. Barkarmo S, Wennerberg A, Hoffman M, Kjellin P, Breding K, Handa P, et al. Nano-hydroxyapatite-coated PEEK implants: a pilot study in rabbit bone. J Biomed Mater Res A. 2013;101A:465–71.10.1002/jbm.a.34358Search in Google Scholar
47. Hahn B-D, Park D-S, Choi J-J, Ryu J, Yoon W-H, Choi J-H, et al. Osteoconductive hydroxyapatite coated PEEK for spinal fusion surgery. Appl Surf Sci. 2013;283:6–11.10.1016/j.apsusc.2013.05.073Search in Google Scholar
48. Akkan CK, Hammadeh ME, May A, Park H-W, Abdul-Khaliq H, Strunskus T, et al. Surface topography and wetting modifications of PEEK for implant applications. Lasers Med Sci. 2014;29:1633–9.10.1007/s10103-014-1567-7Search in Google Scholar
49. Abdullah MR, Goharian A, Abdul Kadir MR, Wahit MU. Biomechanical and bioactivity concepts of polyetheretherketone composites for use in orthopedic implants-a review: Biomechanical and bioactivity concepts of peek. J Biomed Mater Res A. 2015;103:3689–702.10.1002/jbm.a.35480Search in Google Scholar
51. Ratner BD. Biomaterials science: an introduction to materials in medicine. Academic Press, 2004.Search in Google Scholar
53. Holmberg K, Tiberg F, Malmsten M, Brink C. Grafting with hydrophilic polymer chains to prepare protein-resistant surfaces. Colloids Surf A Physicochem Eng Asp. 1997;123:297–306.10.1016/S0927-7757(96)03810-1Search in Google Scholar
54. Petersen S, Strohbach A, Busch R, Felix SB, Schmitz K-P, Sternberg K. Site-selective immobilization of anti-CD34 antibodies to poly(l-lactide) for endovascular implant surfaces. J Biomed Mater Res B Appl Biomater. 2014;102:345–55.10.1002/jbm.b.33012Search in Google Scholar PubMed
55. Hermanson GT. Bioconjugate techniques. London: Academic Press, 2013.Search in Google Scholar
56. Ponader S. In-vitro- und In-vivo-Untersuchungen zur Biofunktionalisierung von Knochenersatzmaterialien, In vitro and in vivo performance of biofunctionalized bone grafts 2009.Search in Google Scholar
©2017 Svea Petersen et al., published by De Gruyter, Berlin/Boston
This work is licensed under the Creative Commons Attribution-NonCommercial-NoDerivatives 3.0 License.