The uses of polymers are ubiquitous in today’s world due to their low synthesis cost, tunable mechanical properties to suit the intended application, non-toxic degradation products, and ease of manufacturing [1
]. One of the most widespread uses of polymers can be found in the field of biomedical engineering as implants or tissue engineering products [1
]. In many fields, polymeric materials satisfy the requirements of many biomedical applications. However, many polymers have a surface that is bioinert and deficient of free reactive functional groups (e.g., –COOH and –NH2
), lacks topographical features, and has poor wettability, affecting their ability to promote cellular functions and biointegration with the surrounding tissue [7
]. These surface properties often render polymers less suitable than natural biomaterials, e.g., collagen or gelatin, as implants for tissue regeneration. The clinical successes of most of such implants heavily depend on sustained material-cell interactions or bioactivity to facilitate the integration process with the surrounding host tissue [11
]. Poor biointegration could lead to device failure or extrusion, which often requires repeat surgeries to replace the loose implants. Surface engineering to create nanoscale or microscale layers of controlled chemical composition, topography and roughness, and balanced hydrophilicity/hydrophobicity on polymeric implants have emerged as a simple, useful, and versatile approach to alleviate the aforementioned biointegration issue [9
]. Another appeal of surface engineering is that the material’s property improvements can be achieved without significant alteration of the bulk properties of the implantable devices.
One of the most widely used applications in surface engineering has been the use of hydroxyapatite (HAp) coating on orthopedic, dental, and middle ear implant surfaces [15
]. HAp is a type of calcium phosphate (CaP) bioceramics and has attracted the most attention due to its close resemblance to the chemical and mineral components of teeth and bone. It has also been described as a bioactive material due to its inherent ability to induce specific biological reactions from cells or living tissues [19
]. As a result of this similarity, HAp has shown good biocompatibility with bone and tooth, and somewhat surprisingly, with the cornea [16
]. In bone tissue engineering, HAp coating has been shown to enhance bone apposition to orthopedic implants, where it prevents the formation of loose fibrous tissue, but instead forms an extremely thin, epitaxial bonding layer with the bone [23
]. Although HAp coating has not been applied to commercially available corneal prostheses, studies have shown that the bioactive material can enhance biocompatibility, adhesion, and proliferation of corneal stromal fibroblasts in vitro [16
]. The HAp has also been demonstrated to be safe when implanted in vivo [16
HAp coating is regularly applied to metals in load-bearing devices [27
]. For this purpose, various methods have been used to deposit HAp coatings, such as thermal spraying, which includes plasma spray [28
], flame spray [29
], and high-velocity oxygen fuel (HVOF) spray techniques [30
], sputter coating [31
], electron beam deposition [32
], electrophoretic deposition [33
], hot isostatic pressing [34
], and sol-gel methods [35
]. Among them, the plasma spray has been the most widely applied coating technique in dentistry and orthopedics [36
]. It is also currently the only U.S. Food and Drug Administration (FDA)-approved method for applying HAp coating on metallic implant surfaces. A common feature of the abovementioned techniques is high processing and/or annealing temperature that can reach a temperature above 1000 °C. This obviously limits their application for biomaterials with relatively low melting temperature (Tm
), such as poly(methyl methacrylate) (PMMA; Tm
= 160 °C), poly(ethylene glycol) (PEG; Tm
= 60 °C), polylactic acid (PLA; Tm
= 160 °C), and poly(ε-caprolactone) (PCL; Tm
= 60 °C), to name a few [37
]. In addition, methods, such as thermal spraying and sputter coating, can only be applied on surfaces that are in the line of sight and, therefore, are not amenable for coating devices with complex dimensions or with pores [38
]. Some of these techniques also require expensive and elaborate equipment to perform.
Although the clinical application of low Tm
polymers in load-bearing prostheses is uncommon, other applications, including as scaffolds for bone, middle ear, and dental tissue regeneration and craniofacial reconstruction, are regularly studied [6
]. Hence, a non-thermal method to deposit HAp on these polymeric substrates is of interest. A facile, non-thermal approach to deposit HAp is also particularly appealing for application on corneal prostheses, which are typically constructed with an acrylic optic cylinder (e.g., PMMA) that acts as the substitute window to the eye [9
]. In the current review, we discuss two different non-thermal HAp coating approaches, namely, the biomimetic deposition and direct nanoparticle immobilization approaches, for low Tm
polymeric substrates. We also discuss the advantages and limitations of each approach to help readers decide on which particular method is more suitable for their intended applications. We end the review with a summary and future perspective of non-thermal HAp coating.
5. Conclusions and Future Perspectives
Over the last three decades, there have been significant advances made in improving the efficiency of CaP deposition on low Tm
polymers via the biomimetic route. The advantages of the biomimetic approach include the simplicity of the method, the obviation of specialized equipment, and the possibility to coat porous substrates or substrates with complex dimensions. However, the lack of SBF use in clinical practice indicates the need for further research and improvements to overcome these translational challenges. The dearth of publications in this specific area of research for the past five years is another proof of the stagnation of the translational efforts. This could be due to several factors. The coating outcomes have been inconsistent between studies in terms of the crystallinity level, CaP phase, and mineral purity. Several studies found that the resulting minerals were non-stoichiometric HAp and amorphous [22
]. CaP minerals with such properties possess higher solubility than HAp and, therefore, are less suitable for potential long-term clinical uses [103
]. Since SBF contains many other ions, chemically pure HAp cannot be precipitated from the incubation in the solution. Ion-substituted CDHA is the predominant CaP form that is precipitated instead. The inconsistencies of the properties of the coating in many reports are an issue that could be largely attributed to the effect of the different surface functional groups that have to be introduced to induce an efficient apatite nucleation process, as well as the fluctuations in pH and temperature of the SBF (in the preparation, storage and/or during the coating process) [98
]. In SBF, the CaP deposition takes a relatively long time (usually days to weeks) to build up and cover the entire surface of polymers. Because the CaP coating is anchored by ionic bonds to the substrate, delamination can easily occur, especially on devices that frequently experience tangential or horizontal forces [9
]. Upscaling a process that is both inconsistent and takes a long time is not feasible from a practical point of view.
The direct immobilization technique via dip-coating was developed to circumvent some of the aforementioned limitations of the biomimetic CaP deposition. An advantage of the technique is that the fidelity of the phase, crystallinity, and purity of the calcined HAp is maintained in the coating. Hence, the expected bioactivity of the HAp or other nanoparticles is not lost or reduced when immobilized on the polymers [102
]. The superior resistance to biodegradation of HAp compared to other CaP phases may allow longer-term stability of the polymeric implants in vivo. Besides this, the dip-coating and subsequent plasma etching can be carried out in 1 day. Since most low Tm
polymers are soluble in organic solvents, the technique is potentially applicable to those polymers. It may also apply to porous polymers provided the pore sizes are not significantly smaller than the nanoparticles. However, a series of optimization of the concentration of nanoparticles and polymer, the size of nanoparticles, the type of solvent, and the length of dip-coating may have to be performed due to the variations in nanoparticle interaction dynamics and the dissolution rate of different polymers.
The therapeutic effects of the coating deposited via the biomimetic route can be diversified by co-precipitating growth factors/proteins or DNA in the SBF [106
]. The therapeutic effect of the coating deposited by the direct immobilization method is dependent on the bioactivity produced by the nanoparticles used. For example, a mixture of immobilized silver and HAp nanoparticles offers anti-bacterial and improved biocompatibility effects to the polymer [105
]. Due to the likelihood of proteins or DNA degradation in the organic solvent, any functional addition with proteins or DNA cannot be carried out simultaneously with the dip-coating. The advantages and disadvantages of direct immobilization and biomimetic HAp deposition are summarized in Table 2
The American Society for Testing and Materials (ASTM) standards specification F1185-03 states that surgical implants require at least 95% of HAp content, established by XRD analysis, while the concentration of trace elements have to be limited to 3 ppm of arsenic, 5 ppm of cadmium, 30 ppm of lead, and 5 ppm of mercury [108
]. The Ca/P ratio of HAp used for surgical implants must be between 1.65 and 1.82 [108
]. Additionally, the International Organization of Standards (ISO) stated in ISO 13779 that it requires the HAp coating on implants to exhibit a crystallinity of at least 45% with the maximum allowable limit of all heavy metals at 50 ppm [109
]. Although the above standards are currently applied to control the quality and safety of thermal sprayed coating, they can serve as a guideline for the application of non-thermal HAp coating on polymeric products intended for tissue engineering in the future. This offers a translational advantage to the direct immobilization method as the coating procedure does not cause any alteration to the intrinsic properties of the HAp nanoparticles, which have been fine-tuned and synthesized conforming to the ASTM and ISO standards.
In summary, the direct immobilization technique offers advantages of a shorter coating time, obviation of the need for surface functionalization of substrates, and consistency of the crystallinity and mineral phase of stoichiometric HAp in the coating. We have previously optimized the coating method in PMMA. However, due to the novelty of the technique, optimization will be necessary to create a uniform coating on other polymers. It also remains to be explored whether it is possible to apply the technique to polymers with higher Tm
, e.g., polyether ether ketone (PEEK; Tm
= 343 °C) or polytetrafluoroethylene (PTFE; Tm
= 327 °C) [110
]. The possibilities to coat porous polymers and polymers with complex dimensions will also need to be researched in the future to expand the applications. In contrast, SBF-mediated CaP deposition has been extensively studied in the literature and can be readily applied to any polymer, although the outcomes can be somewhat unpredictable. The route to the clinical translation of either coating approach is still long and arduous. Currently, in vivo animal experiments are needed to support the in vitro work and determine the safety and performance of both coatings.