**Ion-Substituted Calcium Phosphates Coatings**

Editors

**Gabriela Graziani Enrico Sassoni**

MDPI • Basel • Beijing • Wuhan • Barcelona • Belgrade • Manchester • Tokyo • Cluj • Tianjin

*Editors* Gabriela Graziani IRCSS Istituto Ortopedico Rizzoli, Rizzoli Orthopaedic Institute Italy

Enrico Sassoni University of Bologna Italy

*Editorial Office* MDPI St. Alban-Anlage 66 4052 Basel, Switzerland

This is a reprint of articles from the Special Issue published online in the open access journal *Coatings* (ISSN 2079-6412) (available at: https://www.mdpi.com/journal/coatings/special issues/calcium phosphates coatings).

For citation purposes, cite each article independently as indicated on the article page online and as indicated below:

LastName, A.A.; LastName, B.B.; LastName, C.C. Article Title. *Journal Name* **Year**, *Article Number*, Page Range.

**ISBN 978-3-03943-543-2 (Hbk) ISBN 978-3-03943-544-9 (PDF)**

c 2020 by the authors. Articles in this book are Open Access and distributed under the Creative Commons Attribution (CC BY) license, which allows users to download, copy and build upon published articles, as long as the author and publisher are properly credited, which ensures maximum dissemination and a wider impact of our publications.

The book as a whole is distributed by MDPI under the terms and conditions of the Creative Commons license CC BY-NC-ND.

## **Contents**

#### **Enrico Sassoni, Elisa Franzoni, Milyana Stefanova, Zdravko Kamenarov, Paolo Scopece and Emanuele Verga Falzacappa**

Comparative Study Between Ammonium Phosphate and Ethyl Silicate Towards Conservation of Prehistoric Paintings in the Magura Cave (Bulgaria)

Reprinted from: *Coatings* **2020**, *10*, 250, doi:10.3390/coatings10030250 ................ **145**

## **About the Editors**

**Gabriela Graziani** (PhD). Gabriela Graziani is a Researcher at Rizzoli Orthopedic Institute. She conducts research regarding materials science applied to orthopedics, with a special focus on nanostructured materials, 3D printing and 3D bioprinting, and on the development of calcium phosphate-based consolidants for stone and archeological remains. She obtained her master's degree and PhD at University of Bologna, where she also served as a postdoctoral researcher. She was Visiting Student Research Collaborator at Princeton University. Currently, she is PI of the project "A New Strategy to Address Tumor Relapses by Nanostructured Implantable Devices", financed by the Italian Ministry of Health.

**Enrico Sassoni** (PhD). Enrico Sassoni is tenure track Assistant Professor at the University of Bologna (Italy). Since 2017, he has held the Italian national qualification as Associate Professor in "Materials Science and Technology". In 2015, he was awarded a 3-year Marie Skłodowska-Curie individual fellowship ("HAP4MARBLE" project), thanks to which he worked for 18 months at Princeton University (USA) and 6 months at the University of Gottingen (Germany). He has authored ¨ and co-authored +80 scientific publications, including 50 journal articles (+1100 citations in Scopus and h-index = 21). He is member of the Editorial Board of RILEM Technical Letters and Coatings. He is Convener of Cluster E of RILEM ("Masonry, Timber and Cultural Heritage") since 2018. In 2017, he received the "Gustavo Colonnetti Medal 2017", awarded by RILEM to researchers less than 35 years old who have provided "an outstanding scientific contribution to the field of construction materials and structures".

## **Preface to "Ion-Substituted Calcium Phosphates Coatings"**

Calcium phosphates (CaPs) in the form of bulk materials and coatings have found application in numerous fields spanning from biomaterials to cultural heritage conservation and water treatment. In fact, this class possesses particular and versatile characteristics that make them appealing for different applications. Among CaPs, hydroxyapatite is that most widely investigated as it is similar to the mineral phase of bone, which has boosted its application in biomaterials in past decades. Lately, however, attention has also been drawn to other classes of CaPs having different solubility and degradation profiles, thus achieving a wide spectrum of materials to be selected based on the specific need. Brushite, alfa- and beta-tricalcium phosphate, and octacalcium phosphate coatings are generating increasing interest and are being explored by a constantly growing number of papers. In addition, hydroxyapatite and CaPs can easily accommodate foreign ions, either incorporated into the lattice or adsorbed on the surface, all altering hydroxyapatite crystallinity, solubility, and overall properties. This can be exploited to guarantee additional properties (i.e., antibacterial, antiosteoporotic, antitumor efficacy) but can also affect nucleation when undesired, like in the case of substrate contamination. For these reasons, several research efforts have been devoted to the development of ion-substituted calcium phosphate coatings through wet and plasma-assisted techniques as well as on the study of hydroxyapatite nucleation in the presence of substrate contaminants. This Special Issue is a collection of contributions regarding the main advances in the field of ion-doped CaPs and their applications. The synthesis and deposition of ion-substituted calcium phosphate coatings are explored together with nucleation of calcium phosphate coatings in the presence of substrate contamination. The relevance of specific ion substitutions (e.g., magnesium, strontium, silver, zinc) on the properties and/or the behavior of substituted hydroxyapatites is reviewed with a special focus on their exploitation for the development of antibacterial CaPs. To achieve these aims, different techniques are used to improve the control over ion-doping of calcium phosphates, and advanced characterization techniques are proposed to investigate their properties.

> **Gabriela Graziani, Enrico Sassoni** *Editors*

## *Review* **A Review on Ionic Substitutions in Hydroxyapatite Thin Films: Towards Complete Biomimetism**

#### **Gabriela Graziani, Marco Boi and Michele Bianchi \***

IRCSS Istituto Ortopedico Rizzoli, NanoBiotechnology Laboratory (NaBi), Bologna 40136, Italy<EMAIL_ADDRESS>(G.G<EMAIL_ADDRESS>(M.B.)

**\*** Correspondence<EMAIL_ADDRESS>Tel.: +39-051-6366042

Received: 19 June 2018; Accepted: 28 July 2018; Published: 31 July 2018

**Abstract:** Plasma sprayed coatings composed of stoichiometric hydroxyapatite have been extensively used to improve integration of metallic implants in the host bone, as hydroxyapatite (HA) is normally regarded as similar to the mineralized phase of bone. However, these coatings exhibited several drawbacks that limited their success. On the one hand biological apatite is a carbonated-HA, containing significant amounts of foreign ions, having low crystallinity and a small crystals size. This means that it differs from stoichiometric HA in terms of composition, stoichiometry, crystallinity degree, crystal size/morphology and, as a direct consequence, solubility, and ions release in the peri-implant environment. On the other hand, thick plasma sprayed coatings can undergo cracking and delamination and are scarcely uniform. For these reasons, research is pushing into two directions: (i) Increasing the similarity of apatite coatings to real bone, and (ii) exploring deposition by alternative plasma assisted techniques, allowing to achieve thin films, and having superior adhesion and a better control over the coating composition. In this article, we review the latest advances in the field of plasma-assisted deposition of ion-substituted hydroxyapatite thin films, highlighting the state of the art, the limitations, potentialities, open challenges, and the future scenarios for their application.

**Keywords:** calcium phosphates; ion-substituted apatites; bone regeneration; plasma-assisted deposition; solubility; crystallinity; composition

#### **1. Introduction**

Orthopedic implants are still burdened by high failure/revision rates, posing relevant economical and societal challenges [1–13].

To avoid failure, implants must integrate in the host site, which requires the capability to trigger host bone cells response without eliciting any immune reaction [14,15]. However, implants also need to burden physiological loads, which can be high and multiaxial, and must not be fragile, hence why metallic implants are normally used, that are bio-inert. A common route to overcome bio-inert behavior of metallic implants and boost integration is to coat their surface with layers having osseointegrative capabilities, so as to render the implant bio-active [16,17]. These coatings are traditionally manufactured by plasma spraying, and have quite significant thickness (comprised between ~30 and 200 μm) [17].

According to the "biomimetic principle", the biological behavior of a device for bone replacement can be boosted by using a biomaterial as similar as possible to real bone in terms of its composition, crystallinity, lattice dimensions, and Ca/P ratio [18]. This, in the field of bone regeneration, has pushed towards an extensive use of hydroxyapatite (Ca10(PO4)6(OH)2, HA), as a bone substitute, in view of its similarity to the inorganic phase of mineralized tissues of vertebrates [16,17,19].

However, clinical results obtained by these traditional coatings are not as promising as expected, and generally not superior to those obtained by bare titanium. This can be ascribed to two main reasons:

For this reason, in the last years, research has been pushing into two directions, to overcome these issues:

In addition, an even more recent route is to deposit coatings directly from bone apatite precursors, which is made possible by plasma-assisted techniques.

Here, we review the literature regarding deposition of ion-substituted apatites by novel thin film plasma-assisted techniques, with special attention to deposition from biological precursors, to highlight the potential, issues and challenges connected to this approach. The focus is on coatings aiming at promoting osseointegration, by a composition as similar as possible to the real bone, hence the Review will be centered on substitution by ions present in the bone. However, because ion-substituted apatites coatings are getting increasing attention for several applications and especially for infections, literature regarding silver-doped hydroxyapatite will be briefly discussed in Section 6, which discusses future perspectives on this topic.

#### **2. Plasma-Assisted Methods for Calcium Phosphate (CaP) Deposition**

Plasma-assisted techniques offer some advantaged compared to wet methods and to traditional plasma spraying. Compared to wet synthesis, they allow for much faster coating of the surface thanks to generally high deposition rate, which is an important advantage from an industrial point of view [21]. In addition, they are industrially scalable, which is also boosting their application. Compared to plasma spraying, they allow for an easier transfer of the target composition, a better uniformity in terms of morphology and composition, a lower porosity and a lower tendency to cracking and delamination [16,22].

Among plasma-assisted techniques, the most used in the biomedical field, for manufacturing thin films for bone implants are radio-frequency magnetron sputtering (RF-MS) and pulsed laser deposition (PLD). Matrix-assisted pulsed laser evaporation (MAPLE) and pulsed electron deposition (PED) are emerging techniques, allowing a fine control over film stoichiometry and composition and are gaining increasing interest, hence they will be included in the review.

In magnetron sputtering (MS), deposition takes place starting from the ionization in vacuum of a target material by means of negatively charged magnets [17]. Although two different modes are available, direct current (DC) and radio-frequency (RF); CaP coatings (and, more in general, non-conductive materials), can be deposited only by the latter [17]. MS allows an easy deposition of multi component coatings, by deposition from multi-component ceramic targets or simultaneous deposition from different targets [24]. Coatings can be very thin (average thickness being ~40 nm–3.5 μm) and can effectively cover surface irregularities, while maintaining a uniform thickness. In addition, they are dense and pore-free [16]. As for all plasma assisted techniques, substrate temperature and deposition time are critical in determining the characteristics of the produced films, but in the case of MS those also have a key role on composition [16]. Films can be deposited at room temperature, which allows coating of heat-sensitive substrates; however, scarce adhesion has been reported for deposition at room temperature, in absence of substrate preparation and/or post deposition treatments. As for all plasma assisted techniques, as-deposited films are amorphous, but both deposition at high temperatures or post deposition annealing are effective strategies to obtain the desired crystallinity degree. Post deposition annealing also allows to tailor surface morphology and roughness to a certain extent. Deposited coatings are generally highly adhesive (bonding strength to the substrate >30 MPa) and exhibit suitable bioactivity. One of the main limitations of this technique is that the coatings composition might differ from that of the deposition target material, also depending on sputtering system and parameters. This, for calcium phosphates, can be very relevant, as significantly different properties can be obtained for slight variations in composition, Ca/P ratio and amount of ionic substitutions. In particular, the Ca/P ratio of the films might vary in a very wide range (1.6 up to 2.6 or even higher) and is normally higher than that of the HA target, due to preferential sputtering of calcium and/or to evaporation of P2O5 [16,24].

In PLD, the target to be deposited is ablated by a focused, high power, pulsed laser beam. As a consequence, a plasma plume forms that is accelerated towards the substrate where deposition occurs. Coatings obtained by PLD can be very thin and much below 1 μm. The characteristics of the deposited films are strongly dependent on deposition parameters that determine their morphology, composition and crystallinity. As for MS, as-deposited films are amorphous, but higher crystallinity can be achieved depending on deposition temperature and possible post-deposition heating treatments. However, because deposition can occur at room temperature, coating of heat sensitive substrates is allowed. Substrate temperature and post treatment annealing also have an impact on the final characteristics of the coatings and in particular on the composition and stoichiometry of the coatings. Despite this, PLD allows for a fine transfer of the target composition. This precise control over stoichiometry makes it possible to deposit multicomponent coatings from one same target [16,17,25,26]. Finally, PLD coatings generally exhibit high adhesion to the substrate, good biocompatibility and bioactivity.

When delicate organic or biological matter or thermally unstable phases needs to be deposited, a modification of PLD can be used, namely Matrix-assisted pulsed laser evaporation (MAPLE). MAPLE prevents the laser source from damaging the target, that is a frozen mixture of the "active" biomaterial (up to 5 wt.%) in volatile solvent, gradually removed by a vacuum system as the deposition proceeds [17,27,28]. The solid target is kept frozen during laser irradiation, so as to prevent evaporation; this is achieved by a cooling system inside the chamber and instant-freezing of the targets by immersion in liquid nitrogen. The solvent is chosen to absorb the laser wavelength, that is also tailored to prevent damage of the target, thus avoiding interactions between the laser and the material, allowing the release of the undamaged delicate active matter, that is deposited on the substrate.

More recently, pulsed electron deposition has been proposed as an alternative to PLD also for applications in the biomedical field [29–31]. In PED the target material if ablated by a high energy pulsed electron beam instead of a laser source; a plasma plume is formed at each pulse and the target material is ionized and deposited onto the substrate [32]. Due to the lower cost of the ablation source, PED is more competitive from an economical point of view, compared to PLD. As in the case of PLD, films deposited by PED are composed by globular aggregates and exhibit high roughness, which makes it suitable for promoting host cells behavior [33]. Films obtained are amorphous, but post deposition annealing can be performed to obtain the desired crystallinity degree [34,35]. Films obtained are highly adhesive, adhesion also being influenced by post deposition annealing [36]. Mechanical properties achieved are comparable to those obtained by plasma spraying [36]. Finally, good biocompatibility and bioactivity have been reported [33].

#### **3. The Role of Ionic Substitutions in Hydroxyapatite**

The main ionic substitutions present in bone are reported in Table 1. The most diffused substitution is CO2− 3 for OH− (A-type substitution) or for PO3− 4 (B-type substitution), followed by and Cl− and F− for OH− [19]. In addition, magnesium, strontium, zinc, and manganese substitutions for calcium, and silicates for phosphates are also present in bone [20,21,37,38]. The amount of these substitutions can significantly vary between different tissues (i.e., bone, dentine, enamel), animal species, types of bone, anatomic sites, etc. [21]. In addition, significant differences can be assessed depending on the patient age, sex, and pathologies.

Ionic substitutions are relevant, as each of them have an influence on HA crystal lattice and hence on its solubility. In addition, some of them have significant biological impact and can have therapeutic effects [19,21]. Those have been largely investigated to obtain substitute HAs for bone regeneration and are their role is reviewed below.

Carbonates cause a decrease in *a*-axis and decrease in *c*-axis of HA, this renders the structure is less crystalline and, as a consequence, more soluble and more bioactive [19,21,39]. Magnesium, the main substitution for calcium in HA, causes a decrease in the *c*-axis of the lattice, acts as an inhibitor of HA nucleation and crystallization and destabilizes its structure. Because Mg-HA is less crystalline than stoichiometric HA, it also has higher solubility [19,21,40]. In addition, Mg is an essential element for living organisms, as it is a key factor in the activity of several enzymes [40]. In bone, Mg influences metabolic activity and growth, by acting on osteoblasts and osteoclasts. As a consequence, scarce concentrations of magnesium can cause osteopenia or fragility [40]. Fluorine substitution for OH− causes a decrease in the *a*-axis and in the c-axis, the latter being lower by an order of magnitude compared to the first. As a consequence, the lattice gains a strong symmetry, that results in an increased stability and a lower solubility [21]. The effect on bone cells, instead, is controversial, as F− itself can boost cells attachment, but the lower solubility causes a reduced availability of Ca2+ ions decreasing cells proliferation. Lastly, fluorine is expected to be an effective anti-osteoporotic agent [21,41]. Strontium can be incorporated in bone by both ion exchange and ionic substitution [19]. The effect of substitution of Sr for Ca is dose-dependent: High concentrations of Sr increase crystals dimensions and crystallinity, while low concentrations cause an opposite effect (decrease in crystallinity and crystalline length), and a distortion of the lattice, both leading to a significant increase in solubility [19]. Interestingly, solubility can be further increased by Sr substitution in carbonated hydroxyapatite (CHA), instead of HA. From a biological point of view, Sr increases the mechanical strength of bone [21] and has an anti-osteoporotic effect [19–21]. In fact, strontium ranelate, a strontium-based drug used against osteoporosis, can boost osteoblast activity while inhibiting osteoclasts [20,21,24,37,38,42,43]. Sr capability to improve bone regeneration has also been demonstrated by tests in vivo [17,20,37,38,42]. As for the effect on crystallinity and solubility, also in vitro and in vivo behavior of Sr is dose-dependent. Silicon is normally incorporated in biogenic HA because of SiO4 4– substitution for phosphate groups [19]. This substitution increases HA solubility, also because formation of amorphous phosphate phases is promoted, the apatite microstructure becomes finer and, finally, silicates form defects at grain boundaries [19]. Even at concentrations of about 1 wt.%, silicon can exert significant biological impact and it is reported to enhance in vivo bioactivity and bone regeneration [44]. Zinc is a trace element in bone [19,21], that can substitute for calcium, thus acting as a crystallinity inhibitor. Despite its low concentration, it has a key biological role, as it can act as a co-factor for several enzymes, can impact on nucleic acid metabolism and promote bone formation and regeneration by favoring osteoclasts activity and decreasing that of osteoblasts [19,21]. Manganese is also a trace element, acting as a crystallinity inhibitor [19]. Depending on the concentration, it can alter HA morphology to a different extent, which has an impact on the apatite bioactivity [19]. Mn has a role in bone remodeling and the influence of Mn-doped carbonated hydroxyapatite on osteoblasts activity has been demonstrated [19,29,45].

#### **4. Deposition of Ion-Substituted Apatites by Plasma-Assisted Techniques**

A scheme of literature about plasma assisted deposition of ion-doped apatites in reported in Table 1.

#### *4.1. Deposition by Magnetron Sputtering*

Deposition of carbonated hydroxyapatite (CHA), magnesium-doped hydroxyapatite (Mg-HA), fluorine-doped apatite (FA), strontium-doped apatite (Sr-HA), silicon doped-apatite (Si-HA), and zinc-doped apatite (Zn-HA) has been reported by magnetron sputtering [39,46].

Regarding carbonated hydroxyapatite, high adhesion (measured by pull-out test) has been obtained for different substrates, further increased when co-sputtering of titanium is performed [46]. Quite fine control over stoichiometry of the films has been achieved (Ca/P 1.80) [46].

Sputtered CHAs are biocompatible, as no toxicity is registered towards host cells. Further, suitable adhesion, viability and proliferation of osteoblasts have been reported. In vitro tests showed mineralization after 21 days immersion in simulated body fluid (SBF), and human bone marrow derived MSCs differentiation in human osteoblasts-like cells (hOBs).

Mg-HA coatings have been deposited by targets with variable magnesium content. In vitro tests indicated high adhesion (by pull-out test) and viability of osteoblasts, though not superior to that obtained on pure HA. The authors ascribed this behavior to amorphous nature of both coating, leading to too high a dissolution, making it impossible to discriminate between the effects of the two.

Fluorine-substituted apatite, deposited on Ti and Ti/Si by right-angle MS, indicated a general decrease in HA grain size, independently from the substrate. Crystallinity of the coatings depends on deposition time, as short sputtering times (i.e., 5 min) led to disordered nanocrystalline phases, while longer times allowed for an increase in crystallinity [47]. However, either pre-heating or post treatment annealing (at temperatures up to 300 ◦C) allowed for significant increases in crystallinity of coatings deposited at room temperature. In particular, by pre-heating, suitable crystallinity can be obtained for temperatures lower than those required for post-deposition annealing [47].

Sr-HA has been deposited both by Sr-HA targets with variable Sr/(Sr + Ca) ratio and by co-deposition from different targets. Deposition on titanium alloys and silicon has been reported [1,24,38,43]. Interestingly, the content of Sr was found to influence sputtering rate.

Scarcely crystalline coatings were obtained, where successful incorporation of Sr in the apatite lattice was achieved [43]. Upon annealing, crystallinity of the films increases and morphology and adhesion experience modifications, as do the (Sr + Ca)/P ratio and the presence of by products (CaO and Sr(OH)2). Coatings obtained are calcium deficient, as a Ca/P ratio of 1.18 was reported. A (Ca + Sr)/P ratio of 1.26 was found, that also is modified by annealing, that increases Sr content with respect to calcium. The (Ca + Sr)/P ratio is lower than that of the target, due to a preferential sputtering of Ca compared to Sr, deriving from different atomic weight. Deposition from multiple targets results in a generally higher ratio [1,38,43]. Remarkable adhesive strength to titanium has been reported—higher than that of stoichiometric HA [24].

As expected, Sr does effectively increase osteoblasts activity and decrease that of osteoclasts, which is clearly indicated by the fact that proliferation and differentiation is higher compared to pure HA [43].

Sputtered silicon-doped apatite (Si-HA) was deposited by both silicate-substituted HA targets and by co-sputtering of Si and HA targets, as seen for Sr. Different substitutions levels were obtained (i.e., Ca10(PO4)6−*x*(SiO4)*x*(OH)2−*x*, *x* = 0, 0.25, 0.33, 0.5) [48]. Silicates can be incorporated in significant amounts, as they can prevent incorporation of carbonates [3].

Even though all deposited coatings are amorphous, crystallinity degree depends on the amount of Si (crystallinity decreases for increasing Si content) [49]. Post deposition annealing does not influence the Si content in apatite, so that the desired crystallinity can be obtained without significantly modifying the coating composition. Deposition at high temperature (with progressive increase of temperature from room temperature up to 200 ◦C) has also been reported, causing the formation of a nanocrystalline coating, with some amorphous areas [48]. In our opinion, these latter results are significantly dependent on the annealing conditions chosen, and higher crystallinity could be achieved for higher deposition temperature, as reported elsewhere for other substituted apatites.

Despite the expected lattice distortion caused by incorporation of silicates in the lattice [49], no significant differences are detected between HA and Si-HA by X-rays diffraction. In addition to the relatively low amount of Si, this is ascribed by the authors to the proximity of silicon and phosphorous in the periodic table. Presence of secondary Si and CaO phases has been reported [50]. Despite a uniformity in the elemental composition within one coating [3,51], both stoichiometric and calcium deficient coatings have been produced [50,51], indicating that, on one hand, silicon does not tend to accumulate in specific areas, as seen, for example, in the case of magnesium. However, in our opinion, this also evidences a non-perfect and fully reproducible transfer of the target stoichiometry.

Silicates addition also have a dose-dependent impact on the mechanical properties (by nanoindentation) of the coating, as a general decrease in hardness and elastic modulus is registered and failure due to scarce adhesion (measured by scratch test) is reported for high Si concentrations [49]. In vitro, increased dissolution rate is found compared to HA [51], possibly to be ascribed to the decrease in OH- content due to increasing silicates substitution [52], that results in a decrease in HA grain size [49].

Dissolution in SBF was found to be significantly dependent on post deposition annealing [50], which is expected, as solubility strongly depends on crystallinity degree. Nevertheless, suitable mineralization and good activity towards bone cells were found for both annealed and non-annealed coatings [50]. Although the mechanism of interaction of Si with bone cells is yet to be fully understood, Si-Ha coatings were found to boost attachment, proliferation and differentiation of osteoblasts with respect to HA and uncoated implants and to favor the formation of an extracellular matrix [3,51,53]. In vivo tests on mice are reported, indicating that Si-HA coatings are osteoinductive, while stoichiometric HA coatings are not. Both coatings, however, are biocompatible [51].

Zn substituted HA has also been deposited by RAMS, but literature is very scarce in this regard, and no deposition is reported by any other plasma-assisted technique, to the authors' best knowledge. In vitro tests on Zn-HA, however, show promising data, as an increased adhesion and viability of osteoblasts was observed, compared to stoichiometric HA [54].

#### *4.2. Pulsed Laser Deposition*

Deposition of carbonated hydroxyapatite (CHA), magnesium-doped hydroxyapatite (Mg-HA) and octacalcium phosphate (Mg-OCP), strontium-doped apatite (Sr-HA) and carbonated

hydroxyapatite (Sr-CHA), silicon doped-apatite (Si-HA), manganese-doped carbonated hydroxyapatite (Mn-CHA), and octacalcium phosphate (Mn-OCP) has been reported by PLD [39,46].

Deposition of carbonated hydroxyapatite allowed to obtain nearly stoichiometric films, Ca/P ratio increasing with deposition temperature (30 up to 700 ◦C) [43]. As for MS, also crystallinity degree and mechanical properties (Vickers microhardness) increase upon annealing [39]. Instead, no significant alterations in morphology are detected for annealing at temperatures below 500 ◦C, while they become pronounced at 700 ◦C and above.

Magnesium-doped hydroxyapatite (Mg-HA) at different magnesium content has also been deposited by PLD. Increasing Mg content was found to cause a parallel decrease in crystallinity and crystal size [40,55]. Effective incorporation of Mg in the lattice was verified. Significantly, Mg distribution is non-homogeneous and tends to form Mg-rich areas, where osteoblasts concentrate, attach and grow [40].

In vivo tests on rabbit models have been reported for both Mg-HA and magnesium doped octacalcium phosphate (Mg-OCP), having the same magnesium content (Mg: 0.6 wt.%) [56]. At sacrifice, 6 months after implantation, good bone growth and integration was found for both coated implants and for the uncoated titanium reference. No inflammatory reactions or failure were reported for any of the implants. Bone volume, bone-implant contact length and contact length ratio to implant surface has been found to be higher for Mg-HA and Mg-OCP implants compared to uncoated implants (μCT). Performances of Mg-HA were found to be higher than those of Mg-OCP, in our opinion also due to differences in solubility of the two compounds.

Strontium substituted hydroxyapatite (Sr-HA) [17] and strontium substituted carbonate apatite (Sr-CHA) [42] have been tested. Thin films of Sr-HA were obtained (200 nm up to 1 μm), that were annealed after deposition [17,42]. The presence of Sr in the coatings, which was verified by EDS, did not significantly impact on coatings morphology, as found elsewhere for plasma sprayed coatings [17].

In vitro, osteoblasts adhesion, proliferation, viability and differentiation were found to be boosted, the increase in adhesion (tensile adhesion by ASTM C633-01(2008) [57], ASTM F1147-05:2011 [58], ISO 20502:2005 [59]) since the early steps being proportional to Sr content [17]. Conversely, osteoclasts proliferation was found to be reduced, again depending on Sr content (3%–7%). Interestingly, spreading of osteoblasts is increased compared to stoichiometric HA coatings.

Deposition of Si substituted hydroxyapatite (Si-HA) has been reported, composition of the targets being successfully and finely transferred to the coatings [60]. In a study by Rau et al., it was found that the Si content must be kept below 2 wt.% to avoid a reduction in the decomposition temperature. This ultimately caused the formation of unwanted unstable phases, such as TCP and/or phosphate silicates [61]. Below 2 wt.%, instead, Si does not significantly alter HA lattice, as denoted by XRD spectra, almost identical to those of HA.

Addition of silica can cause an increase in the Ca/P ratio compared to the target, due to the formation of CaO [60,61]. This formation is relevant, as CaO is cytotoxic, hence compatibility of the coatings must be verified. Ca/P ratio is also dependent on the deposition temperature and/or on post-deposition annealing, as seen for MS [61]. Several temperatures have been tested in the literature for deposition of Si-HA (400, 500, 600, and 750 ◦C), all leading to deposition of dense and un-cracked films [61]. Films obtained at these temperatures are crystalline and nanostructured, crystallinity degree being obviously depending on annealing temperature. Interestingly, also thickness and Vickers microhardness are affected by deposition temperature [61]. However, despite mechanical properties generally increasing for increasing deposition temperature, when the temperature is brought above 750 ◦C, formation of byproducts (calcium-titanium mixed phases) hampers adhesion. Also, at these temperatures, Ca/P ratios completely different from those of the target (≈3.5) are obtained, indicating HA decomposition alongside formation of byproducts. As a consequence, deposition temperatures around 500 ◦C are generally recommended.

In vitro mineralization is reported after 1 day of immersion in alfa-MEM, thickness of the HA-layer increasing for increasing soaking times (up to 28 days) [61].

Mn-substituted CaP coatings have been manufactured exclusively by PLD, to the authors' best knowledge. Deposition of both Mn-CHA and Mn-OCP coatings have been described in [62]. While Mn-CHA is more similar to real bone but has higher crystallinity and consequently slower dissolution rate, Mn-OCP is more soluble and is expected to experience faster dissolution—thus possibly being more effective in boosting bone growth at early stages. When tested in vitro, human osteoclasts and fibroblasts exhibited suitable adhesion, proliferation and viability onto both coatings, without any significant difference being assessed. This is ascribed by the authors to a compensation between the beneficial effects CO2− 3 of Mn-CHA, and the lower crystallinity of Mn-OCP.

These coatings were also validated in vivo in rabbit models, in comparison to HA coatings. Tensile strength by pullout test was tested at 8 weeks after implantation [63], indicating that Mn-OCP and, to an even higher extent, Mn-CHA, can better improve implant-bone integration with respect to HA, suggesting a higher osseointegration potential.

#### *4.3. Deposition by Matrix Assisted Pulsed Laser Evaporation*

Deposition of Sr-HA and Sr-OCP has been reported by Combinatorial Matrix-Assisted Pulsed Laser Evaporation (C-MAPLE).

The possibility to deposit metastable Sr-OCP and composites of hydroxyapatite and drugs (here alendronate and zoledronate), evidences the strength of this technique, that allows deposition of heat-sensitive material [27,64]. Sr-doping on OCP was found to improves cells proliferation, activity and differentiation and, hence, the overall behavior of the coating. Strontium zoledronate and HA composites, instead, were found to promote osteoblasts adhesion and boosts proliferation, compared to uncoated references.

#### *4.4. Deposition by Pulsed Electron Deposition*

Deposition of Sr-HA has been very recently reported by PED (in the novel Ionized Jet Deposition version) for variable Sr content (up to 9 wt.%) [36]. Deposition has been carried out at room temperature on a polymeric substrate (PEEK) leading to an increased in wettability and roughness compared to the bare substrate, and suitable mechanical properties. A good transfer of composition (including Sr content) has been assessed from the target to the deposited coatings [36].

#### **5. Multi-Substituted Coatings and Deposition from Biogenic Sources**

Multi substituted HAs have also been explored in the literature, to increase similarity to bone apatite and to exploit the capability of each ion to influence HA lattice, crystallinity and solubility, and the coatings mechanical properties and interactions with cells [24].

In this view, pulsed laser deposition of silicon and strontium co-substituted carbonated hydroxyapatite (SiSrCHA) has been reported [65]. SiSrCHA has been obtained by diatomaceous earth, naturally containing several further foreign ions (Fe, K, Mg, Na, and P), all possibly boosting biological performances [65]. In the study, silica has been kept constant (2.5 at.%), while Sr concentration has been varied in a wide range (0 up to 10 at.%). Deposition has been carried out at high temperature to tune crystallinity and adhesion. Mesenchymal cells adhesion, morphology and osteogenic differentiation and activity have been tested. Results showed that, for multi-substituted coatings, variations in the amount of one of the dopants might affect the content of the others. In fact, aside from an obvious decrease in Ca content, due to a progressive incorporation of Sr, increasing Sr also leads to alterations in carbonates content. This can be seen by a reduction in the intensity of carbonates bands in FT-IR and in the CO2− 3 to PO3− 4 ratio, compared to pure CHA. We ascribe that to the modifications caused by each substitution to the HA lattice, influencing its capability to host further foreign ions.

Instead, doping with Si and Sr does not influence cells morphology compared to pure CHA reference coatings, as good cells coverage is found at 28 days for all the examined group. Instead, increased cells proliferation and differentiation were found for the multi-doped coating, compared

to references. We speculate that examining growth at lower times could be very useful to assess biological behavior.

A further, crucial step in the direction of obtaining coatings having increasing similarity to bone apatite is deposition directly from biogenic targets. Deposition of apatite coatings from targets from a biogenic source have been reported by PLD [66–69] and PED [33,34]. Deposition of doped bone apatite targets has also been reported [67,68].

In particular, by PLD apatite has been deposited starting from simple, LiO2- and CIG-doped ovine and bovine bone, MgF2, and Mg doped bovine bone, simple and titanium doped ovine dentine and human dentine. Further experiments about deposition of "natural apatite" by PLD have also been reported, tested in vitro and in vivo, but without suitable description of the apatite source and the deposition procedures [70,71].

Ovine and bovine bone-derived coatings were obtained by PLD, after calcination of the powders at 850◦C for 4 h in air to remove organic components [66]. Characterizations performed on the targets, obtained by compression of the calcinated apatite powders, indicate the presence of brushite and phosphorous fluoride hydroxide traces, alongside HA. HA (calcium deficient) was also found to be the main phase in the coatings. Traces of Na, Mg, Cl, F and Si were detected in the coatings and the targets, suggesting that ion substitutions were preserved. No in vitro and/or in vivo biological tests are reported for these coatings.

HA powders from sheep dentine were also calcinated (6 h at 850◦ in air) to remove organic components and allow an easy separation between dentine and enamel [67]. After grinding and sieving to obtain particles of a suitable size, part of the powders was admixed with Ti powders. Both Ti-doped and un-doped powders were homogenized and pressed to obtain targets, later heated at 1000 ◦C for 4 h in air, and deposited by PLD.

Morphology and composition of the coatings were investigated, indicating that HA is the main phase in the coatings, alongside Ti minor phase and, possibly, MgCO3 and CaMg2, possibly resulting from calcination. Traces of Na, Mg, C and Ti (in the doped samples) were detected in the coatings by EDS [67].

Both coatings, respectively 1.8 microns (undoped) and 1.2 microns (doped) thick are composed of granular aggregates, providing a high surface roughness, suitable for cells adhesion. Addition of Ti also has an impact on the coatings morphology, as it causes a reduction in average grain size compared to undoped Ti and a denser particulate. Pull-out test (ASTM D4541-17 [72] and ISO 4624:2016 [73]) indicates a good bonding strength, increased in the case of Ti-doping (55.5 ± 5.5 MPa vs. 63.7 ± 5.2 MPa), due to Ti-Ca segregation at the interface between the coating and the substrate, creating a mingled interlayer. Biocompatibility was assessed by evaluating human mesenchymal stem cells (hMSC) morphology and proliferation, and toxicity assay, showing evidence of good biological behavior for both coatings, with and without titanium addition [67]. Bovine apatite from femoral bones was also used, alone and enriched with MgF2 and MgO. As for ovine dentine, prior to deposition, bovine bone powders were calcinated (4 h at 850◦ in air), mixed, pressed in a mold and sintered (4 h at 1000 ◦C) to obtain the deposition targets. Post deposition annealing was performed to enhance crystallinity of the coatings [68].

Morphology, composition and adhesion (ASTM D4541-17 [72] and ISO 4624:2016 [73]) of the coatings were evaluated, then the influence of the targets and the coatings on HEp-2 cells viability was tested and gene expression and antimicrobial activity of the films were evaluated [68].

A close transfer of composition from the target to the coatings was discovered, with the absence of MgF2 peaks and weak MgO peaks, that the authors attributed to Mg2+ and F− substitution in the HA lattice. Peaks characteristic of CHA were shown by FT-IR. MgO and MgF2 enhanced the adhesion to the substrate and to provide anti-biofilm properties to the coatings. All films exhibited good biocompatibility, as none of them exhibited relevant cytotoxicity [68].

Human dentine, bovine-derived HA and ovine-derived HA were also deposited after calcination of the powders (4 h at 850◦), pressing and sintering at 1100 ◦C for 4 h [69]. Deposition was performed at high temperature (500 ◦C) and post-treatment annealing at the same temperature (6 h, water vapor) was also carried out. Composition and crystallinity of the deposited coatings were investigated, showing a good transfer of substrate composition, including that of trace elements. Thickness of about 1.1–3.5 microns was obtained, depending on the precursor used. Pull-out tests (ASTM D4541-17 [72] and ISO 4624:2016 [73]) indicated a good adhesion to the substrate [68].

Bovine bone-deriving apatite has also been deposited by PED, including a more recent development of the technique called Ionized Jet Deposition (IJD) [33,34].

Bone apatite-like films were deposited directly from deproteinized bone targets, without any grinding or calcination. Deposition was performed at room temperature and half of the samples were heat treated after deposition, to improve crystallinity (1 h at 400 ◦C in air) [33,34].

Obtained films were compared to films obtained by sintered stoichiometric-hydroxyapatite targets with the same deposition procedure and parameters, both as deposited and annealed [33,34].

As-deposited and annealed coatings were characterized in terms of composition and crystallinity, microstructure and morphology, and mechanical properties (nanoindentation and micro-scratch, UNI EN ISO 20502:2005 [59]). Obtained coatings are nanostructured and exhibit sub-micrometric thickness. Results showed a close transfer of the targets' composition, including trace elements (Na and Mg and carbonate substitution). As-deposited coatings exhibited low crystallinity, that was significantly increased by post-deposition annealing, up to resembling that of biogenic apatite target. As a result of annealing, mechanical properties increased up to values comparable to those of commercial plasma-sprayed HA-coatings [34].

Biological tests are also reported, evaluating human dental pulp stem cells (hDPSCs) cells attachment, proliferation and the expression of typical osteogenic markers Runx-2, osteopontin, Osx, and Osteocalcin [33].

In vitro biological tests indicated that annealed bone apatite coatings promote hDPSCs proliferation to a higher extent compared to non-annealed bone coating and HA-references. Immunofluorescence and western blot analyses revealed that the typical osteogenic markers were expressed, suggesting that bone apatite-like coatings treated at 400 ◦C alone can efficiently promote the osteogenic commitment of the cells, even in absence of an osteogenic medium [33].

#### **6. Conclusions and Future Perspectives**

Plasma-sprayed (PS) coatings have extensively studied and used in clinical practice but, due some intrinsic limitations, their real effectiveness compared to uncoated implants is still controversial. Instead, good results have been so far obtained for novel plasma-assisted techniques, capable of depositing thin films. In fact, they were found to allow a more controlled transfer of stoichiometry of the deposited coatings and of ionic additions to plasma spraying. However, presence of secondary phases often reported for magnetron sputtering together with preferential sputtering of some ions (i.e., Ca instead of Sr), cause modifications in the Ca/P and substitution rate, and indicates an imperfect transfer of the target composition. Because relatively low changes in the composition of CaPs, especially when substituted can result in significantly different properties, we believe that it is a relevant limitation in the technique, which needs to be addressed. PLD, MAPLE and PED, instead, allow for a more precise transfer of the target composition to the coatings.

Crystallinity degree can also be finely tuned, by deposition at selected temperature or pre/post deposition annealing. Good adhesion and mechanical properties in general, were confirmed by all the examined studies, as well as good behavior in vitro in terms of adhesion, spreading, differentiation, morphology, and osteogenic activity of the cells, although not always higher than those of stoichiometric hydroxyapatite. Some promising results have also been described in vivo, but extremely scarce literature on the topic and differences in trial conditions (animal model, trial time, characteristics of the bone, gap dimensions, etc.) make them scarcely conclusive.

Overall, thin films of ion-substituted apatites are promising materials, having good mechanical properties and biological behavior. Ion-substitution permits the availability of selected ions in selected concentrations in the peri-implant environment, with both a therapeutic effect and/or the capability to boost host cells response in view of the biomimetic principle. Deposition by plasma-assisted techniques allows to obtain thin films, having high adhesion and a fine control over crystallinity and composition, guaranteeing a tailored release.

Deposition from biogenic coatings has been emerging as a very promising alternative, as it allows to obtain multi-substituted coatings with ion-amounts and stoichiometry very close to those of real bone apatite, which is almost impossible to be achieved by artificial apatites. This new development has been made possible by PLD and PED, that allow a great control over the composition of the obtained targets. We believe that, once in vivo trials will be available for these coatings, this topic will obtain even wider interest and will expand its field of application.

Interestingly, however, despite ion-substituted coatings being generally proposed with the aim to better resemble real bone, mimicking of the real composition of bone is very limited, as generally only one or two ions are substituted simultaneously in doped HAs. In addition, there is a lack of correspondence between the concentrations of ions in biological HA and their biological relevance, and the number of studies devoted to their investigation. This can be partially ascribed to the wide variability in the concentration of trace ions in the bone, depending on the anatomical site, age, sex and to patient-specific characteristics such as, for example, diseases or metabolism. However, some ions, such as magnesium (highly present in bone and having significant biological relevance), is far less studied than Sr, whose concentration is significantly lower. Some other ions, such as Cl, Na and K are not even intentionally substituted. The number of studies, instead, seems more driven by the possible therapeutic role of the ions. This is clarified by intentional addition of specific therapeutic ions in biological apatite, which goes in the opposite direction of biomimetism. We believe that this, however, can be a promising route, further expanding the range of use of ion-substituted coatings.

Another important trend that is emerging regarding ion-substituted hydroxyapatite thin films, is that of antibacterial coatings. To this regard, a wide range, and increasing amount, of literature is available about silver-doped hydroxyapatite, for which deposition by magnetron sputtering [74–78], PLD [79–84] and MAPLE [85] has been reported. Being the study of plasma assisted deposition of Ag-doped hydroxyapatite of more recent introduction, the majority of the studies focus on the influence of deposition parameters on the films characteristics and on the evaluation of antibacterial efficacy, while extensive, comparative in vivo studies are yet to be performed. We believe that plasma-assisted deposition of apatites doped with a variety of ions having antibacterial (such as copper [86] and gallium), antitumor or other therapeutic actions (such as selenium [87] and cerium) will be increasingly pursued in the near future.

Overall, what we believe to be extremely significant is the possibility to pursue both aims, i.e., biomimetism and therapeutic functions, thanks to plasma assisted techniques.

We believe that extensive work still needs to be carried out to fully evaluate properties, potentialities and limitations of ion-doped apatite coatings. The main points that should be explored, in our opinion, are the following:

**Funding:** This research was funded by INAIL-Istituto Nazionale Assicurazione Infortuni sul Lavoro (National Institute for Insurance against Accidents at Work), Project METACOS "Trattamento delle amputazioni mediante osteointegrazione" and by Istituto Ortopedico Rizzoli "5 × 1000" year 2015.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

© 2018 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (http://creativecommons.org/licenses/by/4.0/).

#### *Article*

## **Pulsed Laser Deposited Biocompatible Lithium-Doped Hydroxyapatite Coatings with Antimicrobial Activity**

**Liviu Duta 1,\*, Mariana Carmen Chifiriuc 2,3, Gianina Popescu-Pelin 1, Coralia Bleotu 4, Gratiela (Pircalabioru) Gradisteanu 3, Mihai Anastasescu 5, Alexandru Achim 6 and Andrei Popescu 6,\***

Received: 5 December 2018; Accepted: 16 January 2019; Published: 17 January 2019

**Abstract:** Simple and lithium-doped biological-origin hydroxyapatite layers were synthesized by Pulsed Laser Deposition technique on medical grade Ti substrates. Cytotoxic effects of lithium addition and the biocompatibility of obtained coatings were assessed using three cell lines of human origin (new initiated dermal fibroblasts, immortalized keratinocytes HaCaT, and MG-63 osteosarcoma). Antimicrobial properties of obtained coatings were assessed on two strains (i.e., *Staphylococcus aureus* and *Candida albicans*), belonging to species representative for the etiology of medical devices biofilm-associated infections. Our findings suggest that synthesized lithium-doped coatings exhibited low cytotoxicity on human osteosarcoma and skin cells and therefore, an excellent biocompatibility, correlated with a long-lasting anti-staphylococcal and -fungal biofilm activity. Along with low fabrication costs generated by sustainable resources, these biological-derived materials demonstrate their promising potential for future prospective solutions—viable alternatives to commercially available biomimetic HA implants—for the fabrication of a new generation of implant coatings.

**Keywords:** lithium-doped hydroxyapatite coatings; renewable resources for implant coatings; pulsed laser deposition; biocompatibility; inhibition of microbial biofilms development

#### **1. Introduction**

Microbial biofilms development on medical devices and human tissues represents a serious public health problem at global level, leading to a high diversity of biofilm-associated infections, with increasing incidence, although yet underestimated and high rates of morbidity and mortality. The biofilm growth protects microbial cells from antimicrobials and host immune effectors, enabling pathogens to survive in limited conditions and also to disperse and colonize new niches. The biofilm-embedded microbial cells exhibit a high tolerance to conventional antimicrobial agents, explaining the chronic and relapsing character of the produced infections [1–3]. The increasing

incidence of biomaterial-associated infections is in close relationship with the continuously growing demand for implantable medical devices and other invasive diagnostic and therapeutic maneuvers. Therefore, controlling biofilm formation represents a mandatory feature during medical devices manufacturing and their use [4].

Biofilm development is a dynamic and complex process which occurs in several stages [1]. It starts with a reversible adhesion stage, followed by irreversible adhesion [5], maturation and, in the last stage, detachment. The first stages depend on the physico-chemical properties of the surface, conditioning the interactions with microbial structure involved in the adhesion process [6]. The main pathogens involved in the development of biofilm-associated infections include the Gram-positive bacteria, such as *Staphylococcus aureus* (*S. aureus*), and fungal strains, especially *Candida* spp. Reports from the National Institutes of Health demonstrate that around 80% of human infections are due to biofilms [7]. The biofilm-associated infections are characterized by moderate intensity symptoms, chronic evolution and high resistance to antibiotics [8]. Therefore, one main direction bridging materials science to biomedical research is the development of complex therapeutic strategies, including the association of multiple antimicrobial agents [9–11]. However, one should note that the association of multiple antibiotics might lead to high toxicity effects, correlated with a low efficiency on the cells embedded in biofilms [12]. Therefore, the development of preventive measures, based on the design of novel biomaterials, synthesized in the form of thin layers and used for the fabrication of medical devices which are more resistant to microbial colonization and biofilms development could represent a more effective approach [13,14].

Synthetic hydroxyapatite (HA) represents a well-known biocompatible material used as implantable ceramic due to its great chemical and structural resemblance to the inorganic part of human hard tissues [15,16]. Reports from dedicated literature have demonstrated that biological-derived HA (BHA) has some important advantages over synthetic HA. In this respect, to improve mechanical (anchorage to bone) and biological (osseointegration, resistance to microbial colonization) properties of medical implants, simple BHA materials [17] or BHA doped with specific reagents such as ZnO [18], SiO2, MgO, Al2O3 and ZrO2 [19], Ti [20], Li2CO3 and Li3PO4 [21], Li2O [22], commercial inert glass [22,23], CaMgSiO6 [24], MgF2 or MgO [25], and CO2 [26] have been used [27].

The present study is a first attempt in dedicated literature to fabricate, in a single-stage pulsed laser deposition process, BHA coatings to exhibit low cytotoxicity and good biocompatibility, corroborated with an active role in prevention of microbial colonization and biofilm development on prosthetic devices.

#### **2. Materials and Methods**

#### *2.1. PLD Experimental Protocol*

BHA powders were obtained from the femoral bones of freshly slaughtered cattle (received from a Turkish slaughter house, which uses the other animal parts for general consumption). Bones were delivered on ice to the laboratory and, before usage, they were submitted to a veterinary control. Concisely, both heads of femoral bones were cut off (because they contained too much soft tissue) and only femoral shafts were further processed. Bone marrows were extracted and all other soft tissue residues were removed from shafts. Then, femoral shafts were cut into slices, cleaned and washed with distilled water. Cleaned parts were deproteinized for 14 days in an alkali solution of 1% sodium hypochlorite. After washing and drying, bone pieces were calcinated at 850 ◦C for 4 h in air, in order to completely eliminate organic components of bone. We emphasize that as-prepared BHA powders are biologically safe due to the high temperature fabrication route, which not only favors the crystallization of the material, but also totally prevents any risk of diseases transmission, as no pathogen can survive to such extreme conditions [20–22,25]. Resulting calcined bone specimens were first crushed with a mortar and pestle and then ball-milled (for 4 h) to fine powders (i.e., with particles of submicron size). The color of the resulting powders was milky-white. One should note that the

experimental procedure for the fabrication of BHA powders was conducted in accordance with the European Regulation 722/2012 [28] and ISO 22442:2015 standard [29]. Batches of BHA powders were admixed with 1 wt.% of Li2CO3 or Li3PO4 (Sigma-Aldrich GmbH, St. Louis, MO, USA). Sample codes are introduced in Table 1 and will be further used in the text.

**Table 1.** Sample codes and description.

PLD experiments were carried out inside a stainless-steel deposition chamber. A KrF \* excimer laser source COMPex Pro205 (Coherent, Göttingen, Germany) (λ = 248 nm, τFWHM ≤ 25 ns, υ = 10 Hz) was used for target irradiation.

The laser beam was oriented at 45◦ in respect with the target surface. The laser fluence incident onto the target surface was set at ~4 J/cm2 (with a corresponding pulse energy of ~360 mJ).

Before deposition, targets were "cleaned" by short multipulse laser ablation with 1000 pulses. During deposition, the target was continuously rotated with 0.4 Hz and translated along two orthogonal axes to avoid piercing and allow for the growth of uniform films.

The ablated material was collected onto medical-grade titanium (Ti) disks (~0.6 × 0.6 cm2 and 0.1 cm thickness), that were placed parallel to the target at a separation distance of 5 cm.

All experiments were performed in a 50 Pa water vapor atmosphere. During experiments, substrates were heated and maintained at a constant temperature of 500 ◦C using a PID-EXCEL temperature controller (Excel Instruments, Gujarat, India). A heating rate of 25 ◦C/min and a cooling ramp of 10 ◦C/min have been used. For the deposition of each structure, 15,000 subsequent laser pulses were applied.

So as to improve the crystalline status of synthesized films, all samples have been subjected to a 6 h post-deposition thermal treatment in air (in water vapors), at 600 ◦C.

#### *2.2. Characterization of Synthesized Structures*

#### 2.2.1. Atomic Force Microscopy

The surface topography of synthesized structures was examined in high resolution Atomic Force Microscopy (AFM) (Park Systems, Suwon, Korea), performed in the non-contact mode on a XE-100 apparatus from Park Systems. Sharp tips (PPP-NCHR type from NanosensorsTM, Neuchâtel, Switzerland), having 125 μm in length, 30 μm width and radius of curvature of less than 8 nm, were used at surface scanning over areas of 0.5 × 0.5 μm2. The spring constant was of ~42 N/m and the resonance frequency of ~330 kHz, respectively. All AFM measurements were performed at room temperature.

AFM images were processed using Image Processing Program, XEI–v.1.8.0, developed by Park Systems (Suwon, Korea). The horizontal line by line flattening was used as planarization method for displaying purposes and for subsequent statistical data analysis, including the calculation of the roughness. To have a better view of the surface morphology, 2D-AFM images are displayed in the classical view mode (one color used also for the z-scale bar), while 3D-AFM images are shown as topographic maps in the so-called "enhanced color" view mode (using the change of a pixel color relative to its neighbors).

The textural (amplitude) parameters, namely average roughness (*R*a), surface skewness (*R*sk) and surface kurtosis (*R*ku) were determined giving information about the roughness and the degree of skewness and kurtosis of investigated surfaces. *R*a is the area between the roughness and its mean in contrast with the root-mean-squared roughness, which is the standard deviation of the height value in the selected region. In statistics, *R*sk is a measure of the asymmetry of the probability distribution of a real-valued random variable about its mean. In the present case, *R*sk describes the asymmetry of the height distribution (frequency vs. height histogram). *R*ku measures the randomness of profile heights and of the "sharpness" of the height distribution, and has a value of 3 in the case of a perfectly random surface [30]. Kurtosis is not related to the geometry of the distribution, but to the shape of it. Kurtosis is a measure of whether data are heavy-tailed or light-tailed, relative to a normal distribution (*R*ku = 3). Thus, data sets with high kurtosis, *R*ku > 3 (sharper height distributions–so called "lepto kurtosis") tend to have heavy tails, or outliers. Data sets with low kurtosis, *R*ku < 3 (flattened height distribution–so called "platy kurtosis"), tend to have light tails, or lack of outliers.

#### 2.2.2. In Vitro Biological Tests

The potential use of synthesized coatings in the biomedical field was appraised by both safety (cytotoxicity and influence on the cellular cycle of three different human cell lines) and anti-biofilm tests (performed on bacterial and fungal strains).

#### Biocompatibility Assays

#### (i) Cytotoxicity Assays

The cytotoxicity of obtained coatings was investigated using three different cell lines, i.e., MG-63 osteosarcoma ECACC 86051601, new initiated dermal fibroblasts (FBD) and immortalized keratinocytes HaCaT [31].

The cytotoxicity of samples was investigated after microscopic examination of the unstained monolayers developed on their surface. Changes of the general morphology and monolayer confluence were assessed. The cytotoxicity level of the tested samples was classified in accordance to the ISO standard recommendations [32].

The UV-sterilized samples were placed in duplicate in 24-well plates exposing the coated face. A total of 1 × 105 cells in Dulbecco's Modified Eagle Medium:F12 supplemented with 10% fetal bovine serum were added. Plates were maintained for 24 h at 37 ◦C and 5% CO2 atmosphere. After 24 h, samples have been observed with an inverted microscope and photographed in contrast phase with a Zeiss Observer D1 microscope (Zeiss, Oberkochen, Germany). Due to the fact that BHA samples were opaque, the examination and images of cellular monolayers were taken in the vicinity of tested materials. Then, one of the duplicate samples was extracted from the well, fixed in ethanol 70%, stained for 5 min with propidium iodide (PI) (chemical name 3,8-Diamino-5-(3-(diethyl(methyl)ammonio)propyl)-6-phenylphenanthridin-5-ium iodide) 50 μg/mL, and visualized in red channel, using a fluorescence Leica DFC450C microscope (Leica, Allendale, NJ, USA).

#### (ii) Cellular Cycle Assay

For the investigation of the cellular cycle, samples (treated as mentioned in the previous section) were harvested with trypsin, fixed in 70% cold ethanol over night at −20 ◦C, and then washed with phosphate-buffered saline (pH 7.5), treated with 100 μg/mL RNase A for 15 min and stained with 10 μg/mL PI for 1 h at 37 ◦C. After PI staining, the events (number of cells passing through the laser) acquisition was performed with an Epics Beckman Coulter flow cytometer (Beckman Coulter Inc., Indianapolis, IN, USA). The obtained data was analyzed with the FlowJo software 7.6.4 and expressed as fractions of cells found in different cell cycle phases [33].

#### Antimicrobial Assays

Two microbial strains traceable to American Type Culture Collection (ATTC), belonging to two of the most important species known to be involved in the etiology of medical devices biofilm-associated infections, respectively *S. aureus* and *Candida albicans* (*C. albicans*), were used in the experiments. Fresh cultures were obtained from glycerol stocks of *S. aureus* ATTC 6538 and *C. albicans* ATTC 26790 strains on non-selective solid media. Samples were sterilized by exposure to UV irradiation in a laminar flow chamber for 30 min on each side, placed in duplicate in six-well plates (Nunc) in 2 mL of Luria Broth (LB) nutrient medium and then inoculated with 500 μL of standardized microbial suspensions of 0.5 McFarland density corresponding to ~105–106 Colony Forming Units (CFU)/mL. After incubation at 37 ◦C, for different time periods (1, 6, 12, 24, 48 and 72 h, respectively), colonized samples were gently washed with distilled water to remove unattached microbial cells and then immersed in 1 mL sterile saline buffer in Eppendorf tubes, sonicated for 15 s, and then vortexed for 15 s at 3000 rotations/min. Suspensions resulted from the detached biofilms were then ten-fold diluted and 10 μL of each serial dilution were plated in triplicate on tryptic soy agar. After 24 h of incubation at 37 ◦C, colonies were counted and the number of CFU/mL was determined.

The surface characterization of inoculated surfaces was carried out using Scanning Electron Microscopy (SEM) with a FEI Inspect S50 apparatus (FEI, Hillsboro, OR, USA). Prior to this investigation, samples were washed three times in PBS and then fixed in ethanol, for 5 min, at RT.

#### **3. Results**

#### *3.1. AFM Analysis*

Typical bidimensional (2D) and tridimensional (3D) topographic images collected at a scanning area of (0.5 × 0.5) μm2 are presented in Figure 1. The 3D AFM images, which are topographic maps of scanned areas, are presented in the Enhanced ColorTM view mode (Park Systems) in order to better emphasize the morphological surface peculiarities. In 3D-AFM micrographs, dark color (blue) delineates surface "valleys", whilst the light color (orange) corresponds to the salient features ("hills"). The 2D AFM images are presented in the "classical" view mode, with one color gradient (brown in this case) for the z-scale depth. The first row presents the morphological features of the bare substrate (Si), which are located in a z-scale depth (Δz) of less than 1 nm (typical ~0.7 nm), which leads to a very low *R*a value of only 0.1 nm (which is representative for a clean Si wafer), as seen in Table 2. The *R*sk parameter, describing the asymmetry of the height distribution histogram, has a small value in this case, of 0.9, which is corresponding to a flat surface with random (small) protuberances. The *R*ku parameter, which represents a measure of the randomness of the profile heights and of the "sharpness" of the surface height distribution, has a value of ~3, which is characteristic for a near perfect random surface (Gaussian distribution) [30].

Li-0 films (Figure 1–second row) exhibit protruding (*R*sk ~0.46) large particles, randomly distributed (*R*ku ~3.11), having tens of nm in diameter (~35–50 nm), individually displayed or with a tendency to merge. In this case, the average roughness increases to a value of ~5.9 nm, which is higher than the one registered for the bare substrate.

In the case of Li-C structures (Figure 1–third row), the addition of lithium leads to the formation of much larger particles, almost double as compared to Li-0 films and, accordingly, with a double roughness value (10.3 nm). Though, particles are not much more protruding in z direction, as the *R*sk parameter is ~0.56, and relatively uniform distributed as the kurtosis parameter being just a bit higher (*R*ku ~3.49) than in the second case.

Finally, in the case of Li-P structures (Figure 1–last row), the addition of lithium leads to a morphology similar to the one of Li-0 films, but with particles slightly smaller (diameters ~25 nm) and better separated (the fusing tendency being lower). Therefore, it exhibits the highest value for *R*ku parameter in this series (5.58), pointing to a sharp height distribution [34], but with a lower skewness factor (0.17) as compared with Li-0 (0.46) and Li-C (0.56), which means that particles are less

protuberant for Li-P structures. This leads to an average roughness value (4.32 nm) very close to the one characteristic for Li-0 films (5.9 nm).

Overall, based on the results, one should emphasize that the obtained lithium-based structures could be described as homogeneous and uniform, with a morphology that depends on the preparation route and with (average) roughness values in the nanometric range (*R*a ~5–10 nm).

The inferred values of the amplitude parameters, which were discussed above, are summarized in Table 2.

**Figure 1.** The 2D (classical view mode—left column) and 3D (enhanced color view mode—right column) AFM images of Li-0, Li-C and Li-P structures synthesized at 600 ◦C on Si(100) wafers.

**Table 2.** Surface amplitude parameters (*R*a, *R*sk and *R*ku) for a (0.5 × 0.5 μm2) scanned surface area of control (Si) and Li-0, Li-C and Li-P structures.

#### *3.2. Biocompatibility Assays*

#### 3.2.1. Cytotoxicity Test

In the current study, the cytotoxicity of obtained coatings was assessed using three cellular lines of human origin, selected for covering their diverse potential applications in the development of medical devices. Thus, MG-63 osteosarcoma cells exhibit common features with human osteoblast precursors or poorly differentiated osteoblasts, producing type I collagen with no or reduced osteocalcin and alkaline phosphatase basal level, making them appropriate for assessing the biocompatibility of biomaterials designed for orthopedic applications [35,36]. The other two cell lines were represented by human skin new initiated dermal fibroblasts and human skin immortalized keratinocytes, considering that many implanted medical devices or next generation wound dressings are coming for short time into contact with the skin.

Concerning the cytotoxicity of synthesized materials, contrast phase microscopy images (Figures 2–4) have shown that cells developed on plastic well, at the limit with the tested disc surface, are unaffected and preserved an unmodified morphology. This suggests that BHA materials do not release toxic substances in the environment, as secondary product. New initiated dermal fibroblasts exhibited a spindle-shaped morphology (Figure 2), the morphological appearance of keratinocytes showed polygonal cells (Figure 3), whilst MG-63 monolayers were composed of elongated and fusiform cells (Figure 4).

**Figure 2.** Contrast phase inverted microscopy images of the new initiated dermal fibroblasts monolayer grown in the presence of Ti and simple and doped biological-derived HA coatings (Magnification ×200).

**Figure 3.** Contrast phase inverted microscopy images of the immortalized keratinocytes HaCaT monolayer grown in the presence of Ti and simple and doped biological-derived HA coatings (Magnification ×200).

**Figure 4.** Contrast phase inverted microscopy images of the MG-63 osteosarcoma cells monolayer grown in the presence of Ti and simple and doped biological-derived HA coatings (Magnification ×200).

On the other hand, the fluorescence microscopy images (Figures 5–7) have revealed that the number of adhered cells on the bare Ti was similar between all synthesized layers and depends on type of cell line. This observation suggests that deposited layers are not cytotoxic against human skin or bone cells and allow for an unhindered cell development. According to the ISO standard specifications [32], the cytotoxic effect of synthesized layers can be classified as absent, supporting the potential of fabricated materials to be used as coatings for different biomedical surfaces, such as orthopedic devices, dental materials, or catheters.

**Figure 5.** Fluorescence microscopy images of new initiated dermal fibroblasts grown for 24 h on the surface of Ti and simple and doped biological-derived HA coatings (Magnification ×200).

**Figure 6.**

**Figure 7.**  Cellular Cycle Assay

The cell cycle analysis by flow cytometry revealed that the obtained coatings did not influence the cellular cycle progression of the three type of cells (i.e., new initiated dermal fibroblasts, immortalized keratinocytes and MG-63 osteosarcoma cells) (Figures 8–10), confirming that synthesized layers do not interfere with the typical development of normal and tumoral cells.

**Figure 8.** Flow cytometry diagrams of cells cycle analysis of new initiated dermal fibroblasts grown on the tested materials.

**Figure 9.** Flow cytometry diagrams of cells cycle analysis of immortalized keratinocytes grown on the tested materials.

**Figure 10.** Flow cytometry diagrams of cells cycle analysis of MG-63 osteosarcoma cells grown on the tested materials.

#### *3.3. Anti-Biofilm Activity*

3.3.1. Anti-Staphylococcal Biofilm Activity of Obtained Coatings

Compared with Ti, after 24 h of incubation, all tested coatings exhibited a significant inhibitory activity against *S. aureus* biofilm growth, showing a significant increase starting with 48 h of incubation and persisting with high intensity, after 72 h of incubation. This behavior is suggested by the drastic decrease of the number of CFU/mL by 2 to 4 logs at these last two-time intervals (Figure 11b). The

absence of a visible inhibition of the microbial growth in the 1–2 h (Figure 11a) interval reveals that the anti-biofilm effect is possibly due to the slow release of some bioactive components from the synthesized coatings, with potential antimicrobial activity (e.g., oligoelements).

**Figure 11.** Graphic representation of the logarithmic values of *S. aureus* biofilm cells developed on the surface of Ti and simple and doped biological-derived HA coatings: (**a**) 1–12 h and (**b**) 24–72 h.

Regarding comparative efficiency of the tested materials, one can observe that Li-C coatings exhibited a slightly increased antimicrobial activity against *S. aureus* biofilm after 48 h of incubation, as compared to Li-0 and Li-P ones (Figure 11b).

The SEM images of *S. aureus* biofilm developed on the surface of Ti and simple and doped BHA structures are presented comparatively in Figure 12, at a magnification of 10,000×. For comparison reasons only non-inoculated simple and lithium-doped surfaces are presented by the authors of [21]. Although *S. aureus* adhered to all tested samples, it must be noticed that the density of *S. aureus* cells was significantly more reduced on Li-C > Li-P > Li-0 surfaces as compared to Ti ones, which is in good agreement with the quantitative results presented in Figure 11. SEM images of *S. aureus* biofilms are in concordance with others reported in the literature [1,37,38].

**Figure 12.** SEM images corresponding to *S. aureus* biofilm after 72 h of development on the surfaces of Ti and simple and doped biological-derived HA coatings.

#### 3.3.2. Anti-Fungal Biofilm Activity of Obtained Coatings

In the case of *C. albicans* strain, a strong anti-biofilm effect was recorded for all tested coatings as compared to Ti, earlier than in the case of *S. aureus* biofilm, i.e., after 12 h of incubation (Figure 13a), followed by a recrudescence of microbial growth at 24 h and a drastic decrease of the biofilm-embedded cells at 48 and 72 h, respectively (Figure 13b). Moreover, Li-C and Li-P coatings proved a slightly better efficiency (as compared to Li-0 samples) against *C. albicans* biofilms incubated for 48 h, whilst Li-0 and Li-P showed the same antimicrobial efficiency at 72 h of incubation (Figure 13b).

**Figure 13.** Graphic representation of the logarithmic values of *C. albicans* biofilm cells developed on the surface of Ti and simple and doped biological-derived HA coatings: (**a**) 1–12 h and (**b**) 24–72 h.

The efficiency of the anti-fungal biofilm activity of simple and doped coatings decreased in the order Li-C, Li-P, Li-0, after an incubation time of three days (Figure 13).

Similar to the case of *S. aureus* strain, no visible inhibition of microbial development was obtained in the range of 1–6 h (Figure 13a), most probably due to the slow release from the synthesized films of the active elements responsible for the antimicrobial effect.

Following adherence to abiotic substrates (i.e., catheters and other medical devices), *C*. *albicans* develops into a three-dimensional community containing round-budding yeast cells, oval pseudo-hyphal cells and elongated hyphal cells embedded in an extracellular polysaccharides' matrix [39]. *C. albicans* biofilm development follows a sequential dynamic over a 24–48 h period, starting with initial adhesion of single fungal yeast cells, continuing with a cellular proliferation and colonization of the surface and the switching to filamentous hyphal forms, a hallmark of the initiation of biofilm formation, followed by maturation, through the accumulation of the protecting extracellular polysaccharide matrix. Dispersion of biofilm-associated cells from the mature biofilm can initiate formation of new biofilms or disseminate into host tissues, therefore being associated with candidemia and invasive fungal infections [40].

SEM images of *C. albicans* biofilm developed on the surface of Ti and simple and doped BHA structures are presented comparatively in Figure 14 at a magnification of 10,000×. In the case of bare Ti, one can observe many ovoidal or round-shaped cells which have the tendency to form micro-colonies, as well as the existence of filamentous, hyphal forms. Ti samples were completely covered by a continuous layer of cells, which is in good correlation with the quantitative results presented in Figure 13. This observation might be indicative not only for cell-surface type interactions, but also for stronger cell-cell type ones.

**Figure 14.** SEM images corresponding to *C. albicans* biofilm after 72 h of development on the surfaces of Ti and simple and doped biological-derived HA coatings.

In exchange, in case of coated surfaces, particularly Li-C and Li-P ones, the number of yeast cells is drastically decreased, they adhere exclusively in isolated, unicellular form, the filaments being completely absent.

The microscopy results demonstrate that the synthesized coatings inhibit the initiation of fungal biofilm development. SEM images of *C. albicans* are in concordance with others reported in literature [1,41].

We note that results of anti-biofilm quantitative assays, which revealed a slightly superior anti-biofilm effect of Li-C and Li-P structures as compared to control Ti and Li-0 ones, are in good agreement with aspects recorded in SEM examination (Figures 12 and 14). Presented SEM images were considered representative for the entire surface of samples and were selected after examining 8–10 different microscopic fields.

#### **4. Discussion**

#### *4.1. Biocompatibility*

Lithium is considered as a trace metal in organisms and is used for the treatment of bipolar affective disorder [42] and refractory depression [43]. The safe administration of lithium for osteoporosis was reported more than two decades ago [44,45]. When used in the low bone mass and osteoporosis-mice models, lithium was indicated to enhance osteogenesis [46]. Also, in the case of lumbar spine, femoral neck and proximal femur, the corresponding bone mineral density values were reported to be higher than controls [45,47]. Despite its proven efficacy and safety [48], there are cases when lithium administration could be toxic and determine side effects, such as thirst and excessive urination, nausea, diarrhea, tremor, weight gain, sexual dysfunction, dermatological effects, as well as long-term effects on the thyroid gland, kidneys and parathyroid glands [49]. However, recent studies challenge this negative perception of lithium administration, by showing strong evidence that intoxication in patients receiving lithium treatments is rare, it occurs only when the concentration is higher than 1.5 mmol/L, and its main adverse effects can be properly monitored and managed [50–52].

Even though its excellent biocompatibility was proven, the poor mechanical properties of HA limit its applications in implantology to non-load-bearing implants. It has been shown that ion doping represents a good method applied to improve performances of bio-ceramics related to biodegradability, biocompatibility, strength and toughness. Therefore, when lithium is substituted into the HA structure, it plays an important role because it could induce a decrease in HA solubility [53,54], without affecting its biocompatibility [55,56], and an improvement of biomechanical properties, without altering the HA structure. The reason for the increase of strength of lithium doped HA may be due to the fact that metal ions could decrease the porosity of samples which can result in a more compact and hard structure [57]. Also, one notes that, as a substituted trace element, lithium can positively influence the viability [58], proliferation rates and alkaline phosphatase activity of cells [59,60]. Wang et al. [57] reported that when used in a proper low dose (0.5% and 1%), lithium ions can substitute calcium ones in a HA structure, in order to obtain a biomaterial (LiHA scaffold) characterized by high strength, a higher osteoblast-mediated degradation rate and a superior ability to promote cell proliferation as compared to simple HA and controls. On one hand, the lithium incorporation into HA can change the surface topography of original HA material and make LiHA scaffold possess a more compact bulk, which finally can result in a higher adherence and enhanced growth of osteoblasts [61]. On the other hand, the Li+ released from LiHA scaffolds could enhance the activity of osteoblasts, a phenomenon also reported by Tang et al. [62]. In another study, Oh et al. [63] assessed the influence of lithium fluoride (LiF) on in vitro biocompatibility and bioactivity of calcium aluminate-polymethylmethacrylate composite cements. The results indicated that there was no cytotoxicity related to LiF and the addition of 1 wt.% LiF enabled composite cements to form HA on their surface within 15 days of soaking in SBF solutions. One can therefore conclude that these enhanced characteristics strengthen the safe use of lithium and could advance lithium-containing structures as promising materials for the fabrication of bone tissue engineering scaffolds.

The literature data on the cytotoxic ions release rates for lithium-substituted HA materials is rather scarce and reported cytotoxic levels are sometimes contradictory due to the fact that release rates are strongly depending on crystallinity or particle morphology and size of the investigated material [42]. In a current study, Miguez-Pacheco et al. [64] reported on lithium-containing bioactive glasses, with Li2O substituting Na2O in different amounts. The conclusion of their study, in terms of lithium ion release, was that formulations with 2.5 wt.% and 5 wt.% Li2O content proved to be within the therapeutic range.

In vitro cytotoxicity assays of newly developed materials should precede in vivo tests, since the results are more reproducible and allow for the evaluation of cellular and molecular responses triggered by an investigated material [65].

We stress upon that results obtained on human skin new initiated dermal fibroblasts and human skin immortalized keratinocytes could be informative for the potential of the tested materials to be used in oral applications, for example in facial reconstruction, dental implants or periodontal pockets filling. A fully differentiated human gingiva equivalent was constructed entirely from keratinocytes and fibroblasts, being considered a valuable surrogate of animal models for testing new therapeutics [66]. Moreover, one could therefore predict the potential orthopedic implant-related adverse effects, related to the occurrence of hypersensitivity reactions to the implant material components. This evaluation is presently recommended for patients with a clear self-reported history of metal reactions and is done by patch testing before device implant [67].

#### *4.2. Possible Mechanisms Explaining the Antimicrobial Activity of Synthesized Structures*

We stress that infection still remains a major impediment to the long-term use of many implanted devices. Even though microbial adhesion to biomaterial surfaces represents an essential step in the pathogenesis of these infections, molecular and physical interactions that govern this adhesion have not been yet fully understood. Both specific and non-specific interactions may play an important role in the ability of cells to attach to (or to resist detachment from) the biomaterial surface [68]. In the biomedical field, a great number of implant failures is due to the formation of microbes at the implanted site. Microbes, which are able to cause a wide variety of infections in humans by spreading through common places like bathroom tiles, doorknobs, or packing materials, can be controlled by coating implants

with materials with antimicrobial properties. Antimicrobial implantable materials must locally kill microbes or slow down their growth, without being toxic to host tissue cells. Therefore, if an implant material will possess such characteristics, then the problem of failure will be drastically reduced.

Besides the utility of lithium in the treatment of mood disorders, different studies reported that as well as other antidepressants, it could exhibit an immunostimulant role and antimicrobial activity, as it opposes eicosanoids suppressive action upon the immune system [69]. Recent studies have shown that lithium could limit neurodegeneration and ameliorate symptoms of the trimethyltin poisoning in rodents, by inducing pro-survival and anti-inflammatory effects [70]. Moreover, in concentrations of up to 10 mol.%, lithium could enhance the antibacterial activity of bioactive glass against resistant bacteria, such as methicillin-resistant *S. aureus*. The 58S bioactive glass with 5 mol.% Li2O substitution for CaO was considered a promising biomaterial in bone repair/regeneration therapies, showing an enhanced biocompatibility and alkaline phosphatase activity, with a negligible loss in the bioactivity compared to the 58S bioglass [71].

The adhesion ability of *S. aureus* to extracellular matrix components and plasma proteins from the pellicle formed on the surface of implanted biomaterials represents one main concern for infections associated to implanted medical devices. *C. albicans* is considered one of the major etiological agents of hospital-acquired infections, being also frequently associated with contamination of indwelling medical devices [72]. The phenotypic plasticity is a unique feature of many fungal species, including *C. albicans,* which is an extremely versatile, dimorphic yeast that can survive as a commensal in various environments, one of the adapting mechanisms being the switching from the yeast to filamentous forms [73]. *S. aureus* and *C. albicans* can be also found in multispecies biofilms, mutually potentiating their tolerance to antibiotics and disinfectants [74]. The mechanism of indirect pathogenesis, which might lead in the end to therapeutic failure, could also occur inside multispecies biofilms, in which genetically resistant microorganisms are the susceptible pathogens [75]. As an example, the coexistence of *S. aureus* and *C. albicans* inside biofilms determines the enhancement of *S. aureus* resistance to vancomycin [76]. Therefore, the decrease of microbial adhesion represents one of the main concerns of the current research in the field of biomaterials and medical devices.

In the following, we propose some relative contributions of specific and nonspecific mechanisms in order to explain, to some extent, the efficiency manifested by our fabricated coatings to prevent the microbial adhesion. Therefore, surface features like wetting behavior and material surface characteristics (topography), corroborated with metal ions release can stand as key-factors able to modulate the adhesion of microbes and their subsequent survival on the surface of the synthesized structures.

Water is considered as a major interfacial host in all living environments, and it can affect the capacity of protein adsorption [77] and cellular adhesion [78]. Therefore, to interpret the microbial behavior, this parameter must be carefully taken into consideration. In this respect, one possible mechanism able to explain the antimicrobial activity of synthesized structures is based on their wetting properties and, consequently, on their surface energy values [38]. Therefore, a hydrophilic behavior and high values of surface energy [21] might enhance the overall antimicrobial properties of our fabricated coatings. In addition, the anti-biofilm activity might be explained by repulsive forces which will occur between anionic-coated surfaces (associated with the presence of hydroxyl hydrophilic functional groups) and the *S. aureus*/*C. albicans* biofilms matrix, which is also negatively charged, due to its main component, represented by self-produced exopolysaccharides [38]. Moreover, coatings will facilitate the formation of a highly alkaline medium, consequently leading to the destruction of lipids, which are the main component of microbial cell membrane, followed by microbial cells lysis [79]. When discussing anti-biofilm properties of obtained coatings, one must also consider that the adhesion of microbial cells from aqueous suspensions follows the thermodynamic model. The surface energy of microbial cells is typically smaller than the surface energy of liquids in which cells are suspended, causing cells to attach preferentially to hydrophobic materials (i.e., materials with lower surface energies). Therefore, a hydrophilic behavior, such that of our coatings, and the subsequent

formation of hydration layers could disrupt the initial step of microbial biofilm formation, represented by the adhesion of microbial cells to a surface [80].

Another factor to explain the adhesion of microorganisms is related to the surface characteristics of the material. Therefore, the reduction of microorganisms can be investigated as a function of surface topographic parameters. Among the surface properties, surface roughness and topography have been the most investigated topics in dental biofilm researches. There exists a general perception according to which an increase in surface roughness can promote bacterial colonization due to the increase in contact area between the material surface and bacterial cells, in other words, due to extended surface area available for attachment [81]. Thus, smoothening the surfaces to nanometer-scale roughness was demonstrated to reduce biofilm formation [82]. We note that a roughness (*R*a) value of 0.2 μm was reported to be the threshold for maximum reduction of bacterial adhesion on abutment surfaces [83]. In this respect, *R*a values inferred in our cases, which are (19–46) times smaller than this threshold, should be emphasized.

Moreover, it was recognized that *R*ku and *R*sk are parameters of key-importance when describing surface roughness of calcified tissue implants [84] because osteogenic cells' adhesion is sensitive to the stress concentrations which derive from the presence of sharp heights at the implant-tissue interface [85], likewise they can be considered against bacterial proliferation [86]. Inferred values for skewness and kurtosis parameters indicated a sharper heights (*R*ku ~5.6) distribution of surface features in case of Li-P structures as compared with Li-0 and Li-C ones (*R*ku ~3.1 and ~3.5), with more peaks than valleys (*R*sk > 0) and that these peaks and valleys are much closer in height and depth for Li-P coatings (*R*sk ~0.17) than those of Li-0 and Li-C (*R*sk of 0.46 and 0.56, respectively). Thus, the average roughness parameter was found to have a similar value (~5 nm) for Li-0 and Li-P coatings, and a double one (~10 nm), for Li-C structures. The positive values of *R*ku and *R*sk coefficients could also be related to less available surface-active sites [78,80], and, consequently, could stay at the origin of the inhibition of microbial development on the surface of synthesized structures.

Last but not least, there are reports in the literature on the antibacterial potential of antibiotics/antibacterial substances-loaded HA, but very few refer to antifungal-loaded cement drug delivery [87]. We stress upon that there are even fewer reports dedicated to the antibacterial activity of lithium [88]. It was demonstrated that the antibacterial/anti-biofilm effect can become more evident when extending the exposure to metal [89]. Therefore, one of the proposed mechanisms of action could be the metal accumulation inside bacterial cells, followed by the cellular wall disruption and, eventually, cellular lysis. By extending the proposed mechanism, we can speculate that the amount of lithium ions released from our synthesized materials was enough to accumulate in the microbial cells and to alter the culture media. By affecting their multiplication and phenotypic switching capacity, their further progress is slowed down. These assumptions were confirmed by the results of our study, the lithium-doped samples showing a more pronounced inhibitory effect after 48 and 72 h of incubation, respectively.

#### **5. Conclusions**

Simple and lithium-doped biological-derived hydroxyapatite (BHA) coatings were fabricated by Pulsed Laser Deposition onto medical-grade titanium substrates.

In vitro biocompatibility investigations proved that synthesized layers were non-cytotoxic on all tested cell lines, respectively MG-63 osteosarcoma derived cells, human skin new initiated dermal fibroblasts and immortalized keratinocytes, and did not interfere with their cellular cycle phases. In addition, synthesized structures drastically inhibited the ability of *S. aureus* and *C. albicans,* two microorganisms well-known for their complex genetic resistance mechanisms and involvement in the etiology of medical devices-associated biofilm infections, to develop mature biofilms. The anti-fungal biofilm activity has become significant after 12 h of incubation in the case of *C. albicans* and after 24 h for *S. aureus* biofilm, and persisted with high intensity for 72 h. The anti-biofilm activity of obtained coatings could be explained in terms of their surface characteristics, such as hydrophilic character, associated to a high value of the surface energy, or low values of surface roughness (in the range of 5–10 nm).

We consider that these combined characteristics can be important for the protection of biomedical equipment against contamination by nosocomial infections of any sort, as well as for the long life exploiting of various implants and could therefore offer guidance towards the suitability of lithium-doped BHA structures for future biomedical applications.

**Author Contributions:** Conceptualization, L.D. and A.P.; Data Curation, M.C.C., G.P.-P., M.A. and A.A.; Investigation, M.C.C., G.P.-P., C.B., G.P.G. and M.A.; Methodology, L.D. and A.P.; Supervision, L.D. and A.P.; Validation, L.D. and A.P.; Writing—Original Draft, L.D.; Writing—Review and Editing, L.D., M.C.C. and A.P.

**Funding:** This work was supported by a grant of Ministry of Research and Innovation, CNCS-UEFISCDI, No. PN-III-P1-1.1-PD-2016-1568 (PD 6/2018), within PNCDI III. LD, GPP and ACP acknowledge with thanks the partial support of the Romanian National Authority for Scientific Research and Innovation, CNCS-UEFISCDI, under No. PN-II-RU-TE-2014-1570 (TE 108/2015) and Core Programme–Contract 3N/2018. LD and ACP acknowledge the partial support of the Romanian National Authority for Scientific Research and Innovation, CNCS—UEFISCDI, under No. PN-III-P2-2.1-PED-2016-1309 (Contract No. 241/2017).

**Acknowledgments:** All authors thank F.N Oktar for providing simple biological-derived HA powders and G.E. Stan and V. Craciun for useful discussions.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

© 2019 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (http://creativecommons.org/licenses/by/4.0/).

## *Article* **Glancing Angle Deposition of Zn-Doped Calcium Phosphate Coatings by RF Magnetron Sputtering**

#### **Konstantin A. Prosolov 1,2,\*, Olga A. Belyavskaya 1, Juergen Linders 3, Kateryna Loza 4, Oleg Prymak 4, Christian Mayer 3, Julietta V. Rau 5, Matthias Epple 4 and Yurii P. Sharkeev 1,2**

Received: 4 March 2019; Accepted: 25 March 2019; Published: 28 March 2019

**Abstract:** Zn-substituted hydroxyapatite with antibacterial effect was used in radiofrequency (RF) magnetron deposition of calcium phosphate coating onto Ti- and Si-inclined substrates. The development of surface nanopatterns for direct bacteria killing is a growing area of research. Here, we combined two approaches for possible synergetic antibacterial effect by manufacturing a patterned surface of Zn-doped calcium phosphate using glancing angle deposition (GLAD) technique. A significant change in the coating morphology was revealed with a substrate tilt angle of 80◦. It was shown that an increase in the coating crystallinity for samples deposited at a tilt angle of 80◦ corresponds to the formation of crystallites in the bulk structure of the thin film. The variation in the coating thickness, uniformity, and influence of sputtered species energy on Si substrates was analyzed. Coatings deposited on tilted samples exhibit higher scratch resistance. The coating microand nano-roughness and overall morphology depended on the tilt angle and differently affected the rough Ti and smooth Si surfaces. GLAD of complex calcium phosphate material can lead to the growth of thin films with significantly changed morphological features and can be utilized to create self-organized nanostructures on various types of surfaces.

**Keywords:** RF magnetron sputtering; plasma-assisted deposition; biocompatibility; GLAD; ion-substituted apatites

#### **1. Introduction**

The demand for implants that can efficiently conduct bone defect regeneration increases significantly with the aging population [1]. It is known that the bioinert metals that are used in implantology and orthopedics do not provide the desired bioactivity and osteoconductivity; therefore, the surface of such medical devices is usually modified with bioactive coatings [2]. Moreover, the demand for antibacterial or bactericidal surfaces and coatings has increased steadily in recent years. It has been noted that the modern approach to surface modification should take into account both coating functions: osteoconduction properties and antibacterial effect [3]. From that perspective, substituted hydroxyapatites (HAs) are one of the most promising materials in this area of research. It has been shown that the easiness of atomic doping or substitution in HA reveals its potential to be used in various clinical cases [4–6]. Most importantly, it has been shown that inorganic antibacterial metallic ions (e.g., Ag, Cu, Zn, Sr) can be introduced into the HA lattice and improve its antibacterial properties [7–14]. However, there is a physical approach to antibacterial surface production. It has been shown that the creation of nanopatterns is a powerful tool for directing stem cell fate. Another advantage is that high aspect ratio nanopatterns are capable of killing bacteria and preventing biofilm formation [15].

Modern approaches to healthcare are aiming to produce implants with biomimetic properties [16]. These properties are crucial to ensuring a desirable biological response to the newly implanted material, in such a manner that the cells, which are adhered to the surface of such scaffolds, can function in a way that is similar to physiological conditions. From that point of view, the formation of a coating that consists of different types of surface gradient structures and with variations in the level of roughness on the submicro- and nanoscale could be of significant interest for biomimetic purposes [17,18]. The possibility of manipulating nanotopography has attracted much attention in recent years [19–22]. This has fostered the development of film deposition techniques aimed at enhancing these morphological characteristics.

In the field of optics and microelectronics, physical vapor deposition (PVD) of thin films, allowing the deposition of porous and/or columnar-like structured coatings, has been available for some years. For this purpose, the use of an oblique angle geometrical configuration—or as it is also known, glancing angle deposition (GLAD) method—is frequently exploited [23–27].

The GLAD configuration has emerged as a promising tool for the deposition of nanostructured thin films. This method has mostly been applied in the field of metallic coatings or dielectric coatings for optical systems, and the use of such technology in the field of biomaterials started recently [28]. Such coatings are usually represented by inclined columnar-like structures with variations at the nano-roughness level and a direction of growth allowing for the deposition of patterned structures with different shapes and sizes. GLAD has a wide range of applications in various fields, such as sensors [29], optics [30,31], or solar cells [32]. This type of coatings has attracted attention due to the anisotropic nature of the formed films, their peculiar mechanical properties, and the possibility of controlling the level of porosity [33] and roughness [34]. Most researchers who have utilized the GLAD method have used line-of-sight methods of deposition, such as, for example, thermal evaporation. However, the effect of self-shadowing is applicable in many PVD techniques, such as magnetron sputtering.

An emerging method for bioactive coating deposition in the field of PVD is radiofrequency (RF) magnetron sputtering. Magnetron sputtering is widely used in the formation of coatings for various applications. The continuous interest of scientists in this method is due to the possibility of modifying the coating structure and its physicochemical properties by varying the deposition parameters [2]. There is a significant interest in radiofrequency (RF) magnetron sputtering of bioactive calcium phosphate (CaP) thin films [35,36]. This method allows for the deposition of CaP coatings with a high level of adhesion to the substrate. Moreover, RF magnetron sputtering is well known as a highly controllable deposition method. However, to date, a feasible method of calcium phosphate coating deposition in GLAD geometry via RF magnetron sputtering has not yet been reported.

Here, we aimed to develop a method for depositing thin Zn-substituted HA (Zn-HA)-based coatings that are formed by hierarchical gradient surface features with the use of GLAD geometry. This method allows us to manipulate the coating roughness on the submicron and nanoscale level and pursue the possible synergetic antibacterial effect of a nanopattern and the antibacterial activity of Zn.

#### **2. Materials and Methods**

Commercially pure titanium samples (99.58 Ti, 0.12 O, 0.18 Fe, 0.07 C, 0.04 N, and 0.01 H wt %) of 10 × 10 × 1 mm3 size provided by "VSPO-AVISMA" (Perm, Russia) were used as substrates. The samples were sequentially polished by silicon carbide paper of 120, 480, 600, and 1200 grit. Prior to deposition, the substrates were cleaned in an ultrasonic bath of distilled water for 10 min. Furthermore, commercially available Si wafers (1 0 0) provided by "ELMA" (Zelenograd, Russia) were used as substrates in all the deposition runs.

In Figure 1, the deposition process is schematically shown. The samples were mounted on the custom-made sample holder at different angles to the flux of the sputtered material.

**Figure 1.** Schematic image of the deposition process of the Ti and Si samples with varying particle incidence angles at 0◦, 60◦, and 80◦ during the radiofrequency (RF) magneton sputtering of a Zn-substituted hydroxyapatite (Zn-HA) target with the arrows indicating the direction of the flux.

The sample holder allows for the simultaneous placing of substrates at 0◦, 60◦, or 80◦ to ensure the same sputtering conditions for all the deposited samples.

The sputtering target was Zn-substituted hydroxyapatite. This material was chosen because hydroxyapatite resembles the mineral fraction of bone and is used for many implant coatings, while the addition of Zn can provide faster bone growth together with an antibacterial effect [37]. A Zn-HA powder was prepared by mechanochemical synthesis at the Institute of Chemistry and Mechanochemistry, Russian Academy of Sciences, Novosibirsk, Russia. The mechanochemical synthesis was carried out according to the reaction:

$$6\text{CaHPO}\_4 + (4-x)\text{CaO} + x\text{ZnO} = \text{Ca}\_{(10-x)}\text{Zn}\_x(\text{PO}\_4)\_6(\text{OH})\_2 + 2\text{H}\_2\text{O}, \text{ where } x = 0.4.$$

The prepared powder was used as a precursor for the preparation of a target for sputtering. Its detailed characterization can be found in our previous report [38]. The powder was pressed and then sintered in air at 1100 ◦C for 1 h. The chemical composition of the powder and the target was confirmed by X-ray powder diffraction (XRD) and Fourier transform infrared spectroscopy (FTIR) [39].

To identify the optimal deposition parameters for our experiment, we reviewed the literature dealing with GLAD, considering the substrate material, sputtered material, deposition parameters, and type of formed coatings. In Table 1, the deposition parameters for GLAD of various materials and their properties are summarized.

**Table 1.** Deposition parameters and coating properties for glancing angle deposition (GLAD) with magnetron sputtering from the literature.

*L*–target–substrate distance; *t*–deposition time; *P*–pressure; *W*–power.

Even though most GLADs were performed at target-to-substrate distances larger than 70 mm, we decided to perform deposition at a distance of 45 mm, as it is consistent with most published research papers dealing with sputtering of CaPs [47–49] and ensures that the particle-free mean path is significantly higher than the target-to-substrate distance. Moreover, it is known that ZnO and hydroxyapatite share the same type of lattice structure (hexagonal system). Taking into account the fact that there is a correlation between the lattice type of sputtered material and the type of the formed coating structure, we hypothesized that using similar deposition parameters to those used for ZnO would result in the formation of Zn-doped calcium phosphate nanocolumns [42,46,50].

A vacuum installation with a planar magnetron operating at a frequency of 13.56 MHz (TPU, Tomsk, Russia) was utilized for GLAD. The high-quality Zn-HA target was used as the cathode of the magnetron. The sputtering power density was set to 3.15 W/cm2. Prior to the deposition, the system was pumped to a base pressure of 5 × 10−6 Torr, and then, Ar gas was passed through at a controlled rate to maintain the pressure needed for the experimental deposition. The deposition time was set to 3 h, and the target-to-substrate distance was 45 mm. It is important to note that the target-to-substrate distance varied for tilted samples due to the inclination; therefore, the upper part of the sample was as close to the surface of the magnetron as 35 mm. The samples were deposited at a working gas pressure of 0.8 mTorr at a gas flow of 3.7 sccm. Each set of samples with angles varied from 0 to 80◦ was deposited in a single run.

The thicknesses of the deposited films were measured using spectroscopic ellipsometry with an ELLIPS-1891 SAG setup (Rzhanov Institute of Semiconductors Physics of SB RAS, Novosibirks, Russia). The surface morphology and composition of the coatings were examined with scanning electron microscopy using an ESEM Quanta 400 FEG scanning electron microscope (SEM, FEI Company, Hillsboro, OR, USA) operating in a high vacuum. The samples were coated with gold–palladium for 15 s before the SEM, which was equipped with energy-dispersive X-ray spectroscopy (EDX; Genesis 4000, SUTW-Si(Li)detector, EDAX, Inc., Mahwah, NJ, USA) and was operating in a high vacuum study in such a way that no alteration of the surface morphology was induced. The estimated thickness of the deposited Au–Pd layer was 4–5 nm. The acceleration voltage for the SEM was 20 kV. To determine the level of crystallinity of the studied samples and estimate the crystallite size of the deposited thin CaP coatings, a Panalytical Empyrean X-ray diffractometer with a Cu Kα radiation source (λ = 1.54 Å; 40 kV and 40 mA) was used. The coatings were investigated by grazing incidence X-ray diffraction (GIXRD) with an incident beam angle of Ψ = 1.0◦ (with respect to the sample surface) and at 2θ range from 5◦ to 90◦ with a step size of 0.05◦. For qualitative phase analysis with the Bruker software Diffrac.Suite EVA 4.2.2, the patterns of hydroxyapatite (#09-0432) and titanium (#44-1294) from the ICDD database were used as a reference. After the instrumental correction derived from a standard powder sample LaB6 from the NIST ([51]; a(LaB6) = 4.15689 Å), quantitative Rietveld analysis for the calculation of the lattice parameters and the crystallite sizes was performed with the Bruker software TOPAS 5.0. For the morphology visualization and calculation of the surface roughness level of the samples, a JPK NanoWizard Atomic Force microscope (AFM, JPK, Berlin, Germany) was used. The coating-to-substrate adhesion evaluation was carried out with the scratch-test method on a CSM Macro Scratch Tester Revetest (CSM Instruments, Needham Heights, MA, USA) with an indenter of 20 μm in radius. The maximum indentation load was 10 N. The scratch length was set to 5 mm. In order to obtain statistically meaningful data, each measurement was repeated at least three times per sample.

#### **3. Results and Discussion**

#### *3.1. Effect of GLAD on Coating Thickness and Its Uniformity along the Substrate Surface*

The key for variation in the coating surface topography when the GLAD method is used lies in the self-shadowing effect [23]. During the film growth, adatoms migrate to the substrate and encounter the surface in a way that causes peak and valley regions to grow. Due to the inclination of the substrate, the peaks receive more atoms compared with the valleys. As the peaks grow higher, a shadow region, which receives a lower number of atoms, is formed behind them. Consequently, the GLAD coatings have a preferential direction of growth and may even be porous. Also, in most of the reports dealing with GLAD, an increase in root-mean-square (RMS) surface roughness is usually observed when compared with the normal degree deposition [25,26,29].

Even though research on GLAD coatings is evolving extensively, there are few reports on the coating uniformity of the substrate surface and the variation in its thickness. Here, we show the thickness variation on the Si substrates measured by ellipsometry. In Figure 2, the coating thickness of the substrates is placed in parallel to the target of the magnetron (0◦), and inclination angles of 60◦ and 80◦ are presented.

**Figure 2.** Plot of the coating thickness against the substrate tilt angle.

With the increase in the inclination angle, the coating thickness decreases gradually. We believe that this effect is associated with substrate re-sputtering. It is known that when the substrate is in close proximity to the plasma source, impinging energetic ions of Ar+ and O− induce re-sputtering of the coating from the substrate [28]. Moreover, the deposition parameters were chosen in such a way that the mean free path of the sputtered atoms was as large as possible while maintaining a relatively high power density. In this way, the kinetic energy of the sputtered material and the generated ions is close to the original level due to the negligible number of collisions they experience. It not only enables better adatom mobility, but also higher substrate heating and hence higher re-sputtering. In addition, a shorter target–substrate distance will increase the rate of deposition but at the expense of film uniformity [29]. This was the case for the CaP coatings deposited in our experiments. For the samples deposited at less than an inclination angle of 80◦, the coating thickness was measured for the upper part (closer to the magnetron source) and the lower part of the sample. Ellipsometry showed that the coating thickness was inhomogeneous along the sample, and the thickness difference between the upper and lower part of the coating was up to 110 ± 25 nm for some samples. Interestingly, the closer the placement of the sample to the plasma source, the thinner the film deposited. We believe that this effect is due to substrate re-sputtering and energetic ions bombardment.

#### *3.2. Effect of GLAD on Coating Topography and Roughness*

In Figure 3, SEM images show the topography of CaP coatings deposited at different angles of 0◦, 60◦, and 80◦. The coatings deposited without tilting (0◦) show globular-like surface features. The samples inclined at 60◦ during the deposition decreased in size, and the surface structure elements changed their shape. The surface grains were orientated along the incidence direction of the sputtered particles independent of working gas pressure. This effect was even more pronounced for the samples deposited at a tilt angle of 80◦. An alteration in the working gas pressure led to a reduction in surface

elements sizes. When the deposition process was performed at a lower working pressure (here: 0.8 mTorr), surface elements tended to be more homogeneously spread on the surface.

**Figure 3.** Calcium phosphate (CaP) coatings deposited on Ti at a working gas pressure of 0.8 mTorr under tilt angles of 0◦ (**a**), 60◦ (**b**), and 80◦ (**c**).

In Figure 4, the dispersion of the size of the surface features for the cases of GLAD at 0◦ and 80◦ is presented. The mean grain size of the coating decreased from 210 to 190 nm with the increasing tilt angle.

**Figure 4.** Dispersion of coating surface grain sizes for samples deposited under tilt angles of 0◦ (**a**) and 80◦ (**b**) at a working gas pressure of 0.8 mTorr.

It is well known that the coating deposited by magnetron sputtering insignificantly changes the surface topography [52]. In most cases, the deposited coating repeats the topography of the initial substrate surface. Even after polishing, the Ti surface still had some marks after machining that could potentially disturb or alter the growth of the surface coating's elements. Therefore, polished Si samples were used alongside the Ti substrates in all the deposition runs, as they do not provide additional ambiguity in the understanding of coating growth on a non-smooth surface topography. Figure 5 shows SEM images of CaP on the Si substrate deposited at 0◦, 60◦, and 80◦ tilt angles and at a pressure of 0.8 mTorr. This was the pressure that promoted the dispersion of the most homogeneous surface features and minimized the particle collision during deposition most effectively. Interestingly, the deposition of the CaP coating on polished Si resulted in the appearance of surface grains in smaller quantities. Apparently, the Ti surface with defects, cavities, and an overall rougher topography induced more nucleation sites for subsequent grain formation compared with Si.

**Figure 5.** CaP coatings deposited on Si at a working gas pressure of 0.8 mTorr under tilt angles of 0◦ (**a**), 60◦ (**b**), or 80◦ (**c**) on the Si substrates.

Even though the initial surface of the Si samples was smooth and without pronounced defects or impurities, grain nucleation still took place in regions that were energetically favorable. Even though the surface of coated Si deposited at the normal incidence was rather smooth, it was possible to find rare globular-like shaped grains on its surface. In Figure 5, some of the grains that were found on the smooth surface of the sample deposited without tilting are shown. For samples deposited at an angle of 60◦, the concentration of surface elements increased together with their size. Moreover, in some cases, grains were comprised of smaller structural elements and had complex topography. We assume that this could be a transition state before the grains coalescence into bigger surface structural elements. With an increase in tilt angle to 80◦, the effect of self-shadowing became more pronounced and resulted in the appearance of surface grains of different sizes. In this case, the surface grains covered the whole surface of the sample and left no place for a region without complex topography. From this, we can conclude that an RF magnetron GLAD method for CaP can have a topographical effect on smooth and rough surfaces of different materials. Álvarez et al. [23] performed mathematical modeling with experimental validation for Ti thin films deposited by RF magnetron sputtering. The proposed model correlated well with the experimental results. It was solely based on ballistic deposition phenomena, indicating that the simulations realistically reproduced the competitive growth mechanism involved in GLAD experiments. However, a modeling of a far more complex structure (hydroxyapatite) still needs to be performed.

Figure 6 shows a cross-section SEM of thin films deposited under tilt angles of 0◦ (a) and 80◦ (b) at 100,000× magnification and films under tilt angles of 0◦ (c) and 80◦ (d) at 200,000× magnification on Si substrates. For the untilted samples, the cross section is presented by an amorphous calcium phosphate layer without any visible structural features. In contrast, coatings deposited at 80◦ inclination show an anisotropic, elongated type of inner structure orientated towards the particle source represented by small crystallites. EDX data obtained from both types of samples showed that the Ca/P ratio equals

1.8–1.9, while the content of Zn does not exceed 1 at.%. It is worth noting that the inner structure corresponds to the surface topography of the samples. In that way, the surface of the coatings deposited on the untilted samples is smooth with a few globular-like features (Figure 5a), and a cross-section of this sample (Figure 6a,c) is represented by smooth featureless layer. In the case of 80◦, tilting small crystallites do not only belong to the surface (Figure 5b) but are also visible throughout the coatings' inner structure (Figure 6b,c). Moreover, predominant growth towards the particle flux is detectable, and the angle of structural inclination β is calculated to be in the range of 60◦–65◦. As it is mentioned in the literature [53], the tilt angle is always smaller than the vapor incidence angle (β < θ). In the paper by Grüner et al. [53], it is stated that, as a first order approximation, a linear relationship between the tilt angle and the angle of incidence β ∼= 0.71θ can be obtained for GLAD of silicon. A similar, linear behavior can be found for obliquely deposited germanium and molybdenum thin films covering a wide range of melting points. However, we could speculate that, in our case, this coefficient is going to be in a range of 0.75–0.8 according to the SEM measurements. Taking into account the fact that the melting point of silicon (1414 ◦C) is relatively close to the melting point of hydroxyapatite (1570 ◦C) [54], this coefficient could indicate the differences in the lattices of the sputtered materials. Most published papers dealing with GLAD focused on materials with relatively simple oxide or monocrystal types of structures with a higher level of symmetry when compared with calcium phosphates. In turn, during the sputtering of CaP, various sputtered species travel to the substrate and are later are involved in the condensation step. Therefore, we assume that this fact is responsible for the difference in the calculated coefficient and overall difference in the thin film structure.

**Figure 6.** SEM cross-section images of obliquely deposited Zn-doped CaP thin films deposited on Si under tilt angles of 0◦ (**a**) and 80◦ (**b**) at 100,000× magnification and films under tilt angles of 0◦ (**c**) and 80◦ (**d**) at 200,000× magnification on Si substrates, demonstrating significant changes in the film morphology. The arrows indicate the angle of incidence θ. The columns are tilted by angle β.

In Figure 7, an AFM visualization of CaP coatings deposited on Ti with and without tilting is shown. According to Sarkar and Pradhan [55], the RMS values of film roughness increase with the increasing oblique angle of deposition. However, in our case, the RMS for tilted samples were smaller than for samples without inclination: 17 nm for the 80◦ sample and 28 nm for the sample without inclination (0◦), respectively. We believe that this is because small surface grains, which homogeneously cover the initial rough Ti surface, are reducing the RMS compared with the sample without any inclination. In the case of the normal incidence of sputtered particles, round-shaped grains are formed on top of the Ti surface, artificially increasing the RMS. AFM measurements allow us to have a 3D image of the surface topography and reveal an out-of-plane growth of surface grains (Figure 7c).

**Figure 7.** CaP coatings deposited on Ti at a working gas pressure of 0.8 mTorr under tilt angles of 0◦ (**a**) and 80◦ (**b**) and three-dimensional (3D) visualization of coating deposited under 80◦ showing the out-of-plane growth of grains (**c**).

In order to exclude the impact of the substrates' surface topography on the coating's roughness, AFM measurements were performed for polished Si samples as well (Figure 8). For the samples deposited at the normal incidence, it was almost impossible to detect distinct surface features. In an untilted sample, some structural elements were found. Surface elements are represented in the form of pyramids, which possibly represent artifacts rather than the actual shape. Nevertheless, the RMS value for the untilted sample was very low (down to 200 pm).

**Figure 8.** *Cont.*

**Figure 8.** CaP coatings deposited on Si at a working gas pressure of 0.8 mTorr under inclination angles of 0◦ (**a**) and 80◦ (**b**) and 3D visualization of coating deposited under 80◦ (**c**).

However, the RMS of the GLAD sample was 5 nm, i.e., was significantly higher than that of the untilted sample. Moreover, these results are in good correlation with GLAD-related research, as RMS values are usually found to increase with the increasing oblique angle of deposition [24]. Such changes in morphology may lead to a significant decrease in the values of the contact angle of the coatings.

#### *3.3. Effect of GLAD on Coating Crystallinity and Scratch Resistance*

It is not rare that CaP coatings deposited by RF magnetron sputtering have a low level of crystallinity [56]. Frequently, CaP-deposited coatings are annealed in order to reveal peaks from crystalline Zn-HA [7]. However, in our case, we decided to perform a GIXRD of the untreated sample, as the temperature treatment will significantly change the internal structure of the CaP. Figure 9 shows representative X-ray powder diffraction patterns of the coatings deposited on Ti substrates at oblique angles of 0◦ and 80◦. The GIXRD measurements indicate a formation of mostly amorphous CaP coating with a broad halo at 2θ = 30◦, which can be attributed to hydroxyapatite with a crystallite size 2(1) nm as determined by the Rietveld refinement. The slightly increased intensity of the amorphous halo of the hydroxyapatite coating deposited under 80◦ (Figure 9b) compared with the hydroxyapatite coating under 0◦ might be caused by the increased crystallinity of the HA coating. The calculated lattice parameters of Ti (*a* = 2.951 and *c* = 4.685 Å) agree well with the ICDD database, confirming that the RF magnetron sputtering does not change the structure of the metallic substrate (as expected). Only a small difference between the Ti substrates was observed in the intensity of diffraction peak (002) at 2θ = 38.4◦, which was caused by the typical rolling texture in hexagonal metals, including Ti [57–59]. It is worth noting, that the increased crystallinity of the samples inclined to 80◦ correlates with the formation of small-sized crystallites on the surface and in the cross-section structure of the GLAD samples (Figure 6b,d).

**Figure 9.** *Cont.*

**Figure 9.** Representative X-ray powder diffractograms (with Rietveld refinement) of CaP coatings deposited on the Ti substrate perpendicular to the flux at 0◦ (**a**) and inclined at 80◦ (**b**). The diffraction peaks of HA and Ti (with Miller indices) are shown.

In Figure 10, the results of CaP coating scratching are shown. Thin film adherence is one of the most important aspects when dealing with physical vapor deposition methods. *L*c (critical load) is a function of coating–substrate adhesion, but also involves other parameters related to the shape of the stylus, the mechanical properties of the substrate and the coating, the thickness, and the internal stress. In Figure 10a, a micro-scratch on the surface of the Zn-doped CaP coating on the untilted sample is represented. The *L*c was determined during the scratch test to be a relatively small load of 10 N. The first adhesive cracks were observed after *L*c = 5.3 N for the untilted sample. The average value of the friction coefficient during the measurements was calculated to be 0.3.

**Figure 10.** Optical microscopy of the tracks after scratch testing of the CaP coatings deposited on Ti at a working gas pressure of 0.8 mTorr under inclination angles of 0◦ (**a**), 60◦ (**b**), and 80◦ (**c**).