uncorrected proof

0 downloads 0 Views 3MB Size Report
(polyetheretherketone – PEEK) and determined the optimal regimes of the laser powder ... around the implant; b) targeted delivery systems, i.e. scaffolds whose.
Composite Structures xxx (2018) xxx-xxx

Contents lists available at ScienceDirect

Composite Structures

PR OO F

journal homepage: www.elsevier.com

Nano-size ceramic reinforced 3D biopolymer scaffolds: Tribomechanical testing and stem cell activity Igor Shishkovskya⁠ ,⁠ b⁠ ,⁠ c⁠ ,⁠ ⁎⁠ , Vladimir Sherbakova⁠ , Ildar Ibatullinb⁠ , Vladislav Volchkovc⁠ , Larisa Volovac⁠ a b c

Lebedev Physics Institute of Russian Academy of Sciences, Samara Branch, Novo-Sadovaja Str. 221, Samara 443011, Russia Samara State Technical University, Molodogvardeiskaia 244, Samara 443110, Russia Samara State Medical University, Chapaevskaja Str. 89, Samara 443099, Russia

ABSTRACT

UN CO RR EC TE D

ARTICLE INFO Keywords: Ceramic-polymer composites Selective laser sintering/melting 3D matrix and scaffolds PEEK Mesenchymal stromal stem cells (MSSC) Tribological and mechanical properties

The present study was aimed to study the nanosize ceramic reinforced 3D biopolymer scaffolds on the stem cells viability, proliferative activity, and phenotypic changes during the testing. We used the biocompatible polymer (polyetheretherketone – PEEK) and determined the optimal regimes of the laser powder bed fusion process. The powders of biocompatible nano-size reinforced oxide ceramics – TiO2⁠ , Al2⁠ O3⁠ , ZrO2⁠ , and hydroxyapatite were carried out as a basis of the additives. The influence of additional heating during PBF process on the mechanical and tribological properties and stem cell behavior near ceramoplastic scaffolds reinforced by nano oxide’s inclusions were carried out.

1. Introduction

Tissue engineering is an interdisciplinary scientific trend providing a solution to the urgent problem of replacing lost tissues that applies the principles of the traditional materials science, construction mechanics, biochemistry, anatomy and pathology for developing biological substitutes able to restore, preserve and improve the functions of damaged tissues [1–4]. In natural conditions, the place where the cells are located is an extracellular matrix which is a complex of both organic and inorganic compounds that fill the intercellular space. One of the major directions of tissue engineering is the development and manufacturing of three – dimensional (3D) implants, i.e. scaffolds. Modern scaffolds are analogues of the extracellular matrix and can be used both as a substrate for their filling with the body's own cells and for the making of tissue-engineered constructs containing the cells cultured in vitro [5–7]. The tissue-engineered constructs can be roughly divided into the following three groups [8–10]: a) scaffolds that are “empty” matrices for adhesion and proliferation of cells migrating from the tissue around the implant; b) targeted delivery systems, i.e. scaffolds whose bioactivity is artificially increased due to changes in the pore architectonics and by including into their composition of various growth factors that are being gradually released



from the matrix; c) scaffolds with pre-grown cells cultures grown in vitro inside them. Additive technologies allow the formation of three-dimensional (3D0 objects of almost any shape using their digital model. A large number of methods are available for the implementation of the additive manufacturing of 3D implants – scaffolds, but within the scope of this work, we will only report on the technology of powder bed fusion (PBF) [1–3]. Selective laser sintering/melting (SLS/M) as one of the good examples of PBF-technology allows to obtain 3D functional products with a precisely specified structure by their digital model. To obtain scaffolds by this method, a thermoplastic powder material is required, capable of melting under the influence of infrared laser radiation of CO2⁠ laser with the wavelength of 10.6 µm. This material is placed on the platform, and then the laser beam forms the bottom layer of the specified object. After that, the platform is lowered by the height of one layer and the cycle repeats. The spatial resolution of the object and the accuracy of detailing depend on the laser beam diameter and the powder fraction parameters. Polyamide, polycarbonate, polypropylene, polyetheretherketone (PEEK) are the main thermoplastic materials for the SLS/M. Due to the small number of biocompatible polymers suitable for this method of scaffolds fabrication, its modifications are being searched and developed [11]. An indirect SLS/M is a new approach to the formation of objects with no direct thermal effect on the polymer, thus allowing the use of more

Corresponding author at: Lebedev Physics Institute of Russian Academy of Sciences, Samara Branch, Novo-Sadovaja Str. 221, Samara 443011, Russia. Email address: [email protected] (I. Shishkovsky)

https://doi.org/10.1016/j.compstruct.2018.03.062 Received 15 January 2018; Received in revised form 14 March 2018; Accepted 16 March 2018 Available online xxx 0263-8223/ © 2018.

I. Shishkovsky et al.

Composite Structures xxx (2018) xxx-xxx

ment, the structure of the scaffold must simulate the bone by its mechanical, biological and functional characteristics. Hydroxyapatite (HA), previously widely used as a coating for metal implants, proved to be very suitable for the stimulation of reparative processes of the bone tissue [30,31]. This was due to its high ability to facilitate the bone growth along the surface of the biologically inert implant, therewith improving the integration of the implant into the bone and preventing the secondary loosening or slacking, and hence – the need for repeated operations [32,33]. However, the SLM – manufacturing of scaffold only on the basis of oxide ceramics, e.g. from aluminum and zirconium [34], is difficult due to the incapacity to simultaneously achieve both high porosity and high strength, that is why the use of polymer as a matrix filler is absolutely justified [3,6,13]. The polymer scaffolds with the HA included inside them have been studied for a long time [35–39], and their ability to stimulate the osteoblasts differentiation both in vitro and in vivo was confirmed in the course of studies [38–42]. In recent years, the use of the combined scaffolds consisting of alternating layers of biopolymer and oxide ceramics in different proportions has been reported [43,44]. It should be noted that the use of the combination of collagen and hydroxyapatite allows the formation of the bone and cartilaginous tissue regenerate within one implant. Note that the bone formation occurs in deep layers of the scaffold, while the cartilage regeneration takes place on the most superficial part [30,45,46]. The possibility to regenerate the complex of bone and cartilaginous tissue with the help of combined scaffolds determines a broad range for their application in comparison with the single-component matrices [47–49]. The effect of admixtures of oxide ceramics nanoparticles in the biopolymer matrix after the additive manufacturing is also being thoroughly studied [50–56]. Thus, by combining different materials of both natural and artificial origin in the scaffold composition, it is feasible to achieve its optimal characteristics in order to control the bioresorption rate of the composite, as well as its bioactivity and mechanical strength. Recent composite materials have improved and raised to a new level such parameters of modern tissue-engineering structures as biocompatibility, mechanical and strength parameters, bioresorption rate. The use of composite materials for the scaffolds formation allows to obtain unique combinations of properties: light weight, high mechanical strength, high porosity with a small pore size [57]. In our earlier studies [13,26] we showed how the PBF approach of 3D biosynthesis was applied to fabricate a porous polyetheretherketone (PEEK) shell structures including encapsulated micron & nano-sized oxide Al2⁠ O3⁠ , ZrO2⁠ , TiO2⁠ and hydroxyapatite (HA) particles, distributed heterogeneously in the matrix of the PB-Fused polymer. Under the present study, we continue to further develop our approach to the biocompatible polymer PEEK as the polymer matrix, and the nano-sized powders of biocompatible oxide ceramics – TiO2⁠ , Al2⁠ O3⁠ , ZrO2⁠ and/or HA were evaluated as the additives. The studies of the influence of an additional heating during the PBF process on the mechanical and tribological properties and the stem cell behavior near ceramoplastic scaffolds reinforced by nano oxides inclusions were carried out.

UN CO RR EC TE D

PR OO F

materials, including additives [12,13]. The principal difference of the approach is that a small amount of the material (e.g. carbon or metal), capable of intensive absorption of laser radiation at another wavelength in the near infrared region, is added to the polymer. The direct effect of the laser is only observed on the particles of the additive, it consists in the heat transfer to the polymer matrix, while the polymer itself and its useful properties remain unchanged. Several types of metal-polymer compositions for these purposes have been developed and patented by the authors [12]. Currently, the synthetic materials with a whole number of positive characteristics, in particular, such as the simplicity of manufacturing and chemical modification, high versatility, ability to biodegradation and possibility of modulation of their mechanical properties (friction, wear resistance, biological stability) are of great interest for the tissue engineering. For example, PEEK is proposed as the material for making a hip joint implant for elderly people suffering from osteoporosis, so that to reduce the mechanical load at the place of the implant fixation on fragile bones and at the same time to save the functional qualities of the product [14–16]. In addition, an increased percentage of the filler or a decrease in the filler size (dispersion) results in the surface area increase, thus leading to the raise of the scaffold bioresorption rate [17–19]. From the mechanical standpoint, synthetic composite materials are anisotropic structures, since their properties vary both in the direction of growth in the chamber of the additive facility, and also depend on the direction of the operating load. As a rule, the inclusion of fillers into the polymer scaffold can increase its mechanical strength [20,21]. Nevertheless, this can only be considered reasonable if the filler does not exceed the threshold load, that depends on the ratio of the composite and the substance included in it [13,18,23,24]. The filler size can also have an effect on the composite mechanical properties. As shown in the researches of S.K. Misra [22] and J.H. Jo [25], the addition of nanostructured ceramics to the polymer scaffold provides a higher mechanical strength as compared to the polymer matrix with the built-in structure of micron-size bioactive ceramics [26]. Thus, by using the synthetic composite matrices it is possible to try to model the mechanical parameters of composite materials with due regard to the biomechanical particular qualities of the sphere of intended use of the scaffold. And this is a clear advantage over the natural polymers, especially for the stimulation of reparative processes in the skeleton connective tissues. A distinctive feature of scaffolds used to stimulate regeneration of the bone tissue, consists in the fact that they are not just a temporary matrix for the stimulation of differentiation and cells growth, they also provide the mechanical strength of the operated segment under load, thus stimulating eventually the bone tissue remodeling [27,28]. On taking into consideration all the features enumerated above, it can be concluded that the scaffold intended to replace defects and stimulate the regeneration of the bone tissue should have some special mechanical properties allowing to endure an early load immediately after the surgery. At the same time, along with the increased strength, one more obligatory requirement for such matrices is their high porosity (>80%), necessary for the integration of the body's own cells into the scaffold structure. Besides, this high porosity is also necessary to ensure an adequate delivery of nutrients and trophies to the regenerated implant [4,29]. It should also be mentioned that the previously developed scaffolds were used to guarantee the mechanical strength of the bone tissue, i.e. a special attention was paid to osteoconductive properties. Combined scaffolds currently being developed on the base of polymers possess both osteoconductive and osteoinductive properties, provide the possibility of a mechanical load and owing to the internal porous microarchitecture stimulate not only the processes of growth but also osteogenic differentiation of cells [27,28]. It can be concluded from the above mentioned that in order to produce an optimal implant for the bone tissue replace

2. Materials and experimental procedure 2.1. Materials

In this study, a commercially available polyetheretherketone (PEEK – Victrex Co., UK) was used. For better sinterability, the size of the particles taken for the polymer fractions, ∼50 µm, was comparable with the laser beam diameter [5]. As the filling material, oxide ceramics powders (TiO2⁠ ∼ 50, Al2⁠ O3⁠ ∼ 15, ZrO2⁠ ∼ 20, HA ∼ 60 nm, EvNanoTech.com. Ltd., China) were used. In earlier studies [6,12,13,26], the ceramoplast powder mixture compositions were optimized and in the present study, we prepared the mixture in the 1:10 ratio by wt. (the

2

I. Shishkovsky et al.

Composite Structures xxx (2018) xxx-xxx

the polymer sample being examined. During the tests, a load of 90 N contact was provided, and a normal pressure of about 6 MPa, typical for human joints, was created in the friction zone. The rotation speed ⁠ 1 of the control sample was 600 min− , taking into account the fact that the sliding speed in the joints, as a rule, does not exceed 0.2 m/s. As a model lubricant for the comparative analysis of tribotechnical properties, a graphite lubricant was used. In the lubricant composition, the proportion of a coarse graphite as an antifriction additive was 10%. The 5 3D samples were printed for each of the five PEEK+ nano ceramic powder compositions being studied. The surface of the samples was ground to a roughness Ra < 1 μm (without the use of coolant). The specimen was fixed in a mandrel and lubrication (1 mg) was applied to the working surface. We ran a working program in the computer and lowered the counter to the sample surface, including the stand. Within the test duration of 10 min, an automatic data collection (i.e. normal load, friction torque and self-heating temperature of the friction unit under test) was carried out from the sensors. As soon as the test period expired, the setup was turned off, data collection was stopped and the test results were stored in the PC database. After the removal of the tested sample from the cup, its linear wear was evaluated with the help of the profilograph-profilometer Mikron-01. Calculations of the friction coefficient, f, and frictional force were made, taking into account the geometry of the counter-sample and the normal load (90 N) according to the above-mentioned equations.

PR OO F

first figure in the ratios is for oxide ceramic). Our experimental setup for the laser PBF process via the SLS/M technique was earlier described in Refs. [13,26] as a procedure of the laser parameter optimization. We applied an additional heating of the 3D samples for the two reasons: first, heating during the SLS up to the glass transition temperatures of the PEEK (∼150 °C) changes the conditions of the layer-by-layer manufacturing; and second – a slight heating of the implant during the operation (∼40 °C) is also possible. The appearance of the test specimens is shown in Fig. 1. 2.2. Microstructure characterization After the etching, the cross sections of multi-layered sintered samples were subjected to a morphological microanalysis with the optical microscope (Neophot 30 M, Carl Zeiss) equipped with a digital camera. The 3D samples obtained under the optimized regimes were analyzed by a microhardness tester PMT-3M (OKB SPECTR Ltd., St. Petersburg, Russia). The microstructures were studied by a scan electron microscopy LEO 1450 (Carl Zeiss Company) equipped with an energy-dispersive X-ray microanalyzer (INCA Energy 300, Oxford Instruments). 2.3. Tribomechanical testing

UN CO RR EC TE D

The universal complex Universal-1B (Samara-Balance Ltd., RF) was used for the tribotechnical tests, in order to collect and display in the process of friction the data on the normal contact load, self-heating temperature of the friction unit and frictional moment Mt⁠ r (N·m). After the tests, a visual and metallographic analysis of the samples was made, a calculation of the friction coefficient – f was carried out taking into account the reference sample geometry and the normal load (90 N) by the formula f ≈ 4.44Mt⁠ r. The formula Ft⁠ r ≈ 400Mt⁠ r (N) was used for estimation of the frictional force. For the sample microgeometry assessment after the SLS/M process the profilograph-profilometer Mikron-01 (Samara-Balance Ltd., RF) was used. This profilograph allows determination of the roughness within the range of ±300 μm, as well as working with materials of different classes – metals, ceramics, polymers. In view of the large roughness of the 3D printed surface (see Fig. 1) the base was chosen to be of ∼1 mm long. Mechanical tests were carried out with a tensile machine Samson-01 (Samara-Balance Ltd., RF), which allows testing of the miniature samples of 3D printed parts (Fig. 1) with the cross-sectional area of 6 mm2⁠ and base length of 3 mm. The machine allows the samples to be heat-stabilized before testing. Tests with heating are necessary to detect the ultimate changes in strength and deformation properties under the forced operation regimes of the scaffolds. Taking into consideration a probable but undesirable self-heating of the friction unit in the bonding of the implant with the bone that can lead to the temperature increase at the contact point, we conducted the breaking tests with the sample heating up to 40 °C. For the tribotechnical tests, a ring-plane sliding friction scheme was used, which allowed the simulation of the work of heavy-loaded joints operating under the conditions of boundary sliding friction. As the material for the control sample, steel 40X (in the tempering state HRC45) was selected. At the same time wear is provided during friction of only

2.4. Stem cell culture preparation and treatment The scheme of the experiment included three parts of the works carried out [6,30]. The first one was the preparation of the materials for the research, the second part consisted in the toxicity testing in separated cameras, and the third one – cultivating of the multipotent mesenchymal stem cells (MMSC) on the substrate. Pure growth mediums with stem cells, both without samples and with biocompatible PEEK sample, were used as a reference group. All the samples from polymer + nanoparticles, inserted into the growth medium with the stem cells, were considered as testing groups. All the 3D bioprinted samples were first washed in a sterile phosphate salt buffer (PBS) by reiterated immersion and active stirring of the samples in the solution. After that, the samples were sterilized in a Sterident steam sterilizer at the temperature of 121 °C and pressure of 120 kPa for 20 min. The third passage of MMSC was chosen. The cells were taken from the umbilical cord of an adult donor, who previously subscribed the informed consent to donate the biomedical material in the agreement with the international legislation. The isolation of the cells was accomplished on the Ficcol consistence gradient 1.033. The cells inoculation on the material was carried out in a new 6-cavities plotter. The materials under research were placed into the clean cavities, then the MMSC with the concentration of 100 thousand cells per 1 cm2⁠ and maximal volume of 500 mcl of the growth medium were applied right upon them. Then the material was incubated for 3 h to ensure the adhesion of the cells to the material (the preliminary incubation). The cultivation lasted for 7 days using the standard nutrient mediums including aMEM (Sigma), 2 mM alanil-glutamine (Invitrogen Co.) and 10% selected calf serum (Gibco Co.) by the design scheme. We changed

Fig. 1. Appearance of samples after the SLS process (‘n’ – is nano): (a) pure PEEK; (b) PEEK + nAl2⁠ O3⁠ ; (c) PEEK + nZrO2⁠ ; (d) PEEK + nTiO2⁠ ; (e) PEEK + nHA. 3

I. Shishkovsky et al.

Composite Structures xxx (2018) xxx-xxx

elongation. According to Ref. [58], the strength limit of pure and reinforced casting polymers is in the range of 40…100 MPa and the elongation achieves 25%. It is noted that the flow area is not expressed, and residual deformations arise only in the area of the critical on-load action. It is known from biomechanical studies, that cartilages in human joints experience pressures from 0,2 to 6 MPa. The most loaded hip joints under a fast walking experience pressures up to 18…21 MPa, that corresponds to the conditions of operation of heavily loaded friction units. In this case, the friction energy predominantly passes into heat, thus causing a change in the strength properties of the polymers. Therefore, mechanical tests were carried out at two temperatures – ambient temperature and elevated one, so that to reveal the ultimate changes in strength and deformation properties under the forced operating modes of the polymer products. Since burns and tissue necrosis occur at temperatures above 55 °C, the maximum allowable temperature of 50 °C was used in our tests with interruptions between heatings. The stress-strain diagrams obtained as a result of tensile tests for the five types of sintered PEEK-based scaffolds with nano-additives are shown in Fig. 3. The results of the undertaken evaluation of the samples mechanical properties are summarized in Table 1. On the whole, it can be noted that the strength properties of the studied ceramic-polymer compositions obtained by the 3D-printing method, they are inferior to the casting PEEK by almost 50%. Nevertheless, our samples are suitable for use as implants for the field of heavy loads, including a prosthesis for the hip joints. The type of the obtained diagrams indicates that the yield plateau at the room temperature is not clearly visible, but with heating, it becomes more distinct in many cases. Unevenness in the plots (multiple peaks) arises as a result of several nucleus fractures occurring due to the increased defect distribution. The best results (σs⁠ = 33 ± 4.5 MPa) were shown for the samples sintered from pure PEEK. As a whole, the PEEK with nano-additives showed a decrease in strength and deformation properties, progressing with the increasing concentration of the nano-inclusions as it was in Ref. [13]. The tests with heating showed that the temperature increase leads to a significant decrease in the strength limit and an increase in the PEEK plasticity, that seems quite natural. The obtained results revealed the problem of introduction of nano-inclusions in the 3D-printed samples based on the PEEK. The reducing of the mechanical properties of polymers during sintering, uncharacteristic of nanomodifiers, can be explained by the fact that nanopowders tend to coagulate during storage into aggregates of submicron and micron sizes due to a high surface energy. In the conglomerated form, they lost their high activity normally characteristic of nanoparticles. Moreover, the fragile formation was observed in 3D parts, and in our opinion, the reason was that such conglomerated additives created centers of a high-stress concentration when external loads were applied, especially

PR OO F

the nutrient medium every 4 days or every time we noticed that the medium pH alteration had changed. For each group, a daily morphometry was carried out. The cell-covered area was estimated using the Axio ObserverA1 microscope and software package produced by Carl Zeiss in Axio Vision Inc. The culture cluster analysis was conducted on the ImageJ and Image ProPlusbundled software, and the cells were divided into groups in accordance with their pixel density per object. The immunophenotyping procedure was accomplished with a running FACS Canto (Becton Dickinson Co.) cytofluorometer right before the inoculation and after removal of the stem cells on the 0th and 7th day of the cultivation. The following antigens were studied: CD73+, CD 90+, CD 105+, CD 34−, CD 45. 3. Results and discussion 3.1. Preheating influence

UN CO RR EC TE D

We carried out a study on the effect of a pre- and post-thermal treatment on the results of the layered SLS in systems (PEEK + nanoceramics). The preheating was made directly in the 3D-printing chamber (during the SLS process) while post-processing (i.e. thermal annealing) of 3D samples in a muffle furnace for finished parts. The pre- and post-treatment temperatures were chosen below the melting point, within the glass transition temperature range of the corresponding polymer. It is known that PEEK has a glass transition temperature Tg⁠ = 143 °C [58], it operates up to the temperature of 260 °C, its melting point Tm ⁠ = 343 °C, and PEEK destruction terminates at the temperatures of 575–580 °C [13,30]. That is why we chose the temperature of ∼150 °C in the synthesis chamber. As it is seen, there are no significant changes in the 3D samples. As it turned out, the density measurements (Fig. 2) showed a slight increase of the 3D parts density under heating. The thermal heating (a subsequent annealing) of the samples after the SLM at 150 °C for 1 h practically did not change the initial density of the samples. Moreover, we noted a slight 3D part shrinkage for a series of samples based on the pure PEEK. The X-ray structural analysis, optical metallography and SEM with EDX, earlier conducted by us [13], confirmed that after the SLS/m of such mixtures, no significant modifications were observed. It means that these materials can be recommended for medical applications. 3.2. Mechanical testing

The mechanical properties of the PEEK polymer in the solid state are characterized by two main parameters that determine the structural strength of products: the ultimate tensile strength and maximum ultimate

Fig. 2. Change of 3D part density due to additional heating in the PEEK + nano-oxide systems. 4

Composite Structures xxx (2018) xxx-xxx

PR OO F

I. Shishkovsky et al.

Fig. 3. The stress-strain diagrams of the 3D printed samples from PEEK with nano additives: (a, b) pure PEEK; (c, d) PEEK + nAl2⁠ O3⁠ ; (e, f) PEEK + nZrO2⁠ ; (g, h) PEEK + nTiO2⁠ ; (i, j) PEEK + nHA. The upper row (a, c, e, g, i), the measurements were at room temperature, the lower row (b, d, f, h, j) at heating to −50 °C.

stable coefficient of friction. The addition of hydroxyapatite resulted in 3 times reduction of the surface roughness as compared to the unmodified polymer. Additives of titanium dioxide and zirconium dioxide did not show significant changes in the tribological properties of the polymer. During the friction testing, the sample heating did not exceed 1…2C. Profilograms of the surfaces of five samples of the sintered ceramic plastic based on the PEEK matrix are shown in Fig. 5. The corresponding microgeometry characteristics are given in Table 3. It can be noted that after the PBF use via the SLS/M method, a rather high surface roughness of Rm ⁠ ax, often exceeding 100 μm, is formed. The rough surface facilitates the scaffold osteointegration into the bone tissue and this is a favorable factor. A characteristic wavy surface profile (wave step Sb⁠ = 300…500 μm) can be seen in the profilograms, also reflecting the conditions of the laser sintering of the powder. The use of 3D-printed samples with high values of the parameter Rv⁠ k, characterizing the surface oil consumption, provides good conditions for retaining the lubricant in the cracking zone. A visual and microstructural analysis of the SLS/M samples was carried out after the friction tests and tensile fracture. In Fig. 5 the sample appearance is shown after the rupture (a–d) and friction tests (e–f). The rupture fracture was studied by OM and some images are shown in Fig. 6 both at room temperature and at an elevated temperature of 50 °C. The analysis of macro- and microphotographs (Figs. 6 and 7) proves that the destruction process proceeds mainly by a brittle type, without significant stretching of the polymer filaments. In fact, such stretching is characteristic of polymeric materials and nevertheless, we observed it also on polymers with nano-inclusions, such as Al2⁠ O3⁠ or TiO2⁠ . The value of stretching can be estimated by the decrease in the cross-section of the broken samples (see Fig. 7a, b, d). Virtually all the samples were destroyed in the middle, and only a few of them – in the place of fixation (Fig. 6c). Basically, in the destroyed samples with the PEEK matrix, no delamination was observed (this is generally characteristic of the layered SLS/M process). It testifies to a good adhesion of the layers in our additive 3D-printing process. And at last, the 3D sample delamination was not observed for the loads applied at the point of friction (Fig. 6e, f). In these figures, the imprint of the counter-body is clearly visible.

Table 1 The results of the ultimate strength σs⁠ (MPa) and the specific elongation – ε (%) of the PEEK matrix with nano additives after 3D printing. Mechanical properties

UN CO RR EC TE D

Sample material

Under T = 20 °C

PEEK PEEK + Al2⁠ O3⁠ PEEK + HA PEEK + TiO2⁠ PEEK + ZrO2⁠

Under T = 50 °C

σs⁠ , MPa

ε, %

σs⁠ , MPa

ε, %

37,24 25,3 23,3 15,3 18,1

15,5 8,0 2,7 5,0 5,9

28,45 22,0 5,5 8,5 17,5

14,0 5,8 5,2 7,8 9,8

in segregation zones. Therefore, a careful mixing of nano-additives with PEEK powder should be recommended to prevent these problems. An example of the experimental measurement of a normal load diagram, the average temperature in the friction zone, and frictional moment in the wear test in the PEEK + nAl2⁠ O3⁠ sample is shown in Fig. 4. For the remaining 3D samples, the behavior of these parameters was similar. All the antifriction and antiwear characteristics of the samples obtained during the tests were summarized in Table 2. As it follows from the diagrams (Fig. 4 and Table 2), the frictional moment is characteristic of periodic changes, which indicates that a fatigue mechanism for the PEEK wear is flowing by the thermofluctuation mechanism. Each frictional moment jump corresponds to the kinetic cycle “damage accumulation - material destruction”. If the friction conditions are maintained, the cycle period for each ceramic-polymer mixture remains approximately the same, something about 25 s. Despite these changes in the friction moment, on average, the friction level remained near the same value, indicating that the studied friction pairs had frictional compatibility under given friction conditions. For all the friction pairs tested, the coefficient of wear resistance was about eight, which confirms a sufficiently high resistance of materials to the fatigue wear. The tested samples showed high antifriction properties (friction coefficient ranged from 0.03 to 0.22), close to the region of the lower boundary of the coefficient friction range for the thermoplastic polymers in steel of 0.15–0.40, and also lower by 1.5…2 times than that for the casted PEEK, having the coefficient of friction for steel equal to 0.30…0.38 [58]. The Al2⁠ O3⁠ nanomodifier provided an almost double increase in the antifriction properties of the friction pair, as evidenced by a lower and

3.3. Results of the stem cell culture treatment The proliferative activity and viability of the cells are shown in Fig. 8a, b. In comparison with the reference group samples (MMSC and PEEK only), these parameters of crucial importance meet our expectations and even exceed them in some samples.

5

Composite Structures xxx (2018) xxx-xxx

UN CO RR EC TE D

PR OO F

I. Shishkovsky et al.

Fig. 4. Tribotechnical diagram of PEEK + nAl2⁠ O3⁠ sintered sample.

Table 2 The results of tribological tests for the PEEK matrix with nano additives after 3D printing. Sample material

PEEK PEEK + Al2⁠ O3⁠ PEEK + HA PEEK + TiO2⁠ PEEK + ZrO2⁠ *⁠ h

Friction and wear properties of specimens h, μm

γ, μm/h

J

Mf⁠ r, min/max N·m

fm ⁠ in/max

5 10 3 10 6

30 60 24 60 36

⁠ 8 5,26 · 10− ⁠ 7 1,06 · 10− ⁠ 8 3,15 · 10− ⁠ 7 1,05 · 10− ⁠ 8 6,3 · 10−

0,008/0,051 0,004/0,045 0,01/0,05 0,018/0,056 0,02/0,058

0,035/0,226 0,018/0,198 0,044/0,220 0,079/0,246 0,088/0,255

– linear wear; γ – wear rate; J – wear intensity; Mf⁠ r – frictional torque; f – frictional coefficient.

The OM and structure morphology results for the 3rd and 6th days were presented in Table 4. So we can see that the culture density near the PEEK + TiO2⁠ was 10% on the 3rd day and 15% and 30% only for PEEK + ZrO2⁠ and PEEK + Al2⁠ O3⁠ . The cells had a fibroblast-like morphology and were slowly dividing. Near the pure PEEK, it was 40%, the cells had the fibroblast-like morphology, and their divisions and even apoptosis were observed. The similar behavior was observed earlier in Ref. [6]. So, by the 6th day, we can state the following:

After the medical test, an estimation of the MMSC on the materials was carried out using the SEM with great magnification. All materials had no evident signs of the cell fixation. Such a result can arise from the material properties (e.g., great adhesive ability), but can also be caused by the harshness of the sample preparation procedure (for instance, the cells could have been washed away). L. Gutierrez et al. [59] denoted that a lot of factors could significantly influence the features of the synthesized objects and prospective medical properties, such as the core/ shell types, sizes, values between them, and their redistribution over the matrices.

• The level of confluence in the reference groups ranged from 70 to 90% which stands for the presence of a monolayer; • The same growth was observed in cavities 3, 4, 5 (see Table 4). The culture density here also went up to 80–90% and achieved the level of a monolayer; • In all the other cavities the density was observed below of 40–60%; • The cell morphology is uniform, fibroblast-like without stress-fibers in all the cavities.

4. Conclusions

Thus, in the present work, the MMSC morphology and intercellular contacts were studied, including the case of the samples additional heating during their preparation. The results on the proliferative activity and dynamical differentiation of the MMSC on the SLS-obtained porous biocompatible PEEK matrices with the ceramic fillers were obtained. It

6

Composite Structures xxx (2018) xxx-xxx

UN CO RR EC TE D

PR OO F

I. Shishkovsky et al.

Fig. 5. Profilogram from the surface of 3D printed samples: (a) pure PEEK; (b) PEEK + nAl2⁠ O3⁠ ; (c) PEEK + nZrO2⁠ ; (d) PEEK + nTiO2⁠ ; (e) PEEK + nHA.

can be concluded that the additional heating results in a decrease in the 3D-samples porosity. The denser samples with a less developed porous surface do not contribute to the active proliferation of stem cells, as has been observed earlier. It was shown that nano ceramoplast composites generally retained their shape after the heating, but there was a slight shrinkage of 3D parts for a series of samples based on the pure PEEK matrix. By analyzing the results of our research it can be concluded that:

• heating of the 3D samples during the laser PBF process, ensuring a lower porosity of sample, results in nearly by-half decrease of the proliferation speed; • mixtures of PEEK + nano (Al2⁠ O3⁠ /ZrO2⁠ /or TiO2⁠ ) have a worse proliferative activity in comparison with the pure PEEK or PEEK + nHA.

7

I. Shishkovsky et al.

Composite Structures xxx (2018) xxx-xxx

Table 3 Results of the surface microgeometry evaluation for 3D printed samples.

PEEK PEEK + Al2⁠ O3⁠ PEEK + H PEEK + TiO2⁠ PEEK + ZrO2⁠

Surface microgeometry characteristics Rm ⁠ ax, μm

Rp⁠ , μm

Rz⁠ , μm

Ra⁠ , μm

Sm ⁠ , μm

Rp⁠ k, μm

Rv⁠ k, μm

tm ⁠

99,8 110,0 36,0 134 99,8

43,1 71,5 20,7 86,1 43,1

96,3 107,0 35,8 111,0 96,3

22,6 16,8 5,85 22,9 22,6

518 302 164 482 518

33,9 66,4 18,3 60,2 33,9

36,2 22,2 8,54 42,1 36,2

0,44 0,42 0,47 0,44 0,44

PR OO F

Sample material

The conducted mechanical and tribological testing allowed to obtain new data on strength, linear wear, wear profilometry, fracture rupture for 3D samples in systems (PEEK + nano TiO2⁠ , Al2⁠ O3⁠ , ZrO2⁠ , or hydroxyapatite ceramics) after the laser PBF process which could be useful for future medical application of such materials. Acknowledgement

UN CO RR EC TE D

Authors thanks the Russian Science Foundation, grant No. 15-19-00208. Fig. 6. Sample appearance after the mechanical tests.

Fig. 7. Sample break fractography: upper raw – 20 °C; bottom raw – 50 °C.

Fig. 8. Stem cell viability (%) on the 6th day (a); stem cell culture density on the 6th day (b).

8

I. Shishkovsky et al.

Composite Structures xxx (2018) xxx-xxx

Table 4 Comparative characterization of stem cell culture 3rd/6th days (through slash). PEEK + ZrO2⁠

PEEK + Al2⁠ O3⁠

PEEK + HA

PEEK

Control

1

2

3

4

5

8

10/40% + −+

15/60% + −+

30/70% + −+

40/80% + −+

40/90% + +

50/90% + +

References

PR OO F

Density Morphology Division

PEEK + TiO2⁠

[21] Y. Zhou, et al., In vitro bone engineering based on polycaprolactone and polycaprolactone-tricalcium phosphate composites, Polym Int 56 (3) (2007) 333–342. [22] S.K. Misra, et al., Comparison of nanoscale and microscale bioactive glass on the properties of P(3HB), Bioglass Compos Biomater 29 (12) (2008) 1750–1761. [23] G. Vozzi, C. Corallo, C. Daraio, Pressure-activated microsyringe composite scaffold of poly (l-lactic acid) and carbon nanotubes for bone tissue engineering, J Appl Polym Sci 129 (2) (2013) 528–536. [24] M. Mattioli-Belmonte, et al., Tuning polycaprolactone–carbon nanotube composites for bone tissue engineering scaffolds, Mater Sci Eng C 32 (2) (2012) 152–159. [25] J.-H. Jo, et al., In vitro/in vivo biocompatibility and mechanical properties of bioactive glass nanofiber and poly (ε-caprolactone) composite materials, J Biomed Mater Res B Appl Biomater 91-B (1) (2009) 213–220. [26] I. Shishkovsky, V. Scherbakov, Selective laser sintering of biopolymers with micro and nano ceramic additives for medicine, Phys Proc 39 (2012) 491–499, https:// doi.org/10.1016/j.phpro.2012.10.065. [27] G.O. Cunniffe, F. Brien, Collagen scaffolds for orthopedic regenerative medicine, JOM 63 (64) (2011) 66–73. [28] K.V. Suresh, Clinical use of different biomaterials for cutaneous wound healing in veterinary practice, World Cong Biotechnol (2011) https://doi.org/10.4172/ 2155-9538.10000S1. [29] M.M. Stevens, Biomaterials for bone tissue engineering, Mater Today 11 (5) (2008) 18–25. [30] I.V. Shishkovskii, Yu. G. Morozov, S.V. Fokeev, L.T. Volova, Laser synthesis and comparative testing of a three-dimensional porous matrix of titanium and titanium nickelide as a repository for stem cells, Powder Metall Met Ceram 50 (9/10) (2012) 606–618, https://doi.org/10.1007/s11106-012-9366-9. [31] I.V. Shishkovsky, E.Yu. Tarasova, L.V. Zhuravel’, A.L. Petrov, The synthesis of a biocomposite based on nickel titanium and hydroxyapatite under selective laser sintering conditions, Tech Phys Lett 27 (3) (2001) 211–213, https://doi.org/10. 1134/1.1359830. [32] R.G. Geesink, N.H. Hoefnagels, Six-year results of hydroxyapatite-coated total hip replacement, J Bone Joint Surg Br 77 (4) (1995) 534–547. [33] R.Z. Le Geros, Properties of osteoconductive biomaterials: calcium phosphates, Clin Orthopaed Related Res 395 (2002) 81–98. [34] I. Yadroitsev, I.V. Shishkovsky, P. Bertrand, I. Smurov, Alumina-zirconium ceramics synthesis by selective laser sintering/melting, Appl Surf Sci 254 (4) (2007) 966–970, https://doi.org/10.1016/j.apsusc.2007.09.001. [35] A. Takeuchi, et al., Deposition of bone-like apatite on silk fiber in a solution that mimics extracellular fluid, J Biomed Mater Res A 65 (2) (2003) 283–289. [36] H. Chim, et al., A comparative analysis of scaffold material modifications for load-bearing applications in bone tissue engineering, Int J Oral Maxillofacial Surg 35 (10) (2006) 928–934. [37] M.T. Arafat, et al., Biomimetic composite coating on rapid prototyped scaffolds for bone tissue engineering, Acta Biomater 7 (2) (2011) 809–820. [38] C. Vaquette, et al., Effect of culture conditions and calcium phosphate coating on ectopic bone formation, Biomaterials 34 (22) (2013) 5538–5551. [39] I.V. Shishkovsky, S.E. Volchkov, Ceramics-filled 3D porous biopolymer matrices for tissue-engineering on the stem cell culture: benchmark testing, in: Bartolo, et al. (Eds.), High value manufacturing: advanced research in virtual and rapid prototyping, Taylor & Francis Group, 2014, pp. 121–126, USBN 978-1-138-00137-7. [40] B. Mavis, et al., Synthesis, characterization and osteoblastic activity of polycaprolactone nanofibers coated with biomimetic calcium phosphate, Acta Biomater 5 (8) (2009) 3098–3111. [41] E. Seyedjafari, et al., Nanohydroxyapatite-coated electrospun poly (l-lactide) nanofibers enhance osteogenic differentiation of stem cells and induce ectopic bone formation, Biomacromolecules 11 (11) (2010) 3118–3125. [42] A. Polini, et al., Osteoinduction of human mesenchymal stem cells by bioactive composite scaffolds without supplemental osteogenic growth factors, PLoS One 6 (10) (2011) 26211, https://doi.org/10.1371/journal.pone.0026211. [43] R. Gudas, et al., Ten-year follow-up of a prospective, randomized clinical study of mosaic osteochondral autologous transplantation versus microfracture for the treatment of osteochondral defects in the knee joint of athletes, Am J Sports Med 40 (11) (2012) 2499–2508. [44] E. Kon, et al., How to treat osteochondritis dissecans of the knee: surgical techniques and new trends: AAOS exhibit selection, J Bone Joint Surg 94-A (1) (2012) 1–8.

UN CO RR EC TE D

[1] M. Vaezi, S. Yang, Freeform fabrication of nanobiomaterials using 3D printing, book chapter, in: R. Narayan (Ed.), Rapid prototyping of biomaterials: principles and applications, Woodhead Publishing Limited, Cambridge, UK, 2014, pp. 16–74. [2] J.K. Carrow, et al., Polymers for bioprinting, book chapter, in: A. Atala, J.J. Yoo (Eds.), Essentials of 3D biofabrication and translation, Elsevier Inc., Oxford, UK, 2015, pp. 229–248. [3] I. Volyansky, et al., Laser assisted 3D printing of functional graded structures from polymer covered nanocomposites, book chapter, in: I. Shishkovsky (Ed.), New trends in 3D printing, InTech Publ., Croatia, 2016, pp. 237–258, https://doi.org/ 10.5772/63565. [4] P. Fabbri, et al., Highly porous polycaprolactone-45S5 Bioglass® scaffolds for bone tissue engineering, Compos Sci Technol 70 (13) (2010) 1869–1878. [5] A.A. Dhollander, et al., A pilot study of the use of an osteochondral scaffold plug for cartilage repair in the knee and how to deal with early clinical failures, Arthroscopy 28 (2) (2012) 225–233. [6] I. Shishkovsky, S. Volchkov, Influence of the laser assisted fabricated 3D porous scaffolds from bioceramoplasts of micron and nano sizes on culture of MMSC, Proc SPIE 9065 (2013) 906515, https://doi.org/10.1117/12.2035550. [7] A. Yamashita, et al., Cartilage tissue engineering identifies abnormal human induced pluripotent stem cells, Sci Rep 5 (2013), URL: http://www.nature.com/ srep/2013/130613/srep01978/full/srep01978.htm. [8] Shishkovsky I, Scherbakov V. 4D manufacturing of intermetallic SMA fabricated by SLM process. In SPIE LASE photonics west proceedings, Vol. 10523, Laser 3D Manufacturing V; 1052311; 2018. http://doi.org/10.1117/12.2288176. [9] H. Liu, T.J. Webster, Mechanical properties of dispersed ceramic nanoparticles in polymer composites for orthopedic applications, Int J Nanobiomed 5 (2010) 299–313. [10] A.C. Paz, et al., Tissue engineered trachea using decellularized aorta, J Bioeng Biomed Sci (2011) https://doi.org/10.4172/2155-9538.S2-001, SCI S2:001. [11] A. Mazzoli, Selective laser sintering in biomedical engineering, Med Biol Eng Comput 51 (3) (2013) 245–256, https://doi.org/10.1007/s11517-012-1001-x. [12] I.V. Shishkovsky, I.N. Juravleva, Kinetics of polycarbonate distraction during laser-assisted sintering, Int J Adv Manuf Technol 72 (2014) 193–199. [13] I. Shishkovsky, K. Nagulin, V. Sherbakov, Study of biocompatible nano oxide ceramics, interstitial in polymer matrix during laser-assisted sintering, Int J Adv Manuf Technol 78 (1–4) (2015) 449–455. [14] Carrico JD, Traeden NW, Aureli M and Leang KK. Fused filament additive manufacturing of ionic polymer-metal composite soft active 3D Structures. In ASME 2015 conference on smart materials, adaptive structures and intelligent systems, Colorado Springs, Colorado, USA, September 21–23; 2015. Paper No. SMASIS2015-8895. p. V001T01A004. [15] J. Bauer, S. Hengsbach, I. Tesari, et al., High-strength cellular ceramic composites with 3D microarchitecture, PNAS (2014) https://doi.org/10.1073/pnas. 1315147111. [16] A. Boccaccio, A.E. Uva, M. Fiorentino, et al., Geometry design optimization of functionally graded scaffolds for bone tissue engineering: a mechanobiological approach, PLoS One 11 (1) (2016) e0146935, https://doi.org/10.1371/journal.pone. 0146935. [17] A.M. El-Kady, A.F. Ali, M.M. Farag, Development, characterization, and in vitro bioactivity studies of sol–gel bioactive glass/poly (l-lactide) nanocomposite scaffolds, Mater Sci Eng C 30 (1) (2010) 120–131. [18] S.K. Misra, et al., Effect of nanoparticulate bioactive glass particles on bioactivity and cytocompatibility of poly (3-hydroxybutyrate) composites, J R Soc Interface 7 (44) (2010) 453–465. [19] S.-Z. Fu, et al., In vitro and in vivo degradation behavior of n-HA/ PCL-Pluronic-PCL polyurethane composites, J Biomed Mater Res A 102 (2) (2013) 479–486. [20] X. Shi, et al., Fabrication of porous ultra-short single-walled carbon nanotube nanocomposite scaffolds for bone tissue engineering, Biomaterials 28 (28) (2007) 4078–4090.

9

I. Shishkovsky et al.

Composite Structures xxx (2018) xxx-xxx [53] P.H. Warnke, H. Seitz, F. Warnke, S.T. Becker, S. Sivananthan, E. Sherry, et al., Ceramic scaffolds produced by computer-assisted 3D printing and sintering: characterization and biocompatibility investigations, J Biomed Mater Res B Appl Biomater 93 (2010) 212–217. [54] M. Vaezi, H. Seitz, S. Yang, A review on 3D micro-additive manufacturing technologies, Int J Adv Manuf Technol 67 (2013) 1721–1754, https://doi.org/10. 1007/s00170-012-4605-2. [55] X. Wang, M. Jiang, Z. Zhou, J. Gou, D. Hui, 3D printing of polymer matrix composites: a review and prospective, Compos Part B 110 (2017) 442e458, https:// doi.org/10.1016/j.compositesb.2016.11.034. [56] T. Hanemann, D.V. Szabo, Polymer-nanoparticle composites: from synthesis to modern applications, Materials 3 (2010) 3468–3517, https://doi.org/10.3390/ ma3063468. [57] T. Xu, The advance in self-healing scheme of polymeric materials, J Mater Sci Eng (2012) https://doi.org/10.4172/2169-0022.1000e101. [58] VICTREX™ PEEK Polymers. [accessed on 10 January 2018]. [59] L. Gutiérrez, R. Costo, C. Grüttner, et al., Synthesis methods to prepare single- and multi-core iron oxide nanoparticles for biomedical applications, Dalton Trans 7 (2015) https://doi.org/10.1039/c4dt03013c.

UN CO RR EC TE D

PR OO F

[45] E. Kon, et al., A novel nano-composite multi-layered biomaterial for the treatment of osteochondral lesions: technique note and an early stability pilot clinical trial, Injury 41 (7) (2010) 693–701. [46] G. Filardo, et al., Treatment of knee osteochondritis dissecans with a cell-free biomimetic osteochondral scaffold: clinical and imaging evaluation at 2-year follow-up, Am J Sports Med 41 (8) (2013) 1786–1793. [47] I. Schleicher, et al., Biphasic scaffolds for repair of deep osteochondral defects in a sheep model, J Surg Res 183 (1) (2013) 184–192. [48] Z. Shufang, et al., Bi-layer collagen/microporous electrospun nanofiber scaffold improves the osteochondral regeneration, Acta Biomater 9 (7) (2013) 7236–7247. [49] D. Pingguo, et al., The effects of pore size in bilayered poly(lactide-co-glycolide) scaffolds on restoring osteochondral defects in rabbits, J Biomed Mater Res 102 (1) (2014) 180–192. [50] D. Zhang, B. Gokce, Perspective of laser-prototyping nanoparticle-polymer composites, Appl Surf Sci 392 (2017) 991–1003, https://doi.org/10.1016/j.apsusc. 2016.09.150. [51] B.G. Compton, J.A. Lewis, 3D-printing of lightweight cellular composites, Adv Mater 26 (2014) 5930–5935, https://doi.org/10.1002/adma.201401804. [52] H. Seitz, W. Rieder, S. Irsen, B. Leukers, C. Tille, Three-dimensional printing of porous ceramic scaffolds for bone tissue engineering, J Biomed Mater Res B Appl Biomater 74 (2005) 782–788.

10