Preprint of "The Effect of Pulsed Laser on the Surface State of 3D-Printed Triply Periodic Structures in TiAl6V4 Alloy"

1 The Effect of Pulsed Laser on the Surface State of 3D-Printed Triply Periodic Structures in 1 TiAl6V4 Alloy 2 3 Andrea Školáková1, Jan Pinc1,*, Richard Kubík2, Klára Hosová2, Jaroslav Fojt2, Eva Jablonská3, Petr Slepička4, Karel 4 Tesař5, Jan Drahokoupil1, Vojtěch Hybášek2, Jaroslav Čech5, Jan Blažek2, Emílie Kučerová3, Dinara Sobola6,7, Markéta 5 Straková2, Dalibor Vojtěch2 6 7 1 FZU - Institute of Physics of the Czech Academy of Sciences, Na Slovance 1999/2, Prague 8, 182 00, Czech Republic 8 2 Department of Metals and Corrosion Engineering, University of Chemistry and Technology Prague, Technická 5, Prague 9 6, 166 28, Czech Republic 10 3 Department of Biochemistry and Microbiology, University of Chemistry and Technology Prague, Technická 5, Prague 11 6, 166 28, Czech Republic 12 4 Department of Solid State Engineering, University of Chemistry and Technology Prague, Technická 5, Prague 6 166 13 28, Czech Republic 14 5 Department of Materials, Faculty of Nuclear Sciences and Physical Engineering, Czech Technical University in Prague, 15 Trojanova 13, Prague, 120 00, Czech Republic 16 6 Department of Physics, Faculty of Electrical Engineering and Communication, Brno University of Technology, 17 Technická 2848/8, 61600 Brno, Czech Republic 18 7 Institute of Physics of Materials, Czech Academy of Science, Žižkova 22, 61662 Brno, Czech Republic 19 20 21 * Corresponding author: [email protected] 22 23 Abstract 24 In this study, the effect of laser treatment on the surface state of 3D-printed TiAl6V4 alloy with triply periodic structures 25 was investigated. As-printed and chemically etched samples were used as reference materials to demonstrate surface 26 changes and to illustrate the improvement in surface condition. Microscopic observations of the laser-treated sample 27 confirmed morphological changes caused by the remelting of the initial surface, accompanied by the formation of wave28 like structures during the process. The rapid cooling rate led to the formation of cracks that extended through the entire 29 depth of the newly formed layer. Chemical and phase analyses revealed the predominant presence of TixOy (x = 1–2; y = 30 1–3) compounds, with observed differences between the as-printed and treated samples. Both surface modification 31 approaches successfully removed un-sintered powder particles; however, laser processing did not thin the struts and walls, 32 which positively affected mechanical properties. The best mechanical properties were found in the untreated samples, but 33 laser-treated samples showed only a minimal reduction. For example, an untreated diamond-structured sample exhibited 34 a yield strength of 93 ± 2 MPa, whereas a laser-treated sample showed a yield strength of 89 ± 3 MPa. Furthermore, the 35 hardness of treated surface increased approximately by 60 % in comparison to as-printed state. Samples with a gyroid 36 structure demonstrated significantly increased yield strength, compressive strength, and ductility compared to those with 37 a diamond structure. Corrosion resistance testing did not reveal the occurrence of localized corrosion, indicating that laser 38 surface treatment did not negatively impact the corrosion resistance of TiAlV alloys. The relative metabolic activity on 39 laser-treated materials exceeded the 70% normative limit relative to the metabolic activity of control cells. Thus, all the 40 tested materials can be considered cytocompatible, and therefore the laser treatment is promising for Ti-based 41 personalized 3D implants. 42 43 Keywords: Triply periodic structures; Ti-6Al-4V; Surface treatment; Pulse laser; Characterization 44 45 1. Introduction 46 Titanium and its alloys belong among the most important engineering materials used in various applications. 47 They are particularly attractive for highly demanding applications in the aerospace industry and in medicine, thanks to 48 their exceptional and well-balanced properties, including high specific strength (strength to weight ratio), corrosion 49 resistance, or good biocompatibility [1, 2]. Within the category of metallic materials for biomedical applications, pure 50 titanium (α phase) demonstrates the highest biocompatibility. However, its relatively low mechanical properties preclude 51 its use in load-bearing components, such as hip, knee, or spinal implants, in its unalloyed form. Nevertheless, commercial 52 pure titanium could be still used for skull implants or bone plates because pure titanium is free from toxic alloying 53 elements and exhibits higher ductility than alloys [3]. To enhance its mechanical performance, titanium is often alloyed 54 to create materials with superior strength, albeit at the expense of reduced biocompatibility. The Ti-6Al-4V alloy is indeed 55 one of the most commonly studied metallic materials for biomedical industry, primarily due to its relatively low density, 56 high strength, and good corrosion resistance [4]. The microstructure of Ti-6Al-4V alloy contains mixture of α and β phase 57 with α phase being dominant [5, 6]. The presence of α phase results in e. g. poor wear properties [5]. Another drawback 58 is also the possible toxic effect of Al and V for the body tissues [3]. Despite the mentioned disadvantages, the Ti-6Al-4V 59 alloy remains the most used and studied alloy for implants. 60 2 Stress shielding effect is one of the primary obstacles hindering the seamless use of these materials in 61 implantology. This effect is caused by the mismatch in Young's modulus between the implant and the bone leading to the 62 implant failure [3, 7]. The issue of stress shielding can be mitigated by introducing a porous structure into the materials. 63 A porous structure can be achieved through various methods, such as using powder metallurgy, where a pore-forming 64 agent is added to the powder mixture [7-9], or through 3D printing and foaming technologies [8, 9]. Furthermore, 65 materials with a porous structure enhance osseointegration by promoting the ingrowth of bone tissue into the implants [3, 66 9, 10]. Porosity thus significantly enhances the integration between the implant and the bone because porous materials 67 contain opened or closed voids in the structure resulting in the spaces between them [9]. The pore size is crucial for 68 effective cell adhesion to the implant. The shape of the pores also plays an important role; for instance, pores with a cubic 69 cross-section achieve better cell adhesion than those with a spherical cross-section [9, 11]. Additionally, different tissues 70 have varying requirements for implants, which can sometimes conflict with each other. There is no universal consensus 71 on the ideal porosity, pore size, or their interconnectivity. In designing these porous structures, two main approaches are 72 used: regular and irregular structures. 73 As was mentioned above, the porous titanium and titanium alloys could be produced using powder metallurgy 74 or additive technology. Powder metallurgy is a relatively simple and efficient method for fabricating porous implants. 75 Compared to alternative manufacturing techniques, it is more cost-effective and reduces material waste to a minimum. 76 Moreover, powder metallurgy enables the production of intricate porous structures [9]. On the other hand, the production 77 of porous structures through 3D printing is based on CAD models. This approach achieves high precision and allows for 78 the precise customization of the elemental unit. As a result, it is possible to obtain highly complex yet accurate products. 79 These 3D-printed porous structures are applicable in medicine for the manufacture of patient-specific implants [6, 9]. 80 Additive manufacturing technologies, commonly known as 3D printing, for titanium implants can be categorized into 81 two groups: powder bed fusion (PBF) and directed energy deposition (DED) [3, 6, 12, 13]. Powder bed fusion is further 82 subdivided into laser powder bed fusion (L-PBF), referred to as selective laser melting (SLM), and electron beam powder 83 bed fusion (E-PBF). SLM, aimed at reducing the elastic modulus while preserving biomechanical properties through the 84 incorporation of micropores, is the most widely used method for manufacturing titanium implants [6, 14, 15]. In SLM, a 85 focused laser beam selectively scans and melts a defined layer of the powder bed, causing the powder particles to fuse 86 according to the CAD data of the specified slice geometry [6, 16]. The main drawback is the presence of surface defects 87 and residual stresses [2, 13, 17, 18], which can lead to issues with the subsequent biological response (osteoblastic cell 88 adhesion, growth, or differentiation) [19]. The interactions between biomaterials and tissue, as well as the process of bone 89 healing, are significantly influenced by the chemical composition and surface topography of the implanted material. SLM 90 often results in final products that require post-processing surface treatments, as partially melted particles typically remain 91 on the surface following the printing stage [13]. These particles pose the risk of being released and potentially circulating 92 within the human body [20]. Un-melted particles can be effectively removed by sandblasting or etching, which has been 93 shown to have additional beneficial effects. For instance, surfaces that are roughened through sandblasting and acid94 etching have been shown to promote bone cell proliferation and enhance the secretion of extracellular matrix, thereby 95 supporting bone regeneration [3, 21]. The presence of partially melted particles on the surface can be removed through 96 double acid etching [20]. In addition to the previously mentioned sandblasting and acid etching, the surface can also be 97 modified using a plasma spraying, hot isostatic pressing (HIP), vibratory finishing, or laser polishing [6, 13, 16, 22, 23]. 98 A sandblasted and acid-etched surface is considered one of the most effective modifications for commercially pure 99 titanium, featuring a moderately rough topography that promotes strong interaction with bone tissue. This surface 100 structure has been widely accepted for its ability to enhance osseointegration and improve the overall performance of 101 titanium implants, despite the fact that the process negatively impacts corrosion resistance due to slight surface changes 102 [21, 24]. 103 There are numerous publications dedicated to the modification of the surface of 3D-printed products, with 104 particular focus on chemical etching [24-31]. In work [25] was shown that a mixture of HF and HNO3 is more effective 105 for uniform etching compared to the use of HF alone. Acid etching should be also performed after sandblasting, as the 106 sandblasting process induces the formation of microcracks, which are not readily apparent despite the surface appearing 107 smooth [26]. Moreover, acid etching increases internal porosity, while the thickness of the struts decreases, resulting in a 108 reduction of mechanical properties [28]. Study [27], on the other hand, demonstrated the differences between three distinct 109 surface modification techniques: chemical polishing, electropolishing, blasting, a combination of electropolishing and 110 blasting, and finally, laser polishing. In all cases, the surface roughness was greater when laser polishing was applied. In 111 addition, a laser beam can be focused onto a metallic surface to perform a wide range of treatments, including remelting, 112 alloying, and cladding, which are employed to enhance the wear and corrosion resistance of titanium alloys [23, 32, 33]. 113 Based on the existing publications, laser-based surface modification appears to be the most effective method; however, 114 the majority of research focuses primarily on chemical etching, and thus far, laser technology has only been applied to 115 compact materials, not porous ones. Turning and milling were also investigated and both operations significantly reduced 116 surface roughness [34]. 117 3 The customization of surface texture, along with the post-processing optimization and modification of surfaces, 118 represents the future direction in the development of 3D-printed biomaterials. This study aims to characterize samples 119 with triply periodic structures whose surface was influenced by laser treatment, which was used for the first time on 120 porous alloys. Compared to chemical etching, laser processing offers the advantage of particle removal and melting 121 followed by rapid cooling, leading to surface refinement, and hardening. The surface of the samples was analyzed using 122 advanced techniques such as X-ray photoelectron spectroscopy (XPS) and Raman spectroscopy. Furthermore, the 123 morphology, structure, mechanical properties, and corrosion resistance were investigated, along with essential 124 cytotoxicity tests for biomaterials. 125 126 2. Experimental 127 2. 1 Preparation of samples and surface treatment 128 Atomized powder of the TiAl6V4 (wt. %) alloy was used for the preparation of TiAl6V4 scaffolds with triply 129 periodic minimal surface structures using the SLM method. The alloy powder was prepared by a gas atomization process 130 under a protective Ar atmosphere. The particle size distribution of atomized powders was determined by laser diffraction 131 analysis using SYNC particle analyzer (Microtrac Retsch GmbH) combining laser diffraction with dynamic image. 132 Measurement was performed by the wet (dispersion medium - isopropanol) and dry method. The particle size distribution 133 was characterized by d-values (d10, d50, and d90) which are the intercepts for 10%, 50% and 90% of the mass frequency 134 distribution. The particles had a spherical shape with d10 = 23.2 µm, d50 = 32.9 µm, d90 = 48.1 µm. The SLM process was 135 performed using an M2 Cusing printer, with the process parameters set as follows: laser power of 200 W, scanning spacing 136 of 30 µm, and scanning speed of 800 mm·s⁻¹. The process was performed under Ar protective atmosphere. Samples were 137 manufactured with diamond or gyroid structures and a unit size of 1.35 and 2 mm, respectively (Fig. 1). 138 139 Fig. 1: The structure of a) diamond and b) gyroid unit cells; c) macrostructure of the 3D-printed gyroid sample, 140 with an arrow indicating the build direction 141 142 As-printed materials with diamond (D) and gyroid (G) structures were processed by laser treatment (L) and 143 chemical etching (E) to remove adhered spherical powder particles from the surface of the scaffold and decrease the 144 overall roughness of the last layer. For the laser treatment (DL, GL), an excimer laser, Coherent Leap 100K, was used in 145 pulse mode with a frequency of 1 Hz. Each sample was treated with 50 pulses and an energy of 3 J·cm⁻². Exposed area 146 of the sample in one step was 5 x 3 mm2, i.e. 15 mm2, and the large area exposure was achieved by step motor with sample 147 movement in air. Chemical etching (DE, GE) was performed in a solution with the following composition: 20 mL HF, 148 200 mL HNO₃, and 780 mL H₂O. The composition of the etching bath was adopted from Fojt et al. [35]. The process was 149 carried out in an ultrasonic bath for 10 min with 30 mL of the etching solution at ambient temperature. After etching, the 150 samples were cleaned with distilled water and ethanol in the ultrasonic bath to effectively clean the entire sample. The 151 etched samples were also remeasured and weighed after the process to evaluate the dimensional changes caused by the 152 treatment. 153 154 2. 2 Surface characterization and microstructure 155 4 The morphology of the samples was examined using a JEOL JSM-IT500HR scanning electron microscope. Laser 156 scanning confocal microscopy (LSCM) was employed to measure surface roughness and generate height/laser intensity 157 maps using an Olympus LEXT OLS5000, equipped with a 50x long working distance objective. 158 X-ray diffraction (XRD) measurements were conducted using a PANalytical X'Pert Pro powder diffractometer 159 with a cobalt anode (λ = 1.789 Å). The measurements utilized the Bragg-Brentano geometry with a 1° divergence slit, 160 0.02 rad Soller slits, a β-filter, and a linear detector. Data processing was performed using the Rietveld refinement method 161 within the TOPAS software. 162 X-ray photoelectron spectroscopy (XPS) was performed using an AXIS Supra device with an emission current 163 of 10 mA. For each sample, two separate measurements were taken. The resulting spectra were analyzed and fitted using 164 CasaXPS software. 165 The Raman spectra were measured using a Raman dispersive spectrometer from Thermo Scientific, model DXR 166 Microscope, which is equipped with an Olympus confocal microscope. A 532 nm wavelength laser with an input power 167 of 10 mW was used as the excitation source. A 900 grooves per mm grating was applied. The detector was a multichannel 168 thermoelectrically cooled CCD camera. Samples were measured at 50x magnification with a measuring spot size of 169 approximately 1 µm². Measurements were conducted with a power of 9 mW, a measurement time of 10 s, and with 10 170 accumulations of the spectrum. The Omnic 9 software (Thermo Scientific) was used for spectrum processing. Four spectra 171 were measured for every sample. 172 The microstructure and the number of adhered particles were observed using a Nikon Eclipse MA200 173 metallographic microscope and a Tescan Mira II scanning electron microscope (SEM) equipped with an energy-dispersive 174 spectrometer (EDS). Samples for microstructural observations in the normal direction were embedded in EpoxyCure™ 175 2 resin and ground using P80–P2500 sandpapers. Subsequently, the samples were polished using D2 diamond paste (2– 176 3 µm) and etched with Kroll's reagent (5 mL HNO₃, 10 mL HF, 85 mL H₂O). 177 178 2. 3 Mechanical properties 179 Compressive tests were performed using cylindrical samples with a diameter of 14 mm, a height of 6 mm, and a 180 loading speed of 0.035 mm·s⁻¹ at ambient temperature. A LabTest 5.250SP1-VM universal testing machine was used for 181 testing. Nanoindentation tests were performed to determine hardness (HIT) and indentation elastic modulus (EIT) of surface 182 area affected by laser and bulk material using NHT2 nanoindentation tester equipped with Berkovich diamond tip. 183 Maximum applied load was 2 mN, loading cycle consisted of 10 s loading, 5 s hold at maximum load, and 10 s unloading. 184 Measured load-depth data were evaluated using Oliver-Pharr method [36] according to standard ISO 14577. At least 8 185 valid measurements were performed and statistically processed in every investigated area. 186 187 2. 4 Corrosion tests 188 Before measurement or surface treatment, every sample was degreased for 5 min in an ultrasonic bath in a 189 detergent solution, followed by 5 min in ethanol. 190 The susceptibility to non-uniform corrosion was measured according to the ASTM F2129 standard in phosphate191 buffered saline (PBS) (see Table 1) using a three-electrode setup with a reference Ag/AgCl electrode (3 mol · L-1 KCl, 192 abbreviated as SSCE). Due to the high real surface area of the working electrode, a platinum mesh counter electrode was 193 used during the measurement. Before starting the measurement, dissolved oxygen was removed by bubbling the 194 electrolyte with nitrogen gas for at least 30 min. After stabilizing the open circuit potential (OCP) for 1 h, the polarization 195 scan rate was set to 1 mV · s-1, and the samples were polarized to 0.83 V/SSCE. Once the potential reached 0.83 V/SSCE, 196 the scan rate was reversed until the potential returned to the initial values, at which point the experiment was terminated. 197 Due to the unknown real surface area, a geometrical area of 1.77 cm2 was chosen for all measured samples, which were 198 in compact form. 199 200 Table 1: Composition of phosphate-buffered saline (PBS) 201 Chemical reagent Concentration (g · L-1) NaCl 8 KCl 0.2 Na2HPO4 1.15 KH2PO4 0.2 For longer-term exposure, a physiological saline solution containing 9 g · L-1 NaCl (FS) was chosen due to its 202 higher Clconcentration and lack of possible precipitates that could form on the sample surface. The electrode setup 203 included a reference SSCE electrode with a tip filled with agar gel along with two glassy-carbon counter electrodes. 204 During exposure, the OCP and polarization resistance (± 20 mV/OCP, 0.125 mV · s-1) of each sample were monitored to 205 detect any electrochemical changes occurring on the surface. To assess susceptibility to localised corrosion after exposure, 206 a cyclic polarization scan (CP) was performed at a rate of 1 mV·s-1 with a potential scan reversal at 0.83 V/SSCE. 207 5 2. 5 Test on extracts (ISO 10993-5) 208 Samples were cleaned by sonication in EtOH (96%, 15 min) and distilled water (15 min) and then sterilized by 209 autoclaving (121 °C, 20 min). Replicates of Ti-Al-V material with surface modification (TiAl6V4 - laser) and raw Ti-Al210 V material (TiAl6V4 - as printed) were immersed in the test media – MEM (minimal essential media, Sigma-Aldrich, 211 cat. no. M0446) with 5% foetal bovine serum (FBS) or DMEM/F-12 Ham’s (Sigma-Aldrich, cat. no. D6434) with 5% 212 FBS and 2.5 mmol · L-1 Alanyl-glutamine. The immersed materials were then incubated in tubes at 37 °C with constant 213 shaking (100 rpm) for 24 h. The S/V ratio was 1.25 cm² · mL-1. 214 Cell lines of mouse fibroblasts (L929, recommended by the ISO 10993-5) and human foetal osteoblasts (hFOB 215 1.19) were trypsinized, resuspended in appropriate culture medium (L929: MEM with 10% FBS or hFOB 1.19: DMEM/F216 12 Ham’s with 10% FBS and 2.5 mmol · L-1 Alanyl-glutamine), and diluted to a concentration of 1·10⁵ cells · mL-1 (L929) 217 or 2·10⁵ cells · mL-1 (hFOB 1.19). These cell suspensions were then pipetted (100 µL) into the wells of a 96-well plate 218 and incubated for 24 h in the CO2 incubator under standard conditions. The next day, media were aspirated from the 96219 well plates, and 100 µL of extracts from the samples were pipetted onto the cells. Control cells received 100 µL of the 220 test media alone. A positive control for cytotoxicity was prepared with test media containing 0.2% Tween. The cells were 221 then placed in a CO2 incubator for another 24 h under standard conditions. On the third day, the metabolic activity of the 222 cells was measured using resazurin solutions [37]. The wells in the plates were emptied and the cells were gently washed 223 with 100 μL of sterile phosphate buffer. Then, 100 µL of a solution of resazurin (25 µg · mL-1 in culture media without 224 phenol red) was pipetted to the cells. The cells were incubated for ~2 h in the CO2 incubator until the colour of the media 225 went from blue to purple. Fluorescence of resorufin was then measured with a Fluoroskan Ascent FL (Ascent Software). 226 The excitation wavelength was 544 nm, and the emission wavelength was 590 nm. Cytotoxicity of the extracts was 227 evaluated as the relative metabolic activity of the cells compared to control cells. Extracts that reduced metabolic activity 228 below 70% of the control cells' activity were considered cytotoxic. Statistical evaluation was performed using GraphPad 229 Prism 7.0. The statistical significance of differences was determined by ANOVA followed by Tukey's multiple 230 comparison test. 231 232 2. 6 Direct contact test (quantitative) 233 Samples were transferred to 48-well plates (modified surface facing upwards). hFOB 1.19 cells were trypsinized, 234 resuspended in culture media, and diluted to a concentration of 2·10⁵ cells/mL. This cell suspension was pipetted onto 235 the samples. The 48-well plates were then placed in the CO2 incubator under standard conditions. The metabolic activity 236 of the cells was measured at three time points of cultivation – after 24 h, 72 h, and 6 days, using resazurin [38] as described 237 above. Statistical evaluation of the metabolic activity of the cells on TiAl6V4 - as printed and TiAl6V4 - laser was 238 performed using GraphPad Prism 7.0. The statistical significance of differences was determined by ANOVA followed by 239 Tukey's multiple comparison test. 240 241 2. 7 Qualitative cytocompatibility evaluation by SEM 242 hFOB 1.19 cells were trypsinized, resuspended in culture media and diluted so that 24,000 cells · cm-² were 243 seeded onto the TiAl6V4 - as printed and TiAl6V4 - laser samples in the 48-well plates. The plates were then placed in 244 the CO2 incubator under standard conditions. At three time points (24 h, 72 h, and 6 days, consistent with the evaluation 245 of the contact tests), media were aspirated from the wells, and the samples with cells were transferred to clean wells, 246 washed twice with sterile phosphate buffer, and cell fixation was performed [39]. Cells were fixed using Karnovsky’s 247 solution for 1.5 h at ambient temperature. The samples were then transferred to clean wells and washed twice with 0.1M 248 cacodylate buffer. Cells fixed on the samples were dehydrated through a 10-min incubation in increasing concentrations 249 of EtOH (50%, 70%, 90%), followed by a wash in 99.8% EtOH and final incubation in 99.8% EtOH for 5 min. The 250 samples with cells were then dried using CPD with CO2 and acetone. Finally, the samples with cells were coated with 10 251 nm of gold. 252 253 254 3. Results 255 3.1 Morphology 256 Images of the surfaces of diamond and gyroid samples in their as-built state are shown in Fig. 2. The arrangement 257 of struts and walls in diamond and gyroid structures plays a significant role under loading conditions. The gyroid structure 258 is known for its smooth, continuous geometry, free from sharp edges or transitions [40]. This seamless continuity ensures 259 a more uniform distribution of mechanical stresses and can be advantageous in applications where tensile and compressive 260 strength are critical [41]. Additionally, it reduces stress concentration points, thereby extending the lifespan of 261 components. The diamond structure, on the other hand, consists of individual struts and nodes, resulting in sharp 262 transitions between parts of the structure. These sharp edges can be prone to stress accumulation, which may lead to 263 structural failure at these points under high loading conditions. The print quality can be assessed based on surface 264 morphology, which is influenced by residual un-melted metallic powder. These powder residues are particularly visible 265 6 in the detailed images shown in Fig. 2. The images suggest that the samples were similar in terms of surface quality and 266 morphology. 267 268 269 Fig. 2: SEM images of a) diamond and b) gyroid structures with details of the surface morphologies 270 271 Processed samples 272 The surface of etched and laser-treated sample could be observed in Fig. 3. In the case of DE sample, a significant 273 reduction in the presence of surface particles can be observed, as shown in Fig. 3 a. The sample treated with chemical 274 etching shows a notable reduction in the thickness of the struts / walls forming the structure (Fig. 3 a, 3 b), which 275 subsequently affects its mechanical properties negatively (see Fig. 9). Additionally, the composition of the etchant can 276 influence the chemical composition of the surface, such as the presence of fluorides. Surface analysis using EDS revealed 277 an increased fluorine content of around 1 wt. %. Moreover, nano-submicropores could be observed on the surface (Fig. 278 3 a). On the other hand, the surface was smoother after etching. The application of etching is also associated with mass 279 loss due to the removal of un-melted particles [25]. When comparing chemically etched and untreated surfaces in Fig. 2 280 and 3, a significant improvement in terms of adhered particles is visible, although this improvement is unlikely to be as 281 pronounced in the deeper layers of the material. As shown in Fig. 3 a, a higher concentration of particles is already evident 282 in the second layer. This is due to the reduced flow of the etchant into the deeper layers of the structure, despite the 283 experiment being conducted in an ultrasonic bath. The reduction in the thickness of struts and walls within the material 284 is listed in Table 2. 285 286 Table 2: Reduction of thickness, d1 – thickness before etching, d2 – thickness after etching 287 sample d1 (mm) d2 (mm) reduction of thickness (%) G1 0.273 ± 0.036 0.148 ± 0.024 45.9 D1 0.272 ± 0.026 0.174 ± 0.027 35.8 288 Processing using laser appears to be effective in removing of un-melted particles from the surface layer. Surface 289 melting occurred during laser-treating, which led in the formation of a relatively smooth area and lower surface roughness. 290 Laser processing can also help densify the surface, thereby improving the surface quality, which would lead to better 291 resistance to mechanical damage and material fatigue. However, a limitation of the method is that melting is confined to 292 the region of the laser spot and affects only the upper layer of the printed structure. The laser does not penetrate deeper 293 layers due to the blocking effect of the top layers. Fig. 3 c, d shows the laser spot and provides representative images with 294 the designated area enclosed by a dashed line. 295 296 7 297 Fig. 3: SEM images of a), c) diamond and b), d) gyroid structures modified by a), b) etching and c), d) laser; the area 298 enclosed by dash line shows treated area within the surface of the material 299 300 Typical SEM backscatter electron images and LSCM laser intensity and height maps of the top area of diamond 301 structures are shown in Fig. 4. The impact of etching and laser treatment on surface roughness is evident: etching created 302 an irregular surface characterized by rounded valleys, while laser treatment created a more uniform surface. The untreated 303 surface depicted the residual un-melted metallic powder, apparent even on the most exposed top area of the diamond 304 structure, resulting in the Sa value of (11.8 ± 1.3) µm. When the etched samples are considered, rounded valleys were 305 created as the result of the etching process, with no apparent residual powder in the top layer and the Sa value of (11.1 ± 306 2.5) µm. Although the etched surface possessed a similar roughness to the untreated surface, the weakening of the 307 structure's robustness is observable, as well as some residual powder particles in the deeper layers. The laser-treated 308 structure, when compared to the previous states, possessed the lowest surface roughness of Sa = (7.3 ± 0.2) µm, with a 309 low standard deviation. Typical morphology of the ripple patterns and thermal-induced cracks dominated the newly 310 formed surface already shown in [42], with no residual powder remaining on the top areas and partially melted residual 311 powder particles present in other areas. As for the gyroid structures, the deviations of roughness parameters were rather 312 large, as it is difficult to measure comparable areas between the samples. Yet, the qualitative morphological features were 313 equal to both structures. 314 8 315 Fig. 4: SEM backscatter electron images, LSCM laser intensity and height maps of as-printed, etched and laser-treated 316 samples 317 318 3.2 Surface analysis 319 320 XRD 321 The phase composition of the diamond (Fig. 5 a) and gyroid (Fig. 5 b) as-printed and processed structures was 322 analyzed using XRD. As shown in Fig. 5, clear differences can be observed between the as-printed and laser-treated 323 structures. Specifically, the primary difference observed was the presence of TiO (39 wt.%) in the samples within the 324 interaction volume of laser-affected layer. Importantly, this difference was noted regardless of the macrostructure of the 325 as-printed samples. On the other hand, identical phase composition was found in the as-printed and etched samples. This 326 result is expected, particularly due to the absence of thermal effects during etching. More precisely, preferential 327 dissolution of the material during etching may have occurred due to the microstructure and internal stresses introduced 328 by the melting of particles during printing. As a result, uneven dissolution of material may have occurred throughout the 329 volume, but the formation of nonstoichiometric compounds like TiO—formed by the rapid cooling of the melted surface 330 in an oxygen-rich environment—did not take place. Furthermore, differences in crystallite size and microstrain of 331 hexagonal Ti phase were also observed, as indicated by the varying peak widths in Fig. 5. In particular, significantly 332 smaller crystallite size (11 nm) and higher microstrains (0.32%) were recorded in the laser-processed samples compared 333 to the as-printed samples (26 nm, 0.25%) and etched samples (29 nm, 0.22%). Meanwhile, minimal differences between 334 the diamond and gyroid samples were identified, as was the case for phase composition. Overall, all results can be 335 explained by the cooling rate and environmental conditions. The SLM process was conducted under a protective 336 atmosphere, with the printed sample surrounded by powder particles. This setup, in turn, is believed to reduce the cooling 337 rate by gradually increasing the temperature during the process. Conversely, laser processing was performed in air, where 338 oxygen can immediately react with the melted surface. Faster cooling occurred in this case, not only due to heat dissipation 339 9 into the air but also because the affected volume was significantly smaller. Therefore, a higher cooling rate was achieved 340 in the laser-treated samples, resulting in smaller crystallites and higher strains in the structure, which ultimately led to the 341 formation of cracks in the affected layer. Because the results of diffraction showed minimal changes in the phase 342 composition between etched and as-printed samples, further detailed study was conducted on the alloy whose surface was 343 treated with a laser. The results were consistently compared with the as-printed state. 344 345 Fig. 5: Diffraction patterns of as-printed, etched and laser-treated structures 346 347 XPS 348 XPS measurements were performed using as-printed and laser-treated samples, and the results from these 349 measurements are shown in Fig. 6. The tests revealed that during processing, the thickness of oxides present on the sample 350 surface increased. This is evident from the Ti and Al spectra, where the peak for Ti⁰ disappeared and the Al⁰ peak was 351 significantly reduced after processing. In conjunction with the increase in oxygen content (from 23.3 to 30.4 wt.%), the 352 transformation to oxides due to the increased reactivity of the surface caused by the laser could explain these changes. 353 Furthermore, the results suggest the presence of TiN and TiC on the surface of both as-printed and laser-treated samples. 354 However, the overall content of N and C decreased after treatment, from 5.2 wt.% to 3.7 wt.% and from 44.7 wt.% to 355 31.8 wt.%, respectively. This suggests that the laser may reduce the contamination created on the surface during the 356 printing process. The most significant change was observed within the V spectra, where the concentration of V increased 357 from 0 wt.% in the as-printed state to 0.5 wt.% in the laser-treated state. This change could be attributed to the activation 358 of diffusion, where V atoms are transported toward the surface while Al is transported in the opposite direction. This 359 phenomenon was described by Lindwall et al. [43]. This could also be a contributing factor in the reduction of the Al⁰ 360 peak intensity. The results indicate that the contamination originated from the fabrication of the sample rather than from 361 the laser treatment, suggesting no significant negative effects of the process on the surface layer of the samples. 362 16 491 Fig. 12: Relative metabolic activity of the a) L929 and b) hFOB cell lines incubated in extract of as printed and 492 laser treated TiAl6V4 alloy. Standard deviations were calculated as an average of a minimum of three independent 493 measurements. Statistically significant differences between metabolic activity of Control and TiAl6V4 - as printed (**, p 494 ≤ 0.01), TiAl6V4 - laser (*, p ≤ 0.05) are visible in a) 495 496 From a qualitative perspective, including cell morphology and interactions with the material, there was no 497 difference between hFOB 1.19 cells growing on TiAl6V4 - as printed and TiAl6V4 - laser materials. The cell was spread 498 out, exhibiting normal morphology and proliferated equally well in both cases (see images after 144 h). However, it was 499 noted that on TiAl6V4 - as printed, the cells were growing only between surface particles, not on them. In the image of 500 the TiAl6V4 - laser material at the 72-h time point (Fig. 13 e), it is interesting to focus on the detail of the lasered structure, 501 where the resulting 'wave-like' pattern appears to be particularly favorable for the cells. This structure is approximately 502 the same size as the cell width, thus creating a suitable environment for cell adhesion and growth. The metabolic activity 503 of hFOB 1.19 cells cultured in the presence of TiAl6V4 - as printed and TiAl6V4 - laser materials was not statistically 504 significantly different. A statistically significant difference was observed between the metabolic activity of control cells 505 and cells cultured in the presence of TiAl6V4 - as printed as well as TiAl6V4 - laser materials. Nevertheless, it was above 506 the normative limit of 70% in all cases, relative to the metabolic activity of control cells. However, it can be inferred that, 507 unlike cells cultured in the presence of TiAl6V4 - as printed material, culturing cells in the presence of TiAl6V4 - laser 508 over a longer time interval leads to leveling of metabolic activities with those of control cells (without the presence of 509 material). The difference in metabolic activities of control cells and cells in the presence of materials may also be due to 510 slight movement of the samples in the wells, which could lead to the death of a small portion of cells growing in direct 511 proximity to the materials. 512 513 17 514 Fig. 13: Direct cytotoxicity testing of (a), (b), (c) as-printed and (d), (e), (f) laser-treated samples, with (g), (h), 515 (i) corresponding histograms of hFOB 1.19 cell viability evaluated at different time periods of 24, 72, and 144 h 516 517 4. Conclusion 518 The presented study investigated the effect of pulsed laser on the surface state of 3D-printed triply periodic 519 structures in TiAl6V4 alloy. The results were compared with the as-printed condition and the most widely used surface 520 treatment method to date, namely etching. The findings arising from this publication can be summarized into the following 521 points: 522 523 1) Laser surface treatment successfully removed the adhered un-melted powder particles without affecting the thickness 524 of the supports and walls. 525 526 2) Due to the rapid cooling rate, a new fine-grained layer is formed, and the surface is strengthened by residual stresses. 527 However, cracks were observed in the newly formed layer. 528 529 3) Mechanical property tests revealed that laser surface treatment did not have a negative impact on yield strength, 530 ultimate compressive strength, or ductility. The measured values were close to those obtained for the initial as-printed 531 condition. Additionally, it was found that the gyroid structure exhibits significantly higher mechanical properties 532 compared to the diamond structure. Furthermore, nanointendation tests revealed significant increase (60 %) of hardness 533 value within laser-treated layer. This suggests that laser treatment is superior to etching in this context. 534 535 4) Based on the provided data, all samples exhibited metastable disturbances in the passive layer; however, none showed 536 the development of localized corrosion. The corrosion potential values were within the stability range of titanium oxide. 537 Surface laser treatment resulted in an insignificant decrease in polarization resistance for both types of structures. The 538 applied processing does not have a negative impact on corrosion resistance. 539 540 5) The relative metabolic activity on laser-treated materials significantly exceeded the 70% normative limit relative to the 541 metabolic activity of control cells. 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