Metallic Implant Surface Activation through Electrospinning Coating of Nanocomposite Fiber for Bone Regeneration

There is a critical need in orthopedic and orthodontic clinics for enhanced implant-bone interface contact to facilitate the quick establishment of a strong and durable connection. Surface modification by bioactive multifunctional materials is a possible way to overcome the poor osteoconductivity and the potential infection of Ti-based implants. Ti-25Zr biometallic alloy was prepared by powder metallurgy technique and then coated by Nano-composite fiber using electrospinning. Ceramic Nanocompound (CaTiO3, BaTiO3) was used as filler material and individually added to polymeric matrices constructed from the blend of polycaprolactone/chitosan. Using optical microscopy, scanning electron microscopy (SEM), energy-dispersive X-ray spectroscopy (EDX), Fourier transform infrared spectroscopy (FTIR), and wettability, respectively, the morphology, chemical analysis, surface roughness, and contact angle measurements of the samples were evaluated. The result shows a significant improvement in cell viability, proliferation, and ALP activity for coated samples compared to noncoated samples. PCL/Chitosan/Nano-CaTiO3 (CA1) recorded remarkable enhancement from the surface-coated samples, demonstrating a significantly higher cell viability value after seven days of MC3T3-E1 cell culture, reaching 271.56 ± 13.15%, and better cell differentiation with ALP activity reaching 5.61 ± 0.35 fold change for the same culture time. PCL/Chitosan/Nano-BaTiO3 (BA1) also shows significant improvement in cell viability by 181.63 ± 17.87% and has ALP activity of 3.97 ± 0.67 fold change. For coated samples, cell proliferation likewise exhibits a considerable temporal increase; the improvement reaches 237.53% for (CA1) and 125.16% for (BA1) in comparison with uncoated samples (bare Ti-25Zr). The coated samples resist bacteria in the antibacterial test compared to the noncoated samples with no inhibition zone. This behavior suggests that a Nanocomposite fiber coat containing an active ceramic Nanocompound (CaTiO3, BaTiO3) promotes cell growth and holds promise for orthodontic and orthopedic bioapplication.


Introduction
Bone is a living tissue that is formed of a hierarchically structured composite of highly organized collagen bundles reinforced by hydroxyapatite nanocrystals [1]. A bone defect is a lack of bone tissue in the body where it should normally be. Tis could be caused by several congenital or acquired conditions, such as trauma, tumor resection, or infections [2]. Unlike other tissues, most bone injuries can heal spontaneously and without additional treatment due to the bone's ability to regenerate itself [3].
However, for most surgeons, treating critical bone defects with a length greater than 1.5 times the diameter of the long bone is challenging. Failure to cure a bone defect is primarily caused by bone loss and harm to the physiological environment [4]. Although autologous bone graft and bone allograft are the two most common options for bone regeneration, their clinical benefts are not guaranteed. Tis is because complications and morbidity are frequently encountered in patients. Terefore, research on alternative bone substitutes is still necessary, and synthetic bone substitutes are still considered for bone tissue regeneration. Synthetic bone substitutes are used because of their biocompatibility, osteoconductivity, and potential to overcome the previously mentioned limitations of autologous and allogenic bone grafts [2].
Materials used for bone regeneration should have the characteristics of favorable osteoinduction, good osteoconduction, regulated bioactivity, and an adequate degradation rate to imitate the native bone structure [5].
Because of their exceptional mechanical qualities, superior biocompatibility, and excellent corrosion resistance, commercially pure titanium and Ti-based alloys have been widely employed in hard tissue replacement metallic implants, especially in the orthopedic and dental professions [5][6][7][8]. Among the most recommended reasons for boneimplant failure are stress shielding, inadequate osseointegration, and a high potential for bacterial infections. Since the modulus of elasticity of the implant is signifcantly greater than that of the surrounding bone, stress shielding is typically caused by this diference. Tis substantial variance in the modulus of elasticity value prevents the necessary stress from being transferred to adjacent bone, which leads to bone resorption around the implant and, eventually, loosening and failure. Ti-based alloys with nontoxic alloying elements and a low modulus of elasticity have been widely produced to address the issue of stress shielding. Te Ti-Zr alloy is one of the most commonly investigated titanium alloys with, a low elastic modulus and enhanced mechanical and biological properties. Porous titanium alloys have undergone signifcant development in recent years. Additionally, titanium alloys are produced in regulated porous structures, which successfully lower the implant's modulus of elasticity to levels comparable to those of natural bone. Also, it permits new bone to grow within that porous area and achieve mechanical interlocking [9,10].
However, the increase of zirconium element in the Ti-Zr alloy enhances its mechanical properties, but Zr contents exceeding 25 wt.% prevent the formation of calcium phosphate, which is the main component of human bones [11,12]. Also, caution must be taken when considering high Zr concentrations due to an increased susceptibility to pitting corrosion [13].
Ti-base alloys are bioinert materials, and if the implant has not been adequately integrated within the bone, fbrous tissue can readily form between them, reducing the longterm survival of the alloy in the body [14,15]. Furthermore, postoperative implant infections remain a signifcant problem. Tey are produced by wound contamination after surgery and bacterial strains introduced into the surfaces of metallic implants and surgical instruments due to insufcient sanitation. Infections caused by medical devices increase healthcare expenses, cause patient misery, and, in some difcult situations, result in death [6,16].
Bioactive composite coatings on the surface of implants with multiple functionalities to encourage bone formation while preventing bacterial infection play an important role in strengthening osseointegration, shortening healing time, and extending implant life [17,18].
For this purpose, polymeric matrices flled with bioactive ceramic Nanoparticles are often used as composite biomaterials. With this strategy, the processability, biodegradability, and mechanical characteristics of polymers, which are already good for bone tissue, are made even better by adding a bioactive ceramic phase to mimic the composition of natural bone [3,19].
Polycaprolactone was chosen as the polymeric matrix material due to its thermoelastic nature, low melting point, ease of processing, outstanding mechanical strength, and biocompatibility, as well as the fact that it is an FDAapproved biodegradable polymer. To improve stress crack resistance, hydrophilicity, degradation rate, and cell adherence, PCL can be combined with other polymers. PCL combined with other polymers, such as cellulose propionate and cellulose acetate butyrate, has been proven to alter the drug release rate from microcapsules [20][21][22].
Chitosan (CS) is a polysaccharide derived from chitin deacetylation. CS is a natural polymer with antibacterial properties, a high absorption capacity, biodegradability, and biocompatibility. When CS comes into contact with the living tissue, it interacts with several cellular processes during wound healing and has the potential to speed up the healing process [23,24]. Tus, combining PCL and chitosan for the coating process is a desirable method for enhancing biological and mechanical performance compared to using the components individually [25,26].
Because of its biocompatibility, increased apatite bonding, and stimulation of cell adhesion and proliferation, perovskite calcium titanate has lately been employed as the primary coating component on titanium implants. CaTiO 3 has shown promising applications for bone regeneration because it creates phosphate ions and opposite surface charges in a simulated bodily fuid (SBF), which impact future bone-like apatite formation [5,27].
Barium titanate, BaTiO 3 (BTO), is a smart material with a piezoelectric property that generates electrical polarization in reaction to minute structural deformations. BTO is said to have biological properties, including strong biocompatibility when connected with living cells. As a result, it has been identifed as a promising material for biomedical applications [28,29].
In the present work, low-modulus metallic biomaterial Ti-25Zr was created using the powder metallurgy technique. Te inert surface of the Ti-25Zr was activated with a novel multifunctional Nanocomposite fber. Te electrospinning method was used for fabricating hybrid inorganic and organic Nano-fbrous mats with a good bond to the substrate. Diferent Nano-particle ceramic (CaTiO 3 , BaTiO 3 ), each with attractive characteristics for biomedical application, was added as fler material to the blend of PCL/Chitosan to prepare two solutions (PCL/Chitosan/Nano-CaTiO 3 , PCL/ Chitosan/Nano-BaTiO 3 ) for electrospinning coating the surface of Ti-25Zr. Te coated and uncoated samples' surface morphology and elemental composition were studied. Te cytocompatibility of composite coated Ti-25Zr and the noncoated samples was also evaluated by in vitro cell culture.

Substrate Material Preparation.
A Ti-25Zr disc (15 mm in diameter, 2 mm in height) as the base metal was prepared using the powder metallurgy technique. Titanium powder was mixed with 25 wt.% zirconium powder for 6 hours, then the mixed powder was compressed at 500 MPa. Te green compact disc was then sintered in a vacuumed furna-ce10 −3 torr at a range of 10°C/min, held for 2 hours at 1300°C temperature, and left to cool in the furnace.

Surface Treatment of the Base Sample.
Te Ti-25Zr samples were mechanically polished with 120-1000 grit silicon carbide paper. Te samples were cleaned for 20 minutes using an ultrasonic cleaning path with acetone followed by deionized water. Te surface of the metal implant was made rougher and more energetic by treating it with a combination of H 2 SO 4 : HCl : H 2 O(1 : 1 : 1) at 60°C for one hour. In addition, samples were alkaline-treated with 10 M NaOH at 60°C for 24 hours and left to dry overnight.

Porosity Test.
Te porosity of the sintered Ti-25Zr alloys is determined by an equation where ρ is the apparent density of the alloy, which is determined by the liquid displacement method using Archimedes' principle. ρ 0 is the nominal theoretical density of the corresponding alloy, calculated as follows: where A% and B% are the mass fractions of elements A and B, and ρ A and ρ B are the theoretical density of A and B.

Mechanical Evaluation of the Base Sample.
A Vickers digital microhardness tester (HVS-1000, Laryee Technology, China) with a load of 9.8 N and a dwell period of 15 s was used to measure the hardness of the samples. Te local values from 10 points were used to fnd the average microhardness values. Te compressive stress measurement was done at room temperature by the Brazilian method using (a universal tensile machine made in China by Instron) for a sample having (a 15 mm diameter and an 18 mm height). Lastly, the elastic modulus was found by using (an ultrasonic tester of type CCT-4 UK) and solving the following equations: where C trans : is the wave speed transversely, C long : is the wave speed longitudinally, ʋ: is the Poisson modulus,:is modules of elasticity, ρ density of the material (the density value assumes an isotropic, homogenous, and nondispersive material. Te error was estimated using error propagation with a 95% confdence level).

Chemical and Microstructural
Characterization of the Base Sample. Te crystal phase was characterized by X-ray difraction performed with Cu Ka radiation operated at 40 kV and 40 mA at room temperature (XRD, 6000 Shimadzu, Japan). Te microstructures and surface topography of the samples were examined by scanning electron microscopy (SEM, TESCAN VEGA3, Czech Republic). Te chemical composition and homogeneity of sintered samples were examined using energy-dispersive spectrometry (EDX), efciently combining SEM imaging with elemental composition analysis.

Preparation of Electrospinning Solutions.
Chitosan 2% (w/v) was dissolved in 4/6 acetic/formic (v/v) acid (100 ml) using a hot plate magnetic stirrer for 12 hours at 50°C to make a Chitosan solution. PCL 8% (w/v) was added to the Chitosan solution and stirred for 3 hours until a clear and homogenous solution of the PCL/Chitosan blended forms. Te Nano-CaTiO 3 and Nano-BaTiO 3 were added individually, each with 1% (w/v), and stirred for 1 hour to form two solutions. Each solution was then homogenized for 3 min using a homogenizer (Model 300VT Ultrasonic Homogenizer USA) to make the PCL/Chitosan/Nano-CaTiO 3 (CA1) and PCL/Chitosan/Nano-BaTiO 3 (BA1) solutions ready for electrospinning coating.

Electrospinning.
Te solutions were prepared and placed in a 5 ml syringe ftted with a blunt-end 22 G needle. Using an infusion pump, the fuid was expelled at a rate of 1 ml/h (KD Scientifc Syringe Pump 200, USA). Te needle tips' distance from the grounded sample was maintained at 10 cm. Te needle was subjected to a high voltage of 20 kV. Te relative humidity in the room ranged between 35-55%. Before the examination, the fbers were dried overnight and kept in a desiccator.

Characterization of Coatings.
Te microstructures of the electrospun fbers were sputtered with gold before being examined using a feld emission scanning electron microscope (FESEM) (Inspect F-50, Spain) at an accelerating voltage of 15 kV utilizing secondary electrons (SE). At random points on each fber, the diameters of the resultant fbers were measured. Te existence of Nano (CaTiO 3 , BaTiO 3 ) in the PCL/Chitosan polymer blend was confrmed using a dispersive energy X-ray (EDX).

ATR-FTIR.
Fourier transform infrared spectroscopy ATR-FTIR (Bruker Tensor 27 IR, Germany) was used to investigate the functional chemical groups. Te FTIR spectra International Journal of Biomaterials of pure chitosan, pristine polycaprolactone, and diferent types of composite coating were recorded between (4000-500) cm −1 regions using a universal ATR sampling accessory.
2.11. Wettability Evaluation. Te wettability of the coated/ uncoated samples was tested using the sessile drop technique with DD water (Optical Contact Angle SL200KS, China). Tis procedure included dropping 1 μml of distilled water onto the coated surface and measuring the contact angle of the water for 10 seconds. Te test for wettability was performed in triplicate, and the contact angle was measured using a camera-based contact angle meter.

Cell
Proliferation. MC3T3-E1 preosteoblast proliferation was evaluated by determining the cell number in the samples at days 1, 3, and 7 and using the AlamarBlue ® fuorescent assay. At each time point, samples were transferred to a new plate, AlamarBlue ® was added, and the fuorescence was measured. After performing the AlamarBlue ® assay each day, samples were washed twice with PBS and incubated in the osteogenic medium in a humidifed incubator with 5% CO 2 at 37°C. Data were obtained from samples from three independent experiments (n � 3).
2.13.6. Alkaline Phosphate Enzymatic Activity. Alkaline phosphatase (ALP) assay is an essential method of assessing osteogenesis diferentiations. Alkaline phosphatase (ALP) activity was measured to determine the osteoblastic phenotype of MC3T3-E1 preosteoblasts on coated and noncoated samples. On days 3 and 7 of cell culture on the sample's surface, cells were lysed with milli-Q water and freeze-thawed three times to determine ALP activity, and protein content-nitrophenyl-phosphate (Merck, Darmstadt, Germany) at pH 10.3 was used as the substrate for ALP, as described earlier [30]. Te plate was immediately read at 405 nm using a spectrophotometer (BioTek, Winooski, VT) to obtain an absorbance reading correlated with the expression of (pNPP). ALP activity was calculated by dividing the amount of paranitrophenyl phosphate by the protein content. Results were expressed by calculating the fold changes in comparison with the control.  USA). Data are expressed in the form of Mean ± SD. Oneway (ANOVA), as well as the Bonferroni method, was used for comparison between groups. P < 0.05 was considered statistically signifcant.

Microstructures and Chemical Composition of Ti-Zr Alloy.
X-ray difraction analysis (XRD) was used to determine the phase compositions of the alloy developed (Figure 1). shows the typical XRD profles of Ti-25Zr alloys after 2 hours of sintering at 1300°C. Te XRD data indicate that the Ti-25Zr alloy's main phase was the α hcp phase. Te complete solid solution system of Ti and Zr may explain why Ti-25Zr alloys display the α phase, as shown by the optical microscope ( Figure 2). In addition, the higher atomic radius of Zr (1.62 A°) Compared to Ti (1.47 A°), the addition of Zr causes the phase lattice parameters to be raised, resulting in a shift of the peaks of the XRD chart towards a low angle. Tis fnding agrees with those reported [31,32].
Te chemical composition and uniformity of sintered samples were evaluated using energy-dispersive X-ray spectroscopy (EDX). Te fndings of the semiquantitative chemical analysis conducted by the (EDX) in point are shown in (Figure 3). Te EDX examination revealed the homogeneity and purity, indicating that no additional element is present in the powder combination, and (Table 1) shows the proportion of elements, revealing the appropriate mixing procedure.

Surface Treatment.
Tere were apparent morphological diferences between the chemically treated and untreated surfaces of the Ti-25Zr samples. Ground grooves in the surface of the control sample served as a benchmark for comparison (Figure 4(a)). Te grooves were easily visible after being etched with acid and alkali (Figure 4(b)). At the same time, the pit (Figure 4(c)), after treatment, seemed to deepen and sharpen (Figure 4(d)). Te chemically treated surface has more energy and is rougher, consequently enhancing the surface wettability and the coating layer adhesions.

Mechanicals and Physical Characterizations.
Microhardness, tensile, compressive, and Modulus of Elasticity of Ti-25Zr are listed in (Table 2).
Te impact of zirconium contents on the mechanical characteristics of Ti-25Zr alloys was to enhance all mechanical properties over cp-Ti, as previously indicated [31,33,34]. Te toughness of Ti-25Zr alloys rose inversely with Zr content because the substitution of Zr resulted in crystalline lattice deformation and atomic displacement restrictions [31]. Furthermore, the Ti-25Zr alloys were complete solid solutions with hardness increases, most likely generated by solid solution hardening of the α phase and the contribution of the refned microstructure [35].
A modest quantity of Zr signifcantly improves the alloy's compressive and tensile strengths [31]. Two factors were most likely responsible for the increase in compression and tension strength caused by alloying. First, according to the Ti-25Zr alloy phase diagram, the α phase indicated a total solid solution with no intermetallic compound. As a result,    the solid solution mechanism would create more obstacles for the slip system, increasing its mechanical properties. Second, according to the Hall-Petch formula, fne-grain strengthening increases alloy yield strength. Te phase transition starting temperature decreased as Zr increased, inhibiting α phase expansion. Grain refning increased grain boundary area, leading to more excellent resistance to dislocation glide and improved mechanical properties [36].

Fourier Transform IR Spectroscopy.
Te FTIR absorption spectra of polycaprolactone, chitosan, ceramic nanoadditive, and the composite coating of (PCL/Chitosan/ Nano-CaTiO 3 , PCL/Chitosan/Nano-BaTiO 3 ), electrospun fber are shown in (Figures 7(a), 7(b)) respectively. Te primary peak in the pristine PCL spectra is at 1723 cm −1 , which relates to the carbonyl group of the ester group. In addition, displayed an asymmetric CH 2 stretching peak at 2943 cm −1 and a symmetric CH 2 stretching peak at 2869 cm −1 . Te absorption peak at 1294 cm −1 related it to the C-O and C-C stretching modes, and bands at 1239, 1161, 1107, and 1045 cm −1 attributed to asymmetric and symmetric C-O-C stretching [3,25]. Te intense chitosan peak was identifed at 1726 cm −1 of carboxylate ion, and the chitosan spectra revealed a broad band of about 3245 cm −1 linked with O-H and N-H stretching vibrations.; peaks at 2945 and 2897 cm −1 , corresponding to asymmetrical and symmetrical methylene groups, and an 1180-1063 cm −1 range, characteristic of its saccharide structure, were also detected. Interestingly, the chitosan band associated with the C�O stretching of amide I  International Journal of Biomaterials centered generally at 1615 cm −1 . At the same time, the peak at 1510 cm −1 is attributed to the amideII band. In addition, the characteristic band due to the C-C aromatic stretch occurring at 1420 cm −1 was also present in the sample. Tree peaks are situated between 1020 and 1140 cm −1 related to C-O-C stretching asymmetric and symmetric mode [25,37,38]. (Figures 7(a), 7(b)) for titanite nanocompounds (CaTiO 3 ) and (BaTiO 3 ), respectively, showed comparable vibration beaks with strong indication stretching band in range (410-559 cm −1 ) that attributed to the (Ti-O) [3]. A band also detected at 597 cm −1 corresponds to Ca-Ti-O for CaTiO 3 and the same for BaTiO 3, which belongs to Ba-Ti-O [39,40]. On the other hand, the same weak peak indicated at 1647 cm −1 related to the molecular water content of (O-H) band vibration shows the presence of the hydroxyl group, another weak band of hydroxylate (O-H) also present at 3378 cm −1 for BaTiO 3 and at 3356 cm −1 for CaTiO 3 [41,42]. For BaTiO 3, the peak presented at 1430 cm −1, attributed to BaTiO 3 -OH, while for CaTiO 3, the peak related to CaTiO 3 -OH is located at 1470 cm −1 [42,43].
Te FTIR for both composite coatings (PCL/Chitosan/ Nano-CaTiO 3 , PCL/Chitosan/Nano-BaTiO 3 ) showed a stretching vibration peak in the same range as its component with a slight shift to a low wave number, which confrms the formation of a homogenous and novel composite coat.
3.6. Contact Angle. Te surface of a metallic implant is the primary interface between the implant and the host tissue. Te adsorption of serum proteins and the adhesion behavior of osteoblasts and bacteria may be signifcantly afected by the hydrophilicity of the metal surface. Te contact angle measurements of the coating samples and the control sample, as given in (Table 3), revealed angles ranging from 43.642°to 18.534°, indicating that the surfaces of all samples were hydrophilic.
Control sample Ti-25Zr has low wettability. However, after acid and alkaline treatment that increases surface energy and roughness, the contact angle values became signifcantly lower and changed from hydrophobic to hydrophilic with a contact angle of 43.642°, which is benefcial for cell adhesion and increased biocompatibility [33,44].
Te results further show that the composite coating sample's wettability improved. All the coatings were hydrophilic due to their surfaces' high porosity and roughness, the amino groups present there, and the chitosan's hydroxyl group, which is linked with the hydrogen in water molecules, which reduced their hydrophobicity [45].

International Journal of Biomaterials
As shown in (Table 3), the composite coating containing Nano(CaTiO 3 , BaTiO 3 ) ceramic additive to polymer shows a low angle of 41.402°and 18.534, respectively, which may be attributed to the presence of Nano-ceramics fller particles enhance the hydrophilicity [1,46] and, together with chitosan, improve the wettability.    Figure 7: FTIR spectra of (a) pure (PCL, Chitosan, Nano-CaTiO 3 ) and composite (8% w/v PCL, 2% w/v Chitosan, 1% w/v Nano-CaTiO 3 ) and (b) pure (PCL, Chitosan, Nano-BaTiO 3 ) and composite (8% w/v PCL, 2% w/v Chitosan, 1% w/v Nano-BaTiO 3 ). 8 International Journal of Biomaterials International Journal of Biomaterials 9 3.7. Antibacterial Evaluation. Due to its high biodegradability, nontoxicity, and antibacterial characteristics, chitosan is often used as an antimicrobial agent, alone or in combination with other natural polymers [47,48]. Te addition of the Nano-ceramic fller (CaTiO 3 , BaTiO 3 ) used with chitosan was also reported as an antibacterial material, improving the overall inhibition zone around the coating implant.
Staphylococcus aureus (S. aureus) and Streptococcus mutans (S. mutans) bacteria were used in the antibacterial test. All the composite coatings (PCL/Chitosan/Nano-CaTiO 3 , PCL/Chitosan/Nano-BaTiO 3 ) show good antibacterial efects with comparable results in both kinds of bacteria, and the results are shown in (Figure 8) and (Table 4).
Previous reports suggest using bioactive calcium titanate with incorporated silver ions, while others use iodencontaining calcium titanate to create a highly bioactive surface and simultaneously resist bacterial infection [6,16]. Nano-BaTiO 3 was also found to exhibit antimicrobial activity, which may be attributed to a decrease in ergosterol biosynthesis leading to cell death [49].
Furthermore, another study provides a novel approach for the electrical polarization of piezoelectric and nonpiezoelectric biocompatible ceramics, including (CaTiO 3 and BaTiO 3 ), which have been investigated for the development of antimicrobial implants. On polarized surfaces, the vitality of Staphylococcus aureus (S. aureus) and Escherichia coli (E. coli) bacteria decreases dramatically. Furthermore, the efect of polarization on the antibacterial response has been investigated using a variety of mechanisms, including the formation of reactive oxygen species (ROS), catalase activity, and lipoperoxidation [50].

Cytocompatibility
3.8.1. Cell Growth and Morphology. Te cell viability, proliferation, and diferentiation at the implant-host tissue interface have a signifcant role during implantation. Titanium and its alloys have good biocompatibility and the ability to support cell proliferation. However, because of the inert surface and the potential for microbial infection, these biometallic alloys need to activate their surfaces to promote interaction with cells and resist bacteria. Te cytocompatibility of coated and noncoated Ti-25Zr alloys was studied using the MC3T3-E1 cell line and evaluated quantitatively by the Alamar Blue assay, as shown in Figures 9(a) and 9(b).
( Figure 9(a)) show the MC3T3-E1 cell viability percentage for bare Ti-25Zr alloy and the other of the same alloy with diferently coated surfaces by Nano-composite fber PCL/Chitosan/Nano-CaTiO 3 (CA1) and PCL/Chitosan/ Nano-BaTiO 3 (BA1). As can be seen, the cell viability percentages of all the samples rise over time, rising on day 3 and reaching an even higher value on day 7. At day 7, a higher level of cell viability was observed in cells on the coated sample than in the noncoated ones (CA1 vs.Ti-25Zr n � 3, p < 0.0001; BA1 vs. Ti-25Zr n � 3, p < 0.00017). (CA1) coated sample displayed higher cell viability even more than (BA1) (CA1 vs. BA1 n � 3, p < 0.00034). Tis behavior may be attributable to the time taken to mineralize the coated surfaces. After 7 days of culture, the sample surface-coated (CA1) demonstrated a signifcantly higher value of cell viability, reaching 271.56 ± 13.15%, followed by (BA1), which has 181.63 ± 17.87%, while the noncoated bare Ti-25Zr demonstrated lower cell viability in comparison to the coated sample, with 80.52 ± 1.97%. (Figure 9(b)) show MC3T3-E1 preosteoblast cells proliferation was assessed by determining the cell number in the scafolds at day 1, 3, and 7, using AlamarBlue fuorescent assay. Te result shows a steady increase in cellular proliferation until the higher value is reached at day 7. Te (CA1) coated sample showed the highest cellular proliferation, reaching a 237.53% enhancement rate compared to the noncoated sample, followed by (BA1), which has a 125.16% enhancement rate.
Tis result indicates that a composite coating containing Nano-CaTiO 3 is more conducive to cell viability and proliferation. Tat may be due to the fact that CaTiO 3 can emit the Ca 2+ ion that promotes positive reactions with cells. Calcium titanate was also reported to polarize the coated surface, enhancing cell activity [9,50].
Incorporating BaTiO 3 Nano-particles into a PCL/chitosan blended polymer enhanced the osteogenic diferentiation of MC3T3-E1 cells. Te piezoelectric efect of BaTiO 3 induced in the Nano-composite fber produced by the electrospinning method greatly improved the proliferation, viability, and extracellular matrix deposition of osteoblastlike cells. Tis observation agrees with the results reported previously, suggesting that the polymer's viscous and elastic properties play an essential role in the piezoelectric performance of piezoelectric polymer composites. With a 91.2% deacetylation degree, chitosan has piezoelectric properties and acts with barium titanite to promote cell activity [3,23].
Te ALP assay is a marker for time-dependent early cell diferentiation. Te quantitative analysis of ALP activity is shown in Figure 10. All samples' ALP activity increased with time. After 3 days, there was no signifcant diference in ALP activity among coated and noncoated samples, which may be attributed to the time needed for the coating layer to mineralize and react with the physiological host medium. On day 7, a higher level of ALP activity was noticed in cells on the coated sample than the noncoated one (CA1 vs. Ti-25Zr n � 3, p < 0.00022; BA1 vs.Ti-25Zr n � 3, p < 0.0065). (CA1) coated sample displayed higher ALP activity even more than (BA1) (CA1 vs. BA1 n � 3, p < 0.0114).
ALP is an osteogenic diferentiation marker and plays a key role in reparative bone mineralization. It is expected that ALP for osteoprogenitors would be higher for surfaces with a superior biologic response. We observed an increase in ALP activity on all the surfaces over time, and (CA1) especially had the most signifcant infuence on ALP activity on day 7. Since we observed diferences in the osteoinductive properties of these surfaces, we believe that diferences in surface composition and microstructure with the incorporation of a Nano-active ceramic compound in the composite coating layer play a role in the implant's osteoinductive ability, with the (CA1) alloy showing the greatest efect.
Perovskite calcium titanate CaTiO 3 (CTO) has been used recently as the main coating component on titanium implants because of its biocompatibility, middle thermal expansion coefcient between Ti and HA, enhancement of apatite bonding, as well as its promotion of cell attachment and proliferation. CTO has demonstrated potential applications for bone regeneration due to the fact that it provides the opposite surface charges with phosphate ions in a simulated body fuid (SBF), in which it infuences to grow further bone-like apatite CaTiO 3 coatings is an efective method to enhance the biocompatibility of titanium alloy [27]. Several studies have shown that introducing calcium ions to the titanium surfaces can convert passive oxide to active oxide, resulting in a CaTiO 3 coating and enhancing titanium's biological activity. CaTiO 3 is also reported to induce adhesion, proliferation, and bone-like apatite deposition in osteoblasts [14,15]. Incorporating BaTiO 3 particles into the polymeric matrix signifcantly increased dielectric permittivity and decreased dielectric loss. Te bioactive surface of these compounds promoted osteoblast cells' adhesion and proliferation, with distinctive ALP activity and deposition of osteocalcin and collagen I [3]. Surface composition alone can also infuence cell behavior, e.g. diferentiation of cells. Te addition of Zr as an alloying element to Ti highlights the role of surface composition and grain refnement in the behavior of osteoblast cells, especially at the diferentiation stage of cell-material interactions [33].

Conclusion
(i) Te result shows that the Ti-25Zr alloy has a completed solid solution with α phase, which has better mechanical properties than pure Ti, with the exception that modules of elasticity were reduced by 29% because of the presence of Zr and porosity in the alloy. Tis characteristic is a favorite in biomedical applications because it reduces stress shielding. (ii) Te electrospinning coating method produces Nanofber with nanoroughness, high surface contact area, and porosity, which, in consequence, improves the positive reaction with the physiological medium and promotes cell adhesion, proliferation, and diferentiation. (iii) Te water contact angle shows that all samples are hydrophilic, and the contact angle is lower for the coated surface than for the uncoated one.
(iv) Te antibacterial test shows no inhibition zone for the control sample (bare Ti-25Zr) alloy, while the coated samples show a reasonable and comparable inhibition zone. (v) Both (PCL/Chitosan/Nano-CaTiO 3 and PCL/Chitosan/Nano-BaTiO 3 ) coating flms signifcantly increase cells viability, proliferation, and ALP activity of MC3T3-E1 cells on the coated surface of Ti-25Zr alloy, confrming that the electrospinning coating surface dramatically improves the cytocompatibility of the biometallic alloy.

Data Availability
Te data that support the fndings of this study are available from the corresponding author upon request.

Conflicts of Interest
Te authors declare that they have no conficts of interest.