1. Introduction
Transparent materials such as glass have extensive applications in biomedical science, and since they react with physiological fluids and form tenacious bonds to hard and soft tissues through cellular activity, they have to be compatible with biologic organs and surrounding living tissues [
1,
2,
3,
4,
5]. Surface modification is one of the methods of changing surface properties such as roughness, absorption, chemistry and the general structure of materials to enhance the biocompatibility without altering material bulk properties [
6,
7,
8,
9]. Methods used to modify surface properties include chemical, physical and chemophysical treatments like electrochemical etching or photolithography, and these are costly and toxic [
6,
7,
8,
9,
10]. Another modification process reported previously is to apply an enhanced thin layer on the substrate with a laser without changing the bulk properties [
8,
11,
12]. The primary objective of this research is to compare different parameters of introduced ultrashort pulsed laser processing to enhance the biocompatibility of glass for biomedical use. In this study, nanofibrous titanium (NFTi) on glass substrate was prepared by the high intensity laser induced reverse transfer (HILIRT) method and compared with theoretical results of the process. In this method, high laser energy passes through the transparent glass slides, which is then focused to a very small point on the titanium sheet (target material) in a picosecond range with different laser intensities and pulse intervals. Removing titanium particles from a titanium sheet and accumulating them as nanofibers on the glass surface is the result of the energy input from the laser being absorbed and causing the solid’s binding energy to be broken [
11,
12,
13]. This classic method increases atomic kinetic energy and temperature and finally causes phase transitions. Unlike other heating methods, the focused laser beam primarily impacts a material’s electrons in a short time scale [
13,
14]. When the laser is centralized on a surface, it causes material ablation and removes atoms from the bulk surface, thus enabling these atoms to be emitted and form plasma plume, consequently generating nanofibrous structures by rapid cooling [
14]. In this study we investigated the effect of laser power and laser frequency on deposited nanofibrous titanium (NFTi) thin film and biocompatibility on glass substrates via the HILIRT method. In order to analyze the deposited titanium and enhanced-biocompatibility bioactive glass (EBBG) structures and morphologies, a scanning electron microscope (SEM) was used. X-ray diffraction (XRD), and Raman spectroscopy were utilized for structural identification of deposited titanium coatings and hydroxyapatite-like layers formed on EBBG. To assess biocompatibility, an MTT metabolic activity assay was used to determine human cell viability on these substrates.
3. Results and Discussion
As illustrated in
Figure 1, by increasing laser power from 5 to 12 W, the amount of NFTi increased. It is also clear that a laser power of 5 W is not enough laser intensity for removing particles and ablating materials, since the surface of the microscope glass that went through laser deposition with a laser power of 5 W was similar to bare glass, with only a small amount of NFTi. This means that when a laser power of 5 W was focused over a glass surface, it raised the temperature, but not high enough to produce sufficient material ablation and remove atoms and particles from the bulk surface; therefore, the titanium atoms could not be emitted from the titanium sheet and small amount of nanofibers could be formed via rapid cooling process (
Figure 1A). Also, since by increasing laser power the shape of deposited materials would change from agglomerated to nanostructure, their diameter decreased, which is shown in
Figure 1D. As shown in
Figure 1 and as reported before [
9,
12], the surface roughness in microscale has not been changed significantly.
As per Equation (1), the radial temperature gradient has a direct relation to the
Imax which is directly dependent on the average laser power. This means that by increasing the laser power, the temperature gradient goes up and a larger HAZ forms, as shown in
Figure 2.
As displayed in
Figure 3, by increasing the laser power and decreasing the scanning speed, the ablation depth increases. To further clarify, incrementing laser power while other laser parameters are constant causes the amount of the energy delivered to the Ti substrate to increase. This results in the generation of a larger HAZ and greater ablation volume on the Ti surface, which is in agreement with the results presented in
Figure 2. The laser scanning speed has a similar behavior, as at lower scanning speed (higher number of pulses) more energy would be transferred to the substrate, which results in a deeper ablation profile and thicker nanofibrous layer.
The effect of laser frequency is shown in
Figure 4. By increasing laser frequency from 600 to 1200 kHz, the amount of NFTi also increased. This can be due to the fact that increasing laser frequency means decreasing pulse intervals, which leads to a shorter time between consecutive pulses and more transmitted intensity into the substrate, and more heat accumulation and higher average temperature [
15,
16,
17,
18,
19]. Enhancing the spot temperature causes a denser plume and generates more ablated atoms, thus leading to more consecutive inelastic collisions, which results in the deposition of a titanium nanostructure on glass substrates with reduced diameter [
15].
In the HILIRT method, the effect of pulse repetition needs to be considered along with the number of delivered laser pulses. Generally, at a higher pulse repetition rate and a constant amount of average laser power, the energy of each pulse is less, which leads to the lower maximum temperature for the HAZ, as demonstrated in
Figure 5. However, for multi-pulse processes, the average surface temperature increases by increasing the pulse repetition [
15,
16,
17,
18]. Increasing the laser repetition rate results in decreased pulse intervals (shorter time between consecutive pulses) and a higher number of laser pulses per unit time (delivering more laser intensity into the substrate) [
15,
16,
17,
18].
As stated in the above paragraph, at higher laser repetition rates and shorter pulse intervals, denser laser plumes would be formed as the average surface temperature is higher [
15,
18]; this causes the generation of smaller laser ablation volume (and more fibers) as shown in
Figure 6.
The ability to form the hydroxyapatite layer with a calcium to phosphorous ratio of 1.63 and other compositions of these two elements is known as biocompatibility [
4]. In order to analyze the specimens’ biocompatibility, samples should be immersed in SBF. All the specimens produced with different laser powers and frequencies were immersed in SBF for 4 days, and the SEM results show the hydroxyapatite-like layer, which is a good indication of the samples’ biocompatibility. Additionally, increasing laser power and frequency creates better and more consistent layers of hydroxyapatite which can be the result of more NFTi on the samples. To illustrate further, samples with more NFTi have greater area to volume ratio of suitable places for calcium and phosphorous to nucleate and grow on, which means that they have more biocompatible surfaces, as indicated in
Figure 7 and
Figure 8.
The XRD and Raman results in
Figure 9 show that titanium and titanium oxide phases reflected severe peaks on samples created by laser powers of 5, 9 and 12 W. Generally, increasing laser power results in injecting pulses with higher intensity on the center of the ablation, which leads to a rise in the plume temperature and consequently more particle ablation and fiber generation.
In samples that were immersed in SBF solution for 4 days according to the previous reported results for similar structures [
9,
12] there are primary peaks related to hydroxyapatite-like composition, which can be the result of a good amount of NFTi on the samples and proper places for calcium and phosphorous to grow on, as shown in
Figure 10. The hydroxyapatite and other calcium to phosphorus composition peaks relating to Raman and XRD patterns of samples which were produced by higher powers have higher intensities compared to the samples produced by a laser power of 5 W. This can also be due to little or no generation of NFTi on the specimens produced by a laser power of 5 W, and less suitable places where calcium and phosphorous elements can start nucleating and growing [
20,
21].
The XRD and Raman results in
Figure 11 show that titanium and titanium oxide phases reflected severe peaks on samples created by laser frequencies of 600, 800 and 1200 kHz. Generally, increasing laser frequency means a decrease in the laser pulse interval, which brings about a rise in plume temperature and consequently more particle ablations and more fiber generation.
Also, in samples immersed in SBF solution there are several peaks related to hydroxyapatite-like layers which are the result of a good amount of NFTi on the samples and good places for calcium and phosphorous to develop. As illustrated in
Figure 12, the hydroxyapatite and other calcium to phosphorus composition peaks relating to Raman and XRD patterns of samples produced by higher laser frequencies have higher intensities compared to the sample produced by a laser frequency of 600 kHz. This can be due to little or insufficient generation of NFTi on the sample produced by the laser frequency of 600 kHz, and less desirable places for calcium and phosphorous elements to start growing. Also, similar to laser power, the sample produced with the highest laser frequency has a drastic upper shift in the Raman spectrum which can be due to the thicker NFTi coating generation.
Indirect MTT assay of the produced samples is shown in
Figure 13. Generally, samples with different NFTi coatings did not show high toxicity after three days. Cell viability was constant at 100% over the period due to no toxic ion release into the test solution and no significant effect on its life properties. Although there is no difference between the cell viability of the samples produced with different laser frequencies (almost 100% viability), the solution extracted from the sample produced with a laser power of 5 W showed insignificant toxicity after immersion for 2 and 4 days in the medium. This can be representative of insufficient or less NFTi generation on the glass and no protective or compatible titanium and hydroxyapatite-like layers on top of the glass substrate to prevent toxic ion release, such as silicon ions, from the glass substrate [
20,
21].
SEM images of cell morphology and cell adhesion on samples with different laser parameters are shown in
Figure 14. Attached cells on samples produced with a laser power of 5 W and a laser frequency of 600 kHz can be seen in
Figure 14, in comparison with samples synthesized with higher laser powers and higher laser frequencies. The SEM results illustrate that cells adhered better on NFTi surface than bare glass and they are attached better and seem healthier at higher power and higher frequency samples. This can show the biocompatibility of NFTi on the coating produced with higher frequencies and the preference of cells to be fitted on an NFTi coating rather than bare glass. Additionally, cell filopodia have more attachments on the NFTi layers, which is known as their extra cellular matrix (ECM), with increasing laser power and frequency. Cell bodies show less degradation and disappearance during the fixing process, which means that increasing laser power and frequency brings about more NFTi generation and thicker NFTi layers on glass substrate, with higher surface to volume ratio and more compatible and available places as ECM for cell filopodia to connect and adhere to [
21].