**5. Laser interaction mechanisms**

On the implication of laser light into the transparent media, such as biological tissues, as well as viscoelastic like spider silk, various interaction mechanisms were facilitated. Specifically, the tissue characteristics along with the laser parameters are contributing to this diversity. Most important amongst them are optical properties of tissues that include the coefficient of reflection, absorption, and scattering. Altogether, these parameters determine the total transmission of the tissues and viscoelastic at a certain wavelength.

#### **Figure 5.**

*Z-scan data from a single fiber (150 GW/cm<sup>2</sup> ) along with equations with 2, 3, 4-photon. The mismatch near edges was attributed to diffraction losses from the edges of the silk filament. Data taken from Sidhu et al. [20].*

In contrast, the following parameters that affect the laser ablation, are given by the laser radiations themselves: applied energy, wavelength, exposure time, focal spot size, and energy density. Amongst them, exposure time is a very crucial parameter when selecting a certain type of interaction.

Laser-tissue interactions can be subdivided into five different regimes depending on the laser power and the exposure time. **Figure 6** gives us a rough delineation of the interaction regimes. Two lines represented diagonally show the fluences (energy per unit area) at 1 J/cm<sup>2</sup> and 1 kJ/cm<sup>2</sup> , respectively.

According to this plot, the time scale can be roughly divided into five parts: continuous wave or exposure times >1 s indicates photochemical interactions, 1 min

#### **Figure 6.**

*A double logarithmic map of five basic laser-tissue interactions. The circle gives only a rough estimate of the associated laser parameters [18].*

down to 1μs indicates thermal interactions, 1 μs down to 1 ns indicates photo-ablation, and < 1 ns indicates plasma-induced ablation, as well as photodisruption. The difference between the latter two is ascribed to different energy densities [18]. Precise laser surgery can be achieved when the desired target tissue, such as retina or ocular-tissue, is exposed to the optimal laser conditions following the subsequent interaction pathway.

### **5.1 Optical absorption and light propagation**

Optical absorption takes place when the frequency of incident laser irradiation matches with an electronic excitation frequency of a molecule [18]. The typical absorption coefficient estimated by inverse Monte Carlo analysis for bovine cornea and retina at 800 nm was �0.001 cm�<sup>1</sup> and 2.5 cm�<sup>1</sup> , respectively [29].

The coefficients for 2, 3, 4, and 5 photon absorptions for spider silk were <sup>α</sup><sup>2</sup> = 1 � <sup>10</sup>�<sup>2</sup> cm/GW, <sup>α</sup><sup>3</sup> = 2 � <sup>10</sup>�<sup>5</sup> cm<sup>3</sup> /GW<sup>2</sup> , <sup>α</sup><sup>4</sup> = 4 � <sup>10</sup>�<sup>6</sup> cm5/GW3 , and <sup>α</sup><sup>5</sup> = 5 � <sup>10</sup>�<sup>7</sup> cm7 /GW<sup>4</sup> , respectively (**Figure 7**) [20, 30]. Previously, a 3-photon absorption process was observed in the silk-fibroin solution [31], and enhanced 3- and 4-photon absorption was reported for amyloid fibril solution [30].

In the case of biological tissues, such as retina and cornea-tissue, neither water nor macromolecules are present, it absorbs the light strongly in the near-infrared range (roughly between 600 to 1200 nm), this spectral range is considered a "therapeutic window." Laser radiation in this window can penetrate deeper into biological tissues with lower loss during treatment.

Light propagation inside the retina or corneal tissues can be attributed to their absorption and scattering properties. The effective attenuation coefficient (*Aeff*) quantifies how deep the laser light can penetrate, and is expressed as:

$$A\_{\rm eff} = \frac{1}{\sqrt{3a\_t(a\_t + (1 - g))}}\tag{4}$$

**Figure 7.**

*Nonlinear absorption coefficient of silk fiber versus input peak intensity. Theoretical fits for pure 2-, 3-,4- photon absorption are compared with a mixed fit up to 4th order polynomial [20].*

**Figure 8.** *Laser penetration depth in tissue at different wavelengths [21, 27].*

where *α<sup>a</sup>* and *α<sup>s</sup>* are absorption and scattering constants. *g* is the coefficient of anisotropy which is a tissue-dependent parameter. With g = 1 denotes perfectly forward scattering, such as cornea, the effective penetration depth (*Leff*) of incident light is.

$$L\_{\rm eff} = \frac{1}{\alpha\_{\rm eff}}\tag{5}$$

**Figure 8** Illustrates the approximate penetration depth in tissues of different wavelengths, while both*,α<sup>a</sup>* and *α<sup>s</sup>* are taken into account.

The selection of appropriate wavelength and pulse energy followed by linear or nonlinear propagation of laser irradiations into the matched tissue target allows unprecedented precision to achieve surgical effects.

### **5.2 Photochemical interactions of fs-lasers**

It is a well-known fact that light can induce chemical effects or reactions within biological tissues. In ophthalmology, photochemical interactions can play a vital role in dye-assisted photodynamic therapies (PDT). During PDT, photosensitizers were injected into the target tissues. Laser irradiations of specific wavelength trigger the photochemical reaction at low power densities (�1 W/cm<sup>2</sup> ) and long exposure time in order of seconds to continuous wave (CW) [18, 32]. It results in toxic reaction products that cause irreversible destruction of target cell structures. The reaction could be the generation of reactive oxygen species from the interaction of light, oxygen, and photosensitizers, such as verteporfin, and indocyanine green (ICG). These were commonly used to treat choroidal or corneal neovascularization, that is, abnormal blood vessels underlying retina cell layers or sprouting in the corneal stroma, respectively. However, severe adverse reactions are associated with these photosensitizers, including visual disturbance, stromal haze, and lipid degeneration injection site reactions even after months of treatment [33]. Several joules of energy (�150 J/cm<sup>2</sup> ) continuously deployed into corneal stroma allows long-term reduction of neovascular structure, but can also cause corneal scarring or thermal injuries [34, 35].

We suppose that without the use of chemical agents, the intensity of femtosecond laser pulses could be used to manipulate abnormal blood vessels lower than these conventional photodynamic treatment procedures. We will discuss the selective removal of corneal and retinal blood vessels in the next section of this chapter. The precise cutting and manipulation of ultrafine fibers are also discussed.

### **5.3 Thermal interactions**

Thermal effects can be induced by both CW and pulsed-laser irradiations. The rise in local temperature could be attributed to the absorption of photon energy by protein, pigment, or bound water molecules. On the basis of pulse duration and peak values, different thermal effects can be achieved in tissues, including coagulation, vaporization, carbonization, and melting. Moreover, with the rise in temperature to 50°C, the enzyme activity in tissue cells is reduced, energy transfer slows down, and cell repair mechanisms are disabled. At about 60°C, protein denaturation occurs, leading to more immediate cell necrosis and tissue coagulation. Therefore, photocoagulation is the general procedure to treat retinal disease, such as proliferative diabetic retinopathy. Meanwhile, traditional lasers employed for photocoagulation (spot size: 0.1–3 mm, pulse duration: millisecond) creates a heat wave that spreads beyond the focal volume causing inevitable collateral damage to underlying cell layers [36]. This leads to scarring in the retinal segments. Thermal scarring enlarges progressively up to 300% and can cause significant vision loss if the fovea region is involved [37]. Thus, it is important to minimize the thermal effect during laser irradiations.

If pulse laser irradiations do not undergo photochemical or phase-transition processes, the linearly absorbed energy by the target tissue is entirely converted to the temperature rise. Since the focus of the current work is on femtosecond laser interaction within retinal or corneal tissues, we consider a situation where excessive laser energy, sufficiently higher than the threshold, is applied and absorbed in a small focal volume under single pulse configurations. A single shot of a fs-laser could be the entire source of heat. The temperature rise could be determined by calculating the volumetric energy density gained by the plasma during the laser pulse irradiations [38]. Under adiabatic conditions, the local temperature rise (ΔT(r)) at an arbitrary location (x), is directly related to the local volumetric energy density ε (x), as.

$$
\Delta \mathbf{T}(\mathbf{x}) = \frac{e \; (\mathbf{x})}{\rho \mathbf{C}\_{\mathbf{v}}} \tag{6}
$$

where *ρ* is the tissue density and *Cv* is the specific heat capacity per unit volume [19]. In case of the absence of photochemical or phase-transition, the absorbed laser energy is subjected to undergo spatial redistribution by thermal diffusion leads to collateral damage in adjacent focal volume [39]. The heated volume is a layer of tissue, where penetration length is inversely proportional to an absorption coefficient (*1/μa*) and the thermal diffusion time, (*td*), is expressed as:

$$t\_d = \frac{1}{\kappa \mu\_a^2} \tag{7}$$

where κ is the thermal diffusivity. In order to achieve thermal confinement, the ratio of pulse duration (*t*p) to thermal diffusion time (*td*) should be less than or equivalent to 1 (*tp*/*td* ≤ 1). **Figure 9** shows the normalized temperature profiles in

### **Figure 9.**

*Normalized temperature profiles in water immediately after the laser irradiation with constant irradiance and optical penetration depth (1 μm) at various pulse durations [19, 32].*

water followed by laser irradiations with fixed radiant exposure and optical penetration depth (1 μm) for various pulse durations. At typical *tp* < 3 μs, the temperature profile is confined to the diffraction-limited volume. However, for *tp* ≥ 10 μs, the thermal diffusion has spatially redistributed the energy over the larger volume. The peak temperature also reduced significantly within the sample over time [19, 32]. Thus, pulse durations play a vital role in thermal confinement to achieve localized ablations in the target tissue.

Meanwhile, collagen fibrils and water in corneal stroma often the main chromophores contribute to the absorption of IR and UV irradiations. Collagen fibrils in corneal stroma possess microscopic tissue structures in range of 30 nm in diameter, with corresponding thermal confinement times of 6.3 ns [40]. For micro-scale thermal confinement laser pulses of picoseconds or femtosecond, duration should be employed.

### **5.4 Plasma-induced ablation and photodisruption**

Laser-induced optical breakdown occurs in biological tissues when the applied laser intensity exceeds 10<sup>11</sup> W/cm<sup>2</sup> . In case, an intense laser pulse can excite a large number of electrons and generate plasma, which causes vaporization of the materials [10–14, 18]. When a highly intense ultrashort laser pulse (<10 ps) was focused into the tissue, the associated energy density is high enough to induce nonlinear absorption of laser energy through multi-photon, tunneling, and avalanche ionization [10, 16, 19, 20]. Thus, producing micrometer-sized highly excited plasma in the vicinity of the focal volume. Moreover, within the stipulated pulse duration time, the temperature of the laser-induced plasma can reach several thousand Kelvin.

During the process of optical breakdown, plasma generation allows energy deposition in pigmented, as well as weakly absorbing tissues, such as cornea, lens, or retina. Laser-induced plasma will serve as an absorber of photon energy which leads to an increment in absorption coefficient [18, 32]. Thus, by means of plasma-induced ablations, clean removal of tissue without evidence of thermal or mechanical damage can be achieved even in transparent tissues.

*Ablation of Materials Using Femtosecond Lasers and Electron Beams DOI: http://dx.doi.org/10.5772/intechopen.106198*

**Figure 10.**

Alongside plasma generation, the movement of energetic free-electrons in plasma outward from the focal volume results in secondary effects, such as shockwave generation [18, 19, 32]. The shockwave leaves the plasma boundary at supersonic velocity and then slows down to the speed of sound, refers to as acoustic wave generation. The processes associated with an optical breakdown on the interaction of fs–laser pulse with materials were shown in **Figure 10**. Lastly, if energy density still lasts at the focal volume, it forms a cavitation bubble, which performs several oscillations of expansions and collapses within a period of few hundred microseconds under external pressure [41]. **Figure 11** illustrates the scheme of physical processes associated with optical breakdown. Plasma formation followed by mechanical effects could be referred as "*photodisruption*." It could be distinguished from plasma-induced ablations on the basis of employed laser energy density.

In plasma-induced ablations, the plasma formation vaporizes tissue at focal volume with lower threshold energies, providing a localized surgical effect. While, at higher energy densities, photo-disruption results in shock wave expansion, the expansion of the cavitation bubble, as well as heat diffusion after thermal equilibration, Thus causing all unwanted collateral damage to the surrounding tissue and limiting the surgical precision [10, 18, 19, 42, 43]. Thus, applied energy density, which is defined as energy per unit area, plays a crucial role to limit the collateral damage. The fluence (J/cm<sup>2</sup> ) employed in the sample must be minimized while still maintaining a sufficiently higher intensity to produce photodisruption through plasma formation.

Recent studies on plasma formation in water revealed some vital trends for pulse durations dependence while inducing optical breakdown. First, on reduction of pulse duration from 100 ns to 100 fs, the irradiance threshold for breakdown increases by 1000 folds, but the radiant exposure threshold decreases from <sup>10</sup><sup>3</sup> to 1 J/cm<sup>2</sup> [42, 43]. For nanosecond pulses, seed electrons for plasma formation

*Schematic of the physical effects associated with optical breakdown [32].*

**Figure 11.** *Plasma-induced ablation and photodisruption distinction on the basis of energy density in cornea tissue [32].*

were generated through thermionic emission which requires higher peak intensity. Rather, higher intensity of laser field generated by picosecond and femtosecond laser pulses can cause multi-photon ionization, which supplies the seed-free electron needed to start plasma generation at much lower peak intensities. Thus, the difference in radiant exposure threshold comes from the mode of free-electron generation or the initiation of plasma generation. As a result, nanosecond pulse lasers need higher threshold energy for plasma generation as compared to femtosecond and picosecond lasers. Second, the plasma transmission is small for ns-pulses, increases considerably for picoseconds pulses with maxima around 3 ps, and decreases again for fs-pulses [43, 44]. Lastly, the plasma energy density is more than one-fold smaller with fs-pulses than with ns-pulses [42–44]. Considering the facts, fs lasers could be employed as a microsurgical tool for precise ablations of biological tissues.
