# Orthopedics-Related Applications of Ultrafast Laser and Its Recent Advances
## 2 Background  
Laser ablation of bone tissue has gradually been introduced and accepted clinically as complementary techniques due to its higher precision and non-contact nature, although mechanical methods such as the use of drills, burrs, saws and milling cutters remain the gold standard for bone cutting procedures [16; 17]. In particular, compared to laser ablation, mechanical instruments during bone cutting generate substantial thermal damage, bone fragmentation, amorphous formation and a mineral-rich carbon layer as well as biomechanical stress to adjacent structures; yet surgical protocols are well-established for these tools with acceptable clinical outcomes [1]. Among the many types of clinical laser systems, several common ones found within the field of orthopedics are Er:YAG, Nd:YAG, Argon and CO\({}_{2}\) lasers, which machine bone structures by the means of vaporization and melt expulsion [1; 18; 19] (Table 1). These conventional clinical laser systems, which have been developed often as continuous-wave or with longer wave pulse (in the \(\upmu\)s and sub-\(\upmu\)s range), depend primarily on the linear absorption of light by the target chromophore(s); thus, the penetration depth is significantly affected by the absorption coefficient of water in an inverse relationship [1]. Due to the strong wavelength-dependent nature, non-deterministic ablation can be generated in highly heterogenous tissue structures leading to unnecessary injuries. Collateral damage surrounding the laser focal volume can also be induced by excessive heat release, thus greatly hampering bone healing and regeneration [1; 16; 17; 20]. The potential and efficiency of ultrafast pulsed laser (pulse ranging approximately from 100 fs to 10 ps) for ablating biological hard tissues has been extensively studied to overcome the thermal and thermo-mechanical drawbacks of the conventional laser systems. In particular, an ultrafast laser source is capable of producing extremely high peak radiation intensities (>10\({}^{11}\) W/cm\({}^{2}\)), leading to the initial generation of free electrons through a multiphoton absorption mechanism and band-gap (Zener) tunnelling. A high density of excited electrons is therefore created locally in a small focal volume due to nonlinear absorption, triggering avalanche ionization that supports plasma formation. Ablation of tissue is then generated by the rapidly expanding high-density plasma, forming smoother cavity walls with negligible thermal damage. This process is known as plasma-mediated ablation or laser-induced breakdown. A free electron formed during the multiphoton absorption and avalanche ionization processes can be promoted from the ground state to the valence band by the energy of two, three, four, five and six photons for wavelengths longer than 191, 383, 574, 766 and 958 nm, respectively. Wavelength dependence of the process thus indicates decreased tissue damage threshold along with decreasing wavelength [25]. Thermal damage can remain highly confined in the focal volume during optical breakdown not only because the location of plasma formation can be controlled by precise energy deposition using focused laser radiation but also because ultrafast laser pulses are a few orders of magnitude shorter than thermal diffusion time (heat transfer ranging in the order of picoseconds to nanoseconds from absorbing proteins to tissue water) [26; 27; 28]. A standard state-of-the-art ultrafast laser system normally uses a Ti:sapphire oscillator in the near-infrared region or infrared region (NIR/IR) for tissue ablation, where the pulse width and repetition rate can vary depending on the application. Interested readers can find more details on ultrafast laser interactions and mechanisms with biological hard tissues from literature reviews [20; 29; 30; 31; 32; 33; 34]. Advantages of using ultrafast laser over longer wave pulse laser have also been recognized, specifically including absence of melting,