Laser is an acronym for light amplification by stimulated emission of radiation.

The medical field, quick to realize the possibilities in laser energy, was quick to adapt lasers for medical use. Only 3 years after the first commercial laser was built, McGuff et al. reported on the effect of a ruby laser on melanoma cells transplanted in hamsters [23]. Many more studies examining the utilization of laser energy for treating cancer soon followed [24–28].

The process of creating a laser beam involves exciting a particular material (CO₂, Nd:YAG, etc.) by an external source (light, electricity, chemical reaction). The material is placed between two mirrors which act as an optical resonator which intensifies the interaction between the electromagnetic field and the excited material (amplification). Making one of the mirrors partly transparent allows the resulting laser beam to exit the resonator.

The characteristics that make laser energy well adjusted to medical implications are the following: it is coherent (the wave trains are exactly in phase), collimated (the beam is parallel), and monochromatic (all the photons have the same wavelength, frequency, and energy).

The physical properties of a laser system are dictated by its wavelength, and the wavelength of the laser is dictated by the source that is being excited, for example, the CO₂ laser produces a beam of infrared light with the principal wavelength bands centering around 9.4 and 10.6 μm; the main absorption of biological molecules occurs within the range of wavelength shorter than about 280 nm (ultraviolet). The penetration of light is optimal at wavelengths longer than 1 μm (the near-infrared range of the spectrum). The high water content (60–80%) of most tissue leads to an extensive absorption of infrared radiation and thus to a very efficient energy transfer and heating of the tissue when irradiated with lasers of these wavelengths.

The output of a laser may be a continuous constant-amplitude output (continuous wave) or pulsed. In the pulsed mode of operation, the output of a laser varies with respect to time, typically taking the form of alternating “on” and “off” periods. This application facilitates the depositing of as much energy as possible at a given place in as short time as possible.

The term laser ablation refers to the thermal destruction of tissue by laser. The term “interstitial” laser ablation reflects the fact that the laser fiber is inserted into the tissue as opposed to ablating tissue with a laser while maintaining a buffer medium between the fiber emitting the energy and the tissue being ablated (e.g., air for cutaneous application, saline when applying laser energy in the bladder/ureter).

Interstitial laser ablation was first described by Bown in 1983; he inserted a 400-μm glass fiber into a metastatic skin lesion with the aid of an Nd:YAG laser system and caused local necrosis in the treated area [29].

The basic principles behind laser ablation are the conversion of laser light (excited photons) into heat by tissue. The optical and thermal properties of the tissue as well as the parameters of laser beam influence the extent of the thermal ablation. The optical and thermal properties of the tissue are determined by their structure, water content, and blood circulation. The key concepts are absorption, scattering, reflection, thermal conductivity, and heat capacity.

The prostate as a tissue is suited for FLA due its optical absorption rate without excess vascularity. Highly vascularized tissues such as liver have ample heat conduction therefore limiting the ablation size when performing FLA.

The laser most commonly used for FLA is the Nd:YAG laser, with a wavelength of 1,064 nm, but it is being replaced by more compact and less expensive infrared (800–980 nm) diode lasers. The delivered photons induce an increase in temperature, increasing temperatures to 45 °C and 50 °C, and induce enzymatic changes at the cellular level and an edema develops; temperatures above 60 °C cause rapid coagulative necrosis and instant cell death, but irreversible cell death can also be achieved at lower hyperthermic temperatures (>42 °C), although longer durations are required [30, 31]. Temperatures above 100 °C will cause vaporization of cellular protoplasm, followed by desiccation and shrinkage of the tissue; afterward, any additional laser energy causes a quick temperature rise, and temperatures above 300 °C cause the tissue to burn and carbonization occurs. Carbonization of tissue tends to occur closer to the laser fiber and decreases the optical penetration and heat conduction of the tissue being ablated. It limits the size of lesion produced and due to the irregular and unpredictable light absorption in the tissue causes unpredictable ablation volume; therefore, during FLA, it is advisable to refrain from heating above 100 °C.

Many new modifications have been made to allow for efficient tissue ablation. The introduction of the flexible laser fiber, utilizing quartz fibers of small diameter (250–1,000 μm), allows the application of FLA through flexible fiber-optic devices and through thin needles [32, 33].

Interstitial fibers, which are quartz fibers that have flat or cylindrical diffusing tips and are 10–40 mm long, provide a much larger ablative area of up to 50 mm [34, 35].