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Laser Tissue Interaction

Fundamentals of Laser-Tissue-Interaction

When photons are irradiated on tissue they can be transmitted, reflected or backscattered. During this type of interaction they do not transfer any energy to the tissue. This occurs for example, when visible light passes through the transparent parts of the eye. If the photon energy however is resonant with any energy levels of the tissue, absorption takes place and energy is transferred to the tissue. This is the fundamental process of any therapeutic laser application.

Laser-Tissue Interaction Regimes

Therapeutic laser-tissue interaction processes can be divided into five different regimes, determined primarily by the intensity of the laser irradiation and its interaction time with the tissue (Fig. 1).

Photochemical Reactions

At moderate irradiances and exposures roughly longer than seconds (laser-) light, mainly in the visible or ultraviolet (UV) region, produces photochemical reactions.

Photodynamic Therapy (PDT) is used as a minimally invasive method to treat age-related macular degeneration and malignant cancers. A photosensitizer or photosensitizing agent is used which accumulates at the target tissue. During excitation with light (typically at 600 – 800 nm) the sensitizers transfer their energy to molecular oxygen inside the cell. As a result, singlet oxygen is produced which destroys essential components of the cells by radical reactions. 

Low-level laser therapy (LLLT) low-power lasers or light-emitting diodes are used to stimulate cellular functions such as proliferation, mitosis, collagen synthesis etc. As a consequence enhanced wound healing reaction, reduced pain or other therapeutic effects are expected. LLLT is still controversial in mainstream medicine with ongoing research to determine whether there is a demonstrable effect.

Corneal crosslinking (CXL) was developed in the late 1990s. It has been shown in numerous clinical trials to strengthen the cornea through the application of riboflavin, a form of vitamin-B2, followed by treatment with ultraviolet A (UV-A) light.

Crosslinking has proven to be a first choice treatment for ectatic changes, typically marked by corneal thinning and an increase in the curvatures of the cornea. It often leads to high levels of myopia and astigmatism. The most common form of ectasia is keratoconus and less often ectasia is seen after LASIK.

For all applications associated with photochemical reactions temperature increase can occur as an unwanted side effect. This is unique compared with the following mechanisms, where temperature increase plays a key role.


With exposure times around milliseconds to a second and irradiance up to 10 W/cm2, photocoagulation can be achieved as a therapeutic effect. This interaction results from a thermally induced process associated with protein denaturation. Formerly Argon lasers were the most common laser sources , now frequency doubled Nd:YAG lasers or diodes emitting in the visible spectrum are applied. The photon energy is absorbed by the two main, natural chromophores within the ocular tissue: blood and melanin (Fig. 2), allowing the laser to target its effect on ocular structures such as vessels, the iris or deep retinal layers that contain melanin.

Photocoagulation of retinal vascular disease makes use of blood as a chromophore whereas laser trabeculoplasty and panretinal photocoagulation make use of the absorptivity of melanin to achieve an effect.

Another specific form of photocoagulation is used for photothermal shrinkage of tissue, known as thermo-keratoplasty [1, 2]. The concept is to induce shrinkage of the stromal collagen lamellae by delivering sufficient laser energy in the peripheral cornea, resulting in a circular band of tissue shrinkage with a central corneal steepening. Heating corneal tissue to temperature of around 50-60°C has been shown to shrink corneal collagen lamellae to approximately one third of its original length [3].

Holmium:YAG lasers have been used to induce tissue shrinkage by a noncontact delivery of laser energy using a 2120 nm wavelength and a pulse duration of 250 µs. Although short-term viability of this laser interaction for the correction of hyperopia and astigmatism could be demonstrated, most of the effect was lost in the long-term, as had been the case with previous methods of thermal shrinkage of corneal collagen.


Historically one of the first surgical laser applications was tissue cutting and removal at relatively high laser-beam intensities with exposure times of milliseconds to seconds, resulting in the rapid deposition of heat and subsequent vaporization of the water inside the tissue.

Today, mainly infrared laser light sources with comparatively high average power are used. Their wavelengths are strongly absorbed by water, allowing vaporization of any water-containing tissue, accompanied by thermal denaturation of the adjacent tissue. While this vaporization process is used in conventional surgery, where blood vessels are coagulated and to stop bleeding caused by an incision cut, this technique is barely used in ophthalmic surgery due to the wide collateral thermal damage.


Thermal collateral damage due to heat conduction can be minimized when the pulse duration of the laser is equal or shorter than the Thermal Relaxation Time. The Thermal Relaxation Time TR of a laser-heated region of tissue is the time required for the peak temperature to diffuse over the distance of the optical penetration depth δ of the laser light [4]:

TR = δ2 / (4 κ)

where κ is the thermal diffusivity of the tissue. Typically TR is in the range of microseconds to milliseconds when the optical penetration depth δ of the laser light is in the range of micrometers to millimeters.

If the pulse duration of the laser pulse is shorter than TR we call the process Photoablation.

The absorption coefficient of the corneal stromal tissue for laser light sources in the 3 um range, with its corresponding penetration depth δ, allowed for the production of lesions with greater precision and less thermal damage. The CO2 laser, Er:YAG laser or hydrogen fluoride laser have been tested for their longer strong tissue absorption to produce precise corneal excisions as well as surface ablation and keratectomy [5, 6].

Clinically proved corneal photoablation using the excimer laser was introduced in 1983 by Trokel and co-workers [7]. In their landmark article, they reported how this UV laser light source could precisely remove corneal tissue without any thermal side effects.  This revolutionary finding eventually led to the development of laser vision correction and LASIK as the most frequently performed elective procedure in all of medicine [8].

At that time the unique mechanism of interaction of excimer laser photoablation was understood as a pure photochemical process, where the individual photon energy (6.4 eV) is significantly higher than the energy required to break individual molecular bonds, which is not the case for longer laser wavelengths [9]. The subsequent ablation of tissue with UV laser occurs as the molecular fragments are ejected, driven by kinetic energy provided by the energy of the photon in excess of that required for bondbreaking (Fig. 3).

Today we know that excimer laser photoablation on corneal tissue is a mixture of both uv-bondbreaking of the collagen molecules and thermal vaporization of the water [10].

The Argon fluoride (ArF) gas mixture used to generate the 193nm wavelength achieves the highest level of tissue smoothness and precision in corneal ablation with the lowest amount of thermal damage (Fig. 4).  The 193 nm ArF gas mixture is the current excimer laser source used clinically all over the world to perform refractive surgery in a safe and efficient way.


The mechanism of interaction behind laser photodisruption is best described as plasma-mediated ablation, or optical breakdown. It relies on the nonlinear absorption of laser energy in the target achieved when the material specific radiant exposure is exceeded. Fundamentally, optical breakdown is characterized by three successive major events: plasma formation, shock wave generation and cavitation (Fig. 5).  The plasma is a highly ionized state of matter, and can be generated by laser pulses (from femtosecond to a few nanoseconds) of relatively low energy but high peak power. It has been shown that shortening the pulse duration from nanoseconds to femtoseconds decreases the threshold for plasma formation and reduces mechanical effects [11] (Tab. 1).

The process of optical breakdown can be explained by two different but equivalent models. On the one hand, we can regard the laser focus as an extremely strong electromagnetic field. Under the action of this field, electrons are stripped from their atoms and accelerated by the electric field to high kinetic energy.  The accelerated electrons in turn can collide with further atoms and ionize them. This process, which leads to plasma formation, is called "cascade ionization".

Another way to explain the process of ionization is with the photonic model. A free electron can be generated when the energy absorbed by a photon is higher than the energy necessary to excite the electron to the outermost energy level. Typically this energy is 6 to 10 eV. However, the photon energy at wavelengths within the range of 1.06 μm is only 1.17 eV. Thus, at this wavelength, which is typical for clinical lasers in transparent media, an extra six to ten photons would be necessary to promote an electron to leave it’s atom. Optical breakdown occurs when the irradiance is sufficient to produce a critical density of photons so that the probability of a simultaneous absorption of six or more electrons is considerably high.

After the laser pulse ends, the free electrons transfer their energy to the tissue by locally elevating the temperature that stays confined in the focal volume.  The thermal diffusion is too slow to dissipate the laser energy by heat conduction. So, the created plasma first expands at supersonic velocity emitting a shock wave due to its high temperature and pressure, and then slows down to the speed of the sound,. The elevation of the temperature creates a highly localized tensile stress which exceeds the critical tension for mechanical breakdown, resulting in tissue disruption and cavitation bubble formation. Mass spectroscopy analysis of the residual gas bubbles reveals a mixture of CO, CO2, methane, CH4, together with some fragments of CH3, CH2 and water vapour [12]. Photodisruption with ultrashort pulses achieves a fine and highly localized cutting without collateral thermal effects, but also has the benefit of tissue cleaving, due to the presence of rapidly expanding cavitation bubbles, which help to separate the tissue.  As a result, the sequential placement of tissue cutting pulses is a function of pulses energy. Using very low pulse energies in the range of some nanojoules, the cutting process is confined by the focal spot size of the laser pulse. As a consequence, more pulses are needed to cut the same area. To keep the total operation time at the same level, higher pulse repetition rates of some MHz are required (Fig. 6 left).

At comparatively high pulse energies, the cutting process is driven by mechanical forces which are applied by the expanding bubbles, disrupting the tissue. This cutting process is very efficient because the radius of disrupted tissue is larger than the laser spot size (Fig. 6 right). Hence, the spot separation of the scanned laser pulses can be larger than the spot diameter. 


To summarize, the physical mechanisms ruling the laser-tissue interaction differ strongly depending on the laser operation regime, thus allowing a broad variety of surgical applications. The interplay between wavelength, laser beam intensity and exposure time determine the type of laser-tissue interaction and can induce mechanical, thermal and/or chemical modifications, that result in either hemostatic effects, molecular denaturation (photocoagulation and vaporization), structural changes (photothermal shrinkage), tissue removal (photoablation) or cutting (photodisruption).


[1] Gruenberg P, Manning W, Miller D. Increase in rabbit corneal curvature by heated ring application. Ann Ophthalmol. 1981; 13(1):67-70.

[2] Horn G, Spears K, Lopez O. New refractive method for laser thermal keratoplasty with Co:MgF2 laser. J Cataract Refract Surg. 1990; 16(5):611-616.

[3] Shaw E, Gasset A. Thermokeratoplasty temperature profile. Invest Ophthalmol Vis Sci. 1974;13(1):181-186.

[4] Boulnois JL. Photophysical processes in recent medical laser developments: A review; Lasers in Medical Science, Volume 1, Number 1, 47-66, DOI: 10.1007/BF02030737

[5] Wolbarsht M, Foulks G, Esterowitz L. Corneal surgery with an Er:YAG laser at 2.94μm. Invest Ophthalmol Vis Sci. 1986;27(suppl):93.

[6] Peyman G, Badaro R, Khoobehi B. Corneal ablation in rabbits using an infrared (2.9μm) Er:YAG laser. Ophthalmol. 1989; 96(8):1160-1170.

[7] Trokel S, Srinivasan R, Braren B. Excimer laser surgery of the cornea. Am J Ophthalmol. 1983;96(6):710-715.

[8] Solomon KD, Fernandez de Castro LE, Sandoval HP, et al.  Joint LASIK Study Task Force. LASIK world literature review: quality of life and patient satisfaction.  Ophthalmol. 2009;116(4):691-701.

[9] Srinivasan R. Kinetics of the ablative photodecomposition of organic polymers in the far UV (193 nm). J Vac Sci Technol B. 1983;1(4):923-926.

[10] Staveteig, P.T., Walsh Jr., J.T; Dynamic 193-nm optical properties of water. Appl Optics. 1996; 35(19):3392-3403.

[11] Stuart BC, Feit MD, Herman S, Rubenchik AM, Shore BW, Perry MD.  Nanosecond-to-femtosecond laser-induced breakdown in dielectrics. Phys Rev B Condens Matter. 1996;53(4):1749-1761.

[12] Heisterkamp A, Ripken T, Mamom T, Drommer W, Welling H, Ertmer W, Lubatschowski H; Nonlinear side effects of fs-pulses inside corneal tissue during photodisruption; Appl.Phys.B 2002;74:1-7.

Content Overview

Laser-Tissue Interaction Regimes
Photochemical Reactions

© by Holger Lubatschowski, 2012

Figure 1
Figure 2
Figure 3
Figure 4
Figure 5
Table 1
Figure 6
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