This introduction to Soft-Tissue Laser Oral Surgery Physics is based on American Board of Laser Surgery’s (ABLS) Study Guide, ABLS’s position on Soft Tissue Dentistry, and Peter Vitruk’s work listed on the about page.

Introduction to Soft-Tissue Laser Oral Surgery Physics

Not all lasers are efficient at simultaneous cutting and coagulating of the soft tissue. Some laser wavelengths (such as of Erbium lasers) are great at cutting and ablation, but are not as efficient at coagulating. Other laser wavelengths (such as of diode lasers) are highly efficient coagulators, but are poor scalpels. There are also lasers (such as a CO2 laser) that are efficient at both cutting and coagulating of the soft tissues. On this page we discuss, based on [1-24], how the wavelength impacts the photo-thermal coagulation and ablation efficiencies in oral soft tissues surgeries with dental Near Infrared (Near-IR) Diode, IR CO2 and Mid-IR Erbium lasers.

The photo-thermal (radiant) ablation of the soft tissue has been comprehensively reported [5-7,17], and yet there is an incongruity between (a) the well documented and researched data on INEFFICIENT soft tissue absorption / ablation in the Near-IR 800-1,100 nm spectral range, and (b) the widely proliferated misconception about EFFICIENT Near-IR laser ablation of the oral soft tissue.

Indeed, the statements such as “all currently available dental laser instruments and their emission wavelengths have indications for use for incising, excising … oral soft tissue surgery” [25], and “the key to the usefulness of the Nd:YAG is that this wavelength is highly absorbed in oral soft tissue” [26], are in conflict with [5]: “Lasers whose extinction length is 5 mm or more, and whose δ/α ratio (scattering to absorption ratio) is larger than 10 make good coagulators but poor scalpels. Such wavelengths are all in the near-infrared (700-1400 nm) region.” Furthermore, an observation in [17] “Using laser wavelengths where optical scattering is comparable to or dominant over tissue absorption is not conducive to precise ablation” is directly related to 800-1,100 spectral range nm according to Figure 5 from [17]. “An important complicating factor in the use of this high-power laser light, which penetrates deeply before being absorbed totally, is that it may reach vital structures in the vicinity of the target tissue. These vital structures may preferentially absorb near-infrared laser light because of different optical properties and may be heavily damaged before efficient tissue ablation at the surface initiated” [18].

optical absorption coefficient oral soft tissue lasers

Figure 1a. Optical absorption coefficient spectra at different histologically relevant concentrations of Water, Hemoglobin, Oxyhemoglobin and Melanin based on data from.[1-6] Logarithmic scales are in use.

Optical laser absorption coefficient spectra

Figure 1b. Optical absorption coefficient spectra at different histologically relevant concentrations of water, hemoglobin (Hb), oxyhemoglobin (HbO2), and melanin. Logarithmic scales are in use.

To clarify the above inconsistencies, we turn to the absorption spectra of the four main chromophores of the oral soft tissue [1-7]see Figure 1 – namely: hemoglobin (Hb), oxyhemoglobin (HbO2), melanin, and water.  These spectra form the foundation for the analysis of the photo-thermal coagulation (or photopyrolysis [5]) and photo-thermal ablation (or photovaporolysis [5]) efficiencies for the soft tissue lasers in dentistry: the IR CO2 laser at 9,300 nm and 10,600 nm; the Mid-IR Erbium lasers at 2,780 nm and 2,940 nm; and the Near-IR diodes at 808 – 1,064 nm.

Photo-Thermal Ablation, Coagulation and Heat Affected Zone

Light absorption (see Figure 1) and light scattering (that dominates over light absorption at diode lasers’ 800-1,100 nm wavelength range [3-7, 17]) are the key to understanding how the laser light ablates (e.g. vaporizes) and coagulates the soft tissue.

For practical ablative soft and hard tissue dental lasers on the market today (diode, Erbium and CO2 lasers), the laser light energy is converted, through absorption, into the thermal energy inside the tissue leading to increased tissue temperature that, in turn, may lead to tissue coagulation and ablation. Such laser-tissue interaction is referred to as photo-thermal.

tissue density and laser beam intensity spatial distribution

Figure 2. (a), (b) tissue’ density and laser beam intensity spatial distribution before ablation; laser beam is directed at the tissue surface x=0 from the left; (c), (d) tissue’ density and temperature spatial distributions after the pulse.

Consider, see Figure 2a, a one dimensional presentation of a laser beam entering the tissue from the left at x=0. Laser light intensity immediately below the tissue’ surface is I0. Laser beam’s incident intensity (W/cm2) is IB. The transmission of the tissue’s surface is I0/IB, and the reflectivity of tissue’ surface is (IB – I0)/IB. For x>0, i.e. inside the tissue in Figure 2b, the laser light intensity is exponentially decreasing:

I = I0 Exp [-x/A]                                (1)

where A is the depth of absorption, and 1/A is the coefficient of absorption [1-7].

If the laser intensity I0 is greater than intensity Ia ­needed to ablate the tissue in the thin sub-surface layer 0<x<xa (for a specific pulse duration t), see Figure 2b, the tissue ablation takes place and layer 0<x<xa referred to as “Ablation Zone” in Figures 2c and 2d. Inside the heat affected zone xa<x<xc the tissue temperature is ranging from the ablation temperature Ta down to the coagulation temperature Tc. Coagulation depth H = xc – xa, is defined by 60-100 OC temperature range [7, 19-22] inside the heat affected zone in Figure 2 (d) (i.e. Tc = 60 OC and Ta = 100 OC). Normal body temperature is Tb < Tc.

The soft tissue lasers should not be confused with the hard tissue lasers. The simplicity of the soft tissue 10,600 nm CO2 laser surgery (Figure 3a) is largely based upon the low temperature water vaporization at 100OC, and the collateral damage in heat affected zone is simply heat induced coagulation and hemostasis. In hard tissue cutting applications, however, a very high ablation temperatures Ta (as high as 5,000 OC) could result in extremely bright thermal radiation – see Figures 3b and 3c. Heat induced enamel melting on the margins of the cut may compromise the bonding strengths for adhesives. The hard tissue 9,250-9,600 nm CO2 laser’s “beam interactions with the hard tissue can generate intense plasma emissions …requiring suitable optical filtering for direct viewing” [27], while “plasma emissions … may contain sufficient UV” [28], requiring the UV exposure limits [29] to be addressed. Also, the high brightness of the hydroxyapatite plasma’ in the visible spectrum (see Figures 3b and 3c) may interfere with tooth’ visibility due the high translucence of the teeth [30]. By such comparison, the simplicity of the oral soft tissue lasers can be even more appreciated.

soft and hard tissue ablation

Figure 3. Differences between the soft and hard tissue ablation: (a) soft tissue 10,600 nm laser cutting takes place at low (100O C) ablation temperatures, as opposed to (b) hard tissue 9,300 nm laser cutting with ablation temperatures in excess of 5,000O C. (c) Laser plume emission spectrum recorded with 200-850 nm range spectrometer with 1.5nm DWHM resolution (BWTek Inc., Newark, DE, model BRC115U, SN120911304). The enamel of the freshly extracted (< 24 hours) human tooth was ablated with the experimental 9.3µm CO2 laser (500 Hz with sub-100 µsec pulses with average power of 2.5 Watts with 250µm spot size); water irrigation was used. Emission peak around 530-560 nm corresponds to plume’s temperature in excess of 5,000O C in Plank’s black body thermal radiation approximation. Photos and data are courtesy of LightScalpel LLC, Washington, USA.

Chromophores in Epithelium and in Connective Tissue

soft tissue laser beam

Figure 4. Simplified optical model of Oral Soft Tissue consisting of (1) Water-Melanin rich Epithelium layer, and (2) Water-Hemoglobin-Oxyhemoglobin rich sub-epithelium.

Spatial distribution and concentration of chromophores in epithelium and connective tissue are important factors in understanding oral soft tissue-laser interaction.

The melanin is only present in the epithelium layer. The hemoglobin is only present in the connective tissue (sub-epithelium). Therefore, optical properties of epithelium and sub-epithelium need to be analyzed separately and independently from each other as presented in Figure 4:

  • Optical absorption in the 100-300 µm thin [8] epithelium layer is dominated by water and melanin;

  • Optical absorption in the connective tissue, inclusive of lamina propria and submucosa [9,10] (the sub-epithelium medium), is dominated by water and hemoglobin / oxyhemoglobin.

Epithelial Light Absorption and Scattering

Since the 100-300 µm [8] thin epithelium can be significantly thinner than the absorption depth A in the near-IR, the epithelium optical properties are best described not by absorption (or attenuation) depth A, but rather by the percentage of light absorbed as it passes through the epithelium.  The attenuation depth is defined as an inverse of the sum of absorption coefficient and reduced scattering coefficient.

laser optical absorption

Figure 5. Optical absorption (and estimated Near-IR attenuation) spectra of 200 µm thick epithelium.

Absorption spectrum of epithelium is presented in Figure 5 for three cases of volume fraction of melanin pigmentation of 2%  (very light), 13% (moderate) and 30% (dark), similar to pigmentation in epidermis: 1.3-6.3% for light skinned, 11-16% for well-tanned skin, and 18-43% for darkly pigmented African type skin [3]. Epithelial thickness is chosen as 200 µm – an average value from OCT-measurements of oral epithelium [8]; it is much thinner than the absorption depth A for Near-IR, especially for lightly pigmented epithelium – see inset “A” in Figure 5. In view of strong light scattering in the Near-IR wavelength range [3-6] (reduced scattering coefficient 20 cm-1 [4] for skin and [13] for epithelium), an estimate of the attenuation depth is presented as inset “B” in Figure 5. The inset “C” in Figure 5 illustrates how different the attenuation (affected by both absorption and scattering) is from absorption alone (i.e. without scattering) in inset “A”.

Optical absorption in epithelium relatively low and is highly dependent on pigmentation in 800-1,100 nm spectral range. In sharp contrast to Near-IR wavelengths, the IR wavelengths (CO2 laser) and Mid-IR wavelengths (Erbium lasers) display almost 100% absorption (see Figure 1) and very low penetration depths (see Figure 6) regardless of epithelium pigmentation, which is important for reproducible laser removal of the epithelium layer [11].

light absorption soft tissue lasers

Figure 6. Optical absorption depth spectra of 75% water-rich soft tissue (either epithelium or sub-epithelium connective tissue) at Mid-IR and IR wavelengths based on data from.[1-6]

Sub-epithelial Light Absorption and Scattering

Figure 7 presents the absorption depth spectrum for sub-epithelium with 75% water and estimated 10% blood [12] (assuming whole blood contains hemoglobin (and/or oxyhemoglobin) at normal concentration of 150g/L [5, 6]), which are derived from absorption coefficient spectra (Figure 1) for water [1, 2], hemoglobin and oxyhemoglobin [4-6]. The inset in Figure 7 presents the estimated attenuation depth as an inverse of the sum of absorption coefficient [3-6] and reduced scattering coefficient (estimated through absorption to reduced scattering ratio from [14]). Since the light scattering dominates over absorption [3-6, 14] in the Near-IR spectrum, the attenuation depth is a more accurate representation of laser energy penetration into the tissue for Near-IR wavelengths.

absorption depth soft tissue lasers

Figure 7. Absorption (and estimated Near-IR attenuation) depth spectra of sub-epithelium (connective tissue). Logarithmic scale is in use.

Thermal Relaxation Time

Tissue cooling from the blood flow during laser exposure is insignificant for short pulse durations of practical interest, and especially when tissue coagulation takes place underneath the ablated tissue. Instead, the cooling efficiency of the tissue irradiated by laser light is determined by tissue’s own thermal conductivity (or thermal diffusivity) to dissipate (or diffuse) the heat away from the irradiated tissue.

The thermal diffusivity process was first quantified in [15]: the heat propagation distance is proportional to the root square of time that heat source is on. The Thermal Relaxation Time (presented in Figure 7) as TR = A2/K [7,16,17], where A is optical absorption depth discussed above in (1) , defines how fast the heat diffuses away from the irradiated volume of the tissue. K = λ / C) ≈ 0.155 (+/-0.007) mm2/sec is tissue’s thermal diffusivity. λ ≈ 6.2-6.8 mW/cm oC is tissue’s heat conductivity; C ≈ 4.2 J/g oC is specific heat capacity and ϱ  ≈ 1 g/cm3 is density for liquid water [23].

The irradiated tissue is heated the most efficiently when the laser pulse is shorter than Thermal Relaxation Time TR. The irradiated tissue is cooled the most efficiently between the pulses when the time duration between the laser pulses is much greater than TR. Such laser pulsing is referred to as SuperPulse and is a must-have feature of any state-of-the-art soft tissue surgical CO2 laser that minimizes the char and the depth of coagulation.

Efficiency of Photo-Thermal Ablation

The process of fast vaporization of the intra- and extra-cellular water is at the core of soft tissue photo-thermal ablation (or photovaporolysis) [5, 6, 17]. The volume of the irradiated tissue is proportional to the absorption depth (or Near-IR attenuation depth as defined above). The lesser laser energy is required to ablate the tissue for shorter absorption depths. The greater laser energy is required to ablate the tissue for longer absorption depths.

Photo-Thermal Ablation Energy Density Threshold. We consider the conditions for highest efficiency photo-thermal ablation (pulse duration tTR) with minimum collateral damage to the surrounding tissue (pulse repetition rate f << 1/TR). The minimum energy density requirement to vaporize the irradiated soft tissue can be calculated from the exponentially attenuated spatial distribution of laser light intensity (1); a graphical representation of such laser beam is given in Figure 8, where xa is the ablation depth, and H = xa – xc is the coagulation depth illustrated in Figures 2c and 2d. The threshold ablation energy density ETH is derived in [7] and is illustrated in Figure 9. The threshold energy densities in Near-IR 800-1,100 nm spectrum are 1,000s times greater than Mid-IR and IR wavelengths because of much weaker Near-IR absorption by the soft tissue – see Figure 1.

Figure 10 illustrates the high degree of scattering and absence of tissue ablation at 810 nm and 980 nm. Ablation thresholds for 1,100-1,800 nm range are not calculated in view of scarcity of light scattering coefficient data for oral soft tissue (which are important for this wavelength range [17]) – hence the inability to calculate the attenuation depth at these wavelengths. Unlike the Near-IR wavelengths, the IR and Mid-IR wavelengths (CO2 and Erbium lasers) are highly energy efficient at ablating the soft tissue due to extremely short absorption depths (see Figures 6 and 7).

diode soft tisse laser absorption

Figure 10. Near-IR wavelengths 810 and 980 nm are highly scattered and weakly absorbed by the porcine soft tissue, resulting in slow and wide-spread photo-coagulation and no ablation. Photos are courtesy of LightScalpel LLC, Washington, USA.

Spatial Accuracy of Photo-Thermal Ablation. Dental diodes’ wavelengths 800-1,100 nm are poorly absorbed by the low concentration melanin in the thin epithelium and by the scarce hemoglobin and oxyhemoglobin in connective tissue; this leads to multi-millimeter depth of laser energy penetration into the oral soft tissues (Figures 5 and 7). The multi-millimeter penetration increases the collateral damage risk, which is referred to in [18]: “vital structures … may be heavily damaged before tissue ablation at the surface initiated” and these wavelengths are referred to in [5] as “poor scalpels” and in [17] as “not conducive to precise ablation”.

The IR wavelengths (CO2 lasers) and the Mid-IR wavelengths (Erbium lasers) are characterized by a much shorter absorption depths (see Figures 6 and 7), which makes these lasers preferable for the soft-tissue ablative laser applications.

Efficiency of Photo-Thermal Coagulation

Soft tissue coagulation occurs as a denaturation of soft tissue proteins above 60°C [19-22] leading to less bleeding and less oozing from the lymphatics on the margins of the laser cut. The diameter of blood vessels B (21-40 µm from measurements in human cadaver gingival connective tissue [24]) influences the efficacy of the photo-thermal coagulation and hemostasis (shrinkage of the collagen-rich walls of blood vessels and lymphatic vessels at elevated temperatures).

The coagulation depth value H (illustrated in Figures 2d and 8) relative to the blood vessel diameter B is an important measure of coagulation and hemostasis efficiency, and is derived in [7] for laser pulses shorter than Thermal Relaxation Time (pulse duration tTR) and is presented in Figure 11. Coagulation depths for 1,100-1,800 nm range were not calculated in view of scarcity of light scattering coefficient data for oral soft tissue (which are important for this wavelength range [17]) – hence the inability to calculate the attenuation depth at these wavelengths.

laser coagulation depth

Figure 11. Coagulation depth spectrum for ablation threshold conditions. Logarithmic scale is in use.

For Erbium laser wavelengths in Figure 11, the absorption and coagulation depths are significantly smaller than blood vessel diameters, i.e. H<<B. Coagulation is confined to relatively thin layer on the margin of the laser cut and bleeding control is relatively low. Coagulation depth can be increased by pulse width/rate increase.

For diode laser wavelengths in Figure 11, the absorption (and Near-IR attenuation) and coagulation depths are significantly greater than blood vessel diameters, i.e. H>>B. Photo-thermal coagulation extends far from ablation margins as documented in [18].

For the CO2 laser wavelengths in Figure 11, H ≥ B and the coagulation extends just deep enough into the margins of the laser cut to stop the bleeding from the severed blood vessels. Note the coagulation depth in Figure 11 at 10,600 nm is about 50% thinner than at 9,300 nm, also observed in [31], which is due to 50% stronger absorption at 10,600 nm in Figures 1, 6, 7. Coagulation depth can be further increased by pulse width/rate increase.

Photo-Thermal Ablation and Coagulation Summary

A close match between the oral soft tissue blood capillary diameters and coagulation depth for the IR CO2 laser wavelengths makes the CO2 laser the ONLY practical soft-tissue surgical laser, which uses the laser beam directly to both cut and coagulate the soft tissue photo-thermally.

The Near-IR dental diode lasers cannot cut the oral soft tissue photo-thermally. Instead, the NON-LASER WAVELENGTH-INDEPENDENT thermal ablation by the diode’s charred hot glass tip is the only technique used for the oral soft tissue cutting.

The Mid-IR Erbium lasers wavelengths are the best at cutting, but feature reduced coagulation efficiency, which limits their soft-tissue applications.

Soft Tissue Dental CO2 Laser

dental co2 soft tissue laser

Figure 12. Oral soft tissue 10,600 nm CO2 lasers with Articulated Arm (left) and Flexible Fiber (right) laser beam deliveries. Photo courtesy of LightScalpel, Washington, USA.

The CO2 laser’ radiant energy is used directly to photo-thermally cut, ablate and, at the same time, photo-thermally coagulate the soft tissues. The early days CO2 surgical lasers in 1970-80s employed the bulky 7-mirror articulated arms to deliver the laser beam to the surgical site – see Figure 12. In the 1990s the flexible hollow fibers from Luxar Corp., WA, USA ( significantly improved the ergonomics of the CO2 laser surgery.

Laser power and the rate of tissue ablation. The power density of the focused laser beam is similar to the mechanical pressure that is applied to a cold steel blade: the greater the laser power density, the deeper the laser cuts and the greater the rate of soft tissue removal.

Laser beam spot sizes for cutting and coagulation. Just like the sharpness of the steel blade defines the quality and the ease of the cut, the size of the laser beam focal spot defines the quality of the laser cut. Just like a dull blade cannot produce a quality incision, the large diameter laser beam will not cut. The small diameter focal spot of the beam enables the narrow and deep incision with sufficient laser energy. The laser beam can be defocused through increasing the distance between the laser aperture and the tissue. “Painting” the “bleeder” with the large diameter defocused laser beam is just one of the practical techniques that does not require to change the laser settings from the control panel.

Controlling thermal effects. The so called “SuperPulse” design for CO2 laser pulsing parameters is optimized around the Thermal Relaxation Time concept discussed above and illustrated in Figure 13. The most efficient cutting occurs when laser pulse is shorter than the Thermal Relaxation Time TR.; the most efficient tissue cooling occurs when spacing between the pulses is much greater than the Thermal Relaxation Time TR. The SuperPulse results in less char on the margins of the cut, which helps post-operative healing and reduces scarring of surgical wounds.

superpulse laser explained

Figure 13. SuperPulse explained: high power, short laser pulse duration maximize soft tissue removal rate and keep adjacent tissue cool. Graph courtesy of LightScalpel, Washington, USA.

Hot Tip Cutting with Dental Diodes

The near-IR laser wavelengths of dental diode lasers cannot photo-thermally ablate soft tissue [4-7] (except for high melanin content epithelium), as illustrated in Figure 14.

co2 laser vs diode laser incision

Figure 14. Unlike 10,600 nm CO2 laser, the 810nm near-IR diode laser cannot cut soft tissue photo-thermally at comparable values of laser fluence.

glass tip diode laser

Figure 15. “Initiated” (i.e. charred per manufacturer’s instruction of burning corkwood at 1 Watt) glass tip of the 810 nm near-IR diode laser at 3 Watts.

Since the dental diode Near-IR laser wavelengths are not suitable for the oral soft tissue cutting, there is a different technique of non-optical (non-radiant) tissue cutting with the so called “hot tip”, as illustrated in Figure 15. A critical component of such hot tip-tissue interaction is the creation of the optically dark char deposit on the very end of the glass tip. The diode laser optical energy is then used to heat up the charred glass fiber’s tip up to 900OC [32]. Such hot glass tip heats up the soft tissues through heat conduction (i.e. heat diffusion) from (and through) hot glass tip to (and through) the soft tissue. Such charred glass tip dental diodes perform as a NON-LASER WAVELENGTH-INDEPENDENT ablation thermal devices with coagulation depth in the range 250-1,000 µm [32].

Figure 16 reproduces the hot tip coagulation depths measurements [32] conducted under constant tip temperature (red circles) and under constant power (blue circles) conditions. Also presented are the calculated [33] coagulation depths for constant tip temperature (red line) and for constant power (blue line) conditions. As can be seen, the hot glass tip has the capability to significantly limit the spread of thermal necrosis to less than 1 mm vs. 10+ mm Near-IR photo-thermal coagulation depth from Figure 11 (also presented in Figure 13). Hot tip coagulation depth depends strongly on tip-tissue contact time (i.e. hand-speed) and significantly exceeds the gingival blood vessel diameters and is much greater than CO2 laser coagulation depth from Figure 11.

diode yag laser hot tip coagulation depth

Figure 16. Hot tip coagulation depth as a function of tip-tissue contact time and hand speed. Logarithmic scale is in use.

Since the soft tissue ablative diodes are contact thermal devices, the device-tissue interface (charred hot glass surface) controls the medical efficacy and is highly dependent on a multiplicity of factors such as:

  • The so called “initiation” or “activation”, i.e. placing the light-absorbing CHAR layer on the tip before the surgery; and
  • Maintaining such CHAR layer on the tip as such tip is dragged through the tissue; and
  • Controlling the effectiveness of the light-absorbing char layer on the tip end; and
  • A “partially transmitting” diode laser fiber tip irradiates the soft tissue with near-IR laser radiation, which is not conducive for ablation and cutting, but is a good coagulator whose depth and width are strongly affected by the fluence and by the tissue- and wavelength-dependent absorption and scattering properties; and
  • Degradation of the char on glass tip’ surface creates the “optically leaky” tip with reduced tip’ temperature, and increased risk of Near-IR induced sub-surface thermal necrosis [18] as well as mechanical tearing of the tissue by the cold glass’ sharp edges; and
  • Overheating of the glass may result in thermally induced fractures of the glass [34]; and
  • Biocompatibility [35] of the tip can be compromised by the burning of the fiber’ cladding materials at 500-900°C operating temperatures; and
  • Sterility [36] of the tip can be compromised by non-sterile inks and cork used in tip initiation process; and
  • Thermomechanical thresholds for thermal gradient-induced fractures of the hot glass tip at 500-900°C operating temperatures and biocompatibility of fractured glass fragments [37]; and
  • User’s skill and technique of controlling tip contact with the tissue and hand movement, since the hot tip coagulation depth (250-1,000 µm [32]) depends strongly on tip-tissue contact time and handspeed [32].


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