The Science of Laser Medicine, which includes Laser Surgery, is considered a branch of Medical Physics, which, in its turn, is one of the branches of Medicine. Similar to Applied Physics and Engineering, Medicine is one of the Applied Sciences.
The American Laser Study Club (ALSC) has designed its Surgical Laser educational curriculum based on peer-reviewed literature with its foundation in laser-tissue interaction and laser surgery.[1-24] This curriculum was created by Dr. Peter Vitruk, ALSC founder, and is dedicated to the detailed physics of soft-tissue ablation and coagulation with lasers and hot tip (non-laser) devices. A brief overview of surgical lasers in terms of their ablative and coagulative properties is presented in Video 1 below. The ALSC’s curriculum overcomes the known limitations of many laser dentistry education programs that ignore laser-tissue interaction science.[25,26]
Introduction to Laser Surgery Physics
Not all lasers are efficient at simultaneous cutting and coagulating 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 soft-tissue. On this page we discuss, based on [1-24], how the wavelength impacts the photo-thermal coagulation and ablation efficiencies in oral soft-tissue surgeries with surgical Near-Infrared (Near-IR) Diode, IR CO2 and Mid-IR Erbium lasers.
The photo-thermal (radiant) ablation of 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 surgical laser instruments and their emission wavelengths have indications for use for incising, excising … oral soft-tissue surgery” , and “the key to the usefulness of the Nd:YAG is that this wavelength is highly absorbed in oral soft-tissue” , are in conflict with : “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  “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 . “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” .
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 ) and photo-thermal ablation (or photovaporolysis ) efficiencies for surgical lasers: 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 key to understanding how the laser light ablates (e.g. vaporizes) and coagulates soft-tissue.
For practical ablative soft and hard-tissue surgical 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 an increased tissue temperature that, in turn, may lead to tissue coagulation and ablation. Such laser-tissue interaction is referred to as photo-thermal.
Consider, see Figure 2a, a one-dimensional presentation of a laser beam entering the tissue from the left at x=0. Laser light intensity (or power density) immediately below the tissue’s surface is I0. The laser beam’s incident intensity (W/cm2) is IB. The transmission of the tissue’s surface is I0/IB, and the reflectivity of the tissue’s 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.
Hard-Tissue Dental Lasers Principles
Soft-tissue lasers should not be confused with hard-tissue lasers. The simplicity of 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 the 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” , while “plasma emissions … may contain sufficient UV” , requiring the UV exposure limits  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 to the high translucence of the teeth . By such a comparison, the simplicity of the oral soft-tissue lasers can be even more appreciated.
The 9,250-9,600 nm CO2 laser wavelengths are strongly absorbed by hydroxyapatite in enamel or dentin; their laser beam energy is utilized to vaporize hard-tissue at very high temperatures, leading to strong thermal side effects mentioned above.
Erbium lasers (both the 2,940 nm and 2,780 nm ) can overcome the strong thermal side effects since these wavelengths are absorbed much stronger by water (see Figure 1) relative to hydroxyapatite. Short and high-intensity Erbium laser pulses cause the explosive vaporization of water, trapped near the surface in enamel or dentin, which creates enormous subsurface water vapor pressure that fractures and ejects the surrounding fractured mineral structure. For safety reasons, the explosive nature of hard-tissue removal severely limits the speed and depth of ablation by Erbium lasers, which makes Erbium lasers significantly slower than regular mechanical drill handpieces. Since water vaporization happens at a significantly lower temperature than hydroxyapatite vaporization, hard-tissue Erbium laser plumes do not glow like hard-tissue CO2 laser plumes. However, both Erbium (2,940 nm and 2,780 nm) lasers and CO2 lasers (at 9,250-9,600 nm) require, just like mechanical drills, an intense spray of cooling water on the hard-tissue in order to prevent the pulp from overheating.
For the majority of dentists, the high-speed drill handpiece still remains superior to 2,780, 2,940 and 9,250 nm lasers since these modern-day hard-tissue lasers still do not deliver on the promise of lesser (relative to mechanical handpiece) thermal effects, they do not eliminate the need for anesthesia and also remain prohibitively expensive.
Chromophores in Epithelium and in Connective Tissue
Spatial distribution and concentration of chromophores in the 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, the 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  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  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 the absorption coefficient and reduced scattering coefficient.
The 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 the epidermis: 1.3-6.3% for light-skinned, 11-16% for well-tanned skin, and 18-43% for darkly pigmented African type skin . Epithelial thickness is chosen as 200 µm – an average value from OCT-measurements of the oral epithelium ; 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  for skin and  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 the 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 .
Sub-epithelial Light Absorption and Scattering
Figure 7 presents the absorption depth spectrum for sub-epithelium with 75% water and estimated 10% blood  (assuming whole blood contains hemoglobin (and/or oxyhemoglobin) at a 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 the absorption coefficient [3-6] and reduced scattering coefficient (estimated through absorption to reduced scattering ratio from ). 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.
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 the 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 : the heat propagation distance is proportional to the root square of time that the 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 /(ϱ C) ≈ 0.155 (+/-0.007) mm2/sec is tissue’s thermal diffusivity. λ ≈ 6.2-6.8 mW/cm o C is tissue’s heat conductivity; C ≈ 4.2 J/g oC is the specific heat capacity and ϱ ≈ 1 g/cm3 is the density for liquid water .
The irradiated tissue is heated most efficiently when the laser pulse is shorter than Thermal Relaxation Time TR. The irradiated tissue is cooled most efficiently between pulses when the time duration between 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 photothermal ablation (or photovaporolysis) [5, 6, 17]. The volume of 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 Fluence (Energy Density) Threshold. We consider the conditions for the highest efficiency photo-thermal ablation (pulse duration t ≤ TR) with minimum collateral damage to the surrounding tissue (pulse repetition rate f << 1/TR). The minimum fluence, or energy density, required 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 a 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  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 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 the scarcity of light scattering coefficient data for oral soft-tissue (which are important for this wavelength range ) – 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 soft-tissue due to extremely short absorption depths (see Figures 6 and 7).
Spatial Accuracy of Photo-Thermal Ablation. Laser 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 a 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 : “vital structures … may be heavily damaged before tissue ablation at the surface-initiated” and these wavelengths are referred to in  as “poor scalpels” and in  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 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 ) 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  for laser pulses shorter than Thermal Relaxation Time (pulse duration t ≤ TR) and is presented in Figure 11. Coagulation depths for 1,100-1,800 nm range were not calculated in view of the scarcity of light scattering coefficient data for oral soft-tissue (which are important for this wavelength range ) – hence the inability to calculate the attenuation depth at these wavelengths.
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 a 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 .
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 , 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 soft-tissue photo-thermally.
The Near-IR 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.
Surgical CO2 Lasers
The CO2 laser’ radiant energy is used directly to photo-thermally cut, ablate, and, at the same time, photo-thermally coagulates 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 12a. In the 1990s the flexible hollow fibers from Luxar Corp., WA, USA (www.lightscalpel.com) significantly improved the ergonomics of the CO2 laser surgery.
Laser Power and the Rate of Tissue Ablation
For a laser scalpel, e.g., CO2 or Erbium lasers, the power density of the focused laser beam is equivalent to the mechanical pressure that is applied to a cold steel blade. Greater laser fluence (i.e., power density times the duration is applied to the target) results in greater depth and rate of soft-tissue removal. For short pulse steady-state ablation conditions (xa << A, see Figure 8) the ablation depth is: A (E – ETH)/ETH, where A is the absorption depth and ETH is the ablation threshold fluence from Figure 9, and E is the fluence delivered to the tissue. For repetitive pulses that are scanned across the soft-tissue, the depth of incision is proportional to laser average power and is inversely proportional to focal spot diameter and the surgeon’s hand speed (see Figure 12b).
For a very high power laser beam that is moving over the target tissue (e.g. for fast and deep skin incisions with focused round beams or for shallow ablation with rectangular laser beams in veterinary surgery), the vaporization/ablation depth is illustrated in Figure 12d, and it is also:
- proportional to laser average power, and
- inversely proportional to laser beam width, and surgeon’s hand speed.
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 – see Figure 12c. “Painting” the “bleeder” with the large diameter defocused laser beam is just one of the practical techniques that do 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 the laser pulse is shorter than the Thermal Relaxation Time TR.; the most efficient tissue cooling occurs when the 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.
Hot Tip Cutting with Diode Lasers
The near-IR laser wavelengths of diode lasers cannot photo-thermally ablate soft-tissue [4-7] (except for high melanin content epithelium), as illustrated in Figure 14.
Since the 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 . 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 diode lasers perform as a NON-LASER WAVELENGTH-INDEPENDENT ablation thermal devices with coagulation depth in the range 250-1,000 µm .
Figure 16 reproduces the hot tip coagulation depths measurements  conducted under constant tip temperature (red circles) and under constant power (blue circles) conditions. Also presented are the calculated  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.
Since 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  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 ; and
- Biocompatibility  of the tip can be compromised by the burning of the fiber’ cladding materials at 500-900°C operating temperatures; and
- Sterility  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 ; 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 ) depends strongly on tip-tissue contact time and hand speed .
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