The Science of Laser Medicine, which includes Laser Surgery, is considered a branch of Medical Physics, which, in 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 Erbium lasers) are great at cutting and ablation but are not as efficient at coagulating. Other laser wavelengths (such as diode lasers) are highly efficient coagulators but are poor scalpels. There are also lasers (such as a CO2 laser) that are efficient at 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 a contradiction 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” , 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 using 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 crucial 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 absorption depth, and 1/A is the absorption coefficient [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 ranges from the ablation temperature Ta down to the coagulation temperature Tc. Coagulation depth H = xc – xa is defined by 60-100° C temperature range [7, 19-22] inside the heat-affected zone in Figure 2 (d) (i.e., Tc = 60° C and Ta = 100° C). 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 based mainly on the low-temperature water vaporization at 100° C. The collateral damage in the heat-affected zone is simply heat-induced coagulation and hemostasis. In hard-tissue cutting applications, however, very high ablation temperature Ta (as high as 5,000° C) could result in extremely bright thermal radiation – see Figures 3b and 3c. Heat-induced enamel melting on the cut’s margins may compromise adhesive bonding strengths. 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 the strong thermal side effects mentioned above.
Erbium lasers (2,940 nm and 2,780 nm) can overcome the strong thermal side effects since water absorbs these wavelengths much stronger (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 substantially 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 to prevent the pulp from overheating.
According to the American Dental Association (ADA) Guide on Dental Lasers, “the term ‘hard tissue’ … refers to dentin, enamel, bone, and cartilage. All other biological structures, including dental pulp, are considered to be soft tissue. Lasers with wavelengths highly absorbed in water and protein are ideally suited for osseous recontouring, dental preparations, and conservative caries removal, with minimal healthy tissue loss. Laser pulses can preferentially ablate carious tissue due to the higher percentage of water and protein present in carious tissue at a higher ratio than in normal tissue. Hard tissue ablation with Erbium lasers is accomplished on dentinal and osseous structures utilizing wavelengths between 2,780 and 2,940 nm that are highly absorbed in water. The higher the absorption of the laser energy into the water contained within the structures is vaporized, the more efficiently the water molecule is converted into steam / vaporized. The rapid expansion of the water as it is vaporized disrupts the crystalline matrix of the bone and dentition, thus removing the structures as part of the laser plume. Dental preparations accomplished with an Erbium laser ablation create a rougher surface, enhancing the bonding strengths to the restorative material when compared to a conventional preparation performed with rotary instrumentation. Photo-thermal ablation of hard tissue with carbon dioxide (CO2) lasers, with wavelengths in the 9,200-9,600 nm range, is facilitated by the high absorption coefficient of these wavelengths in the hydroxyapatite. During the photo-thermal laser ablation (or vaporization) of solid materials, the laser energy heats up, melts, and vaporizes the molten material. Accordingly, CO2 lasers with wavelengths in the 9,200 to 9,600 nm range and pulse durations in the microsecond range accomplish hard tissue ablation as the hydroxyapatite strongly absorbs the laser irradiation and heats up and melts it at temperatures greater than 1,300 C. The molten hydroxyapatite is vaporized at several thousand degrees Celsius. Like Erbium wavelengths’ rapid expansion of water, the rapid expansion during the vaporization process disrupts the crystalline matrix of the hard tissues, thus removing the structures as part of the laser plume. Additionally, on the margins of the vaporized hard tissue (such as enamel or dentin), the CO2 laser irradiation heats the hydroxyapatite, which changes the structure and chemical composition of the remaining mineral of enamel and dentin. When the molten material cools, a glazed type of surface occurs”.
Examples of Erbium laser wavelengths include the Waterlase by Biolase at 2,780 nm and the Lightwalker by Fotona at 2,940 nm, and examples of hard tissue CO2 laser wavelengths include the Solea by Convergent Dental at 9,300 nm.
For the majority of dentists, the high-speed drill handpiece 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 Connective Tissue
Spatial distribution and concentration of chromophores in the epithelium and connective tissue are essential factors in understanding oral soft-tissue laser interaction.
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 absorption and reduced scattering coefficients sum.
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. Because 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 the epithelium is 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 shallow penetration depths (see Figure 6) regardless of epithelium pigmentation, which is essential 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 an estimated 10% blood  (assuming whole blood contains hemoglobin (and/or oxyhemoglobin) at an average 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 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. SuperPulse.
Tissue cooling from the blood flow during laser exposure is insignificant for short-term pulse durations of practical interest, mainly when tissue coagulation occurs underneath the ablated tissue. Instead, the cooling efficiency of the tissue irradiated by laser light is determined by the tissue’s 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 the 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 the tissue’s thermal diffusivity. λ ≈ 6.2-6.8 mW/cm ° C is the tissue’s heat conductivity; C ≈ 4.2 J/g ° C 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 called SuperPulse and is a must-have feature of any state-of-the-art soft-tissue surgical CO2 laser that minimizes the char and the coagulation depth.
The 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. A greater laser energy is needed 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 the 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 essential 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 of 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 much shorter absorption depths (see Figures 6 and 7), which makes these lasers preferable for soft-tissue ablative laser applications.
The 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 essential measure of coagulation and hemostasis efficiency. It is derived in  for laser pulses shorter than Thermal Relaxation Time (pulse duration t ≤ TR) and is presented in Figure 11. Coagulation depths for the 1,100-1,800 nm range were not calculated because of the scarcity of light scattering coefficient data for oral soft-tissue (which are essential 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 the ablation margin,s 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 that the coagulation depth in Figure 11 at 10,600 nm is about 50% thinner than at 9,300 nm, also observed in , , 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 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 of CO2 surgical lasers in the 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 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 scanned across the soft-tissue, the incision depth is proportional to average laser power. It is inversely proportional to the 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 average laser 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. 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 by 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 changing the control panel’s laser settings.
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 unsuitable for oral soft-tissue cutting, there is a different non-optical (non-radiant) tissue cutting technique 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 the charred glass fiber’s tip up to 900° C . Such a hot glass tip heats up the soft-tissues through heat conduction (i.e., heat diffusion) from (and through) the 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 of 250-1,000 µm .
Figure 16 reproduces the hot tip coagulation depths measurements  conducted under constant tip temperature (red circles) and under continuous power (blue circles) conditions. Also presented are the calculated  coagulation depths for constant tip temperature (red line) and for continuous power (blue line) conditions. As can be seen, the hot glass tip can 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), 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. It 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 a CHAR layer on the tip as the 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 the 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’s cladding materials at 500-900°C operating temperatures; and
- Sterility  of the tip can be compromised by non-sterile inks and cork used in the 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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