LA&HA - Journal of the Laser and Health Academy
Volume: 2020 | number: 1
ISSN (print):1855-9913 | ISSN (web): 1855-9921
Smooth-Surfacing of Soft Tissues Using Thermal Er:YAG Laser Pulse Sequences
Matjaz Lukac, Anze Zorman, Blaz Tasic, Nejc Lukac
Pages: S05
Ablative skin resurfacing by Er:YAG lasers has proven to be an effective and reproducible method for treating wrinkles [1-4]. In erbium laser procedures, it is the tissue’s water content, not its pigment that plays the role of an absorbing chromophore. The laser induced temperature elevation T is thus not limited to a particular pigment, such as melanin or hemoglobin, but to the superficially irradiated tissue layer with its thickness determined by the laser’s optical penetration depth () and the subsequent thermal diffusion [5, 6]. Very roughly, the elevated temperature lasts for the combined duration of the laser pulse and subsequent conductive cooling, which is for typical Er:YAG single laser pulse durations on the order of one millisecond.
The Er:YAG laser treatments consist of heating the superficial tissue up to the maximally achievable temperature defined by the ablation temperature Tabl where tissue ablation starts, as a result of micro-explosions of overheated tissue water within the elastic skin tissue [6]. Since the water contained within the confined solid tissue cannot expand freely, the ablation temperature is not at the boiling temperature of water under atmospheric pressure of about 100 0C but at a much higher temperature of Tabl ≈ 250 0C [6, 7].
Resurfacing with the Er:YAG laser has been of particular interest since it allows for the so-called “cold” ablation with minimal thermal damage below the ablation front [4, 8]. This may be somewhat surprising considering the high superficial temperatures encountered during ablation. Namely, the heating of tissue is accompanied by the chemical process of protein denaturation as a result of the cellular exposure to the increased temperature [9]. The tissue damage is typically calculated using the Arrhenius damage integral Ω calculated over the time of the thermal exposure [9-12], according to which the tissue injury grows exponentially with the elevated temperature T, and linearly with the time of exposure t. The tissue damage is often characterized by a critical (i.e., damage threshold) temperature (Tcrit), representing the temperature at which the concentration of the undamaged tissue is reduced by a factor of e. According to the Arrhenius model of skin damage developed by Henriques and Moritz [11, 12], based on measurements at longer exposures, the critical temperature for the exposure durations encountered during Er:YAG laser treatments would be around 70 0C, significantly below the ablation temperature.
Recently [7], it was pointed out that the above discrepancy can be explained by noting that during measurements performed at extremely short exposure times (t < 10 ms), the critical temperatures have been found to be significantly higher than what would be expected from the standard single process Arrhenius model [13 -16]. Using a VHS (Variable Heat Shock) model where the tissue thermal response was modeled with two interacting biochemical processes, defining the cell viability at very long and very short exposure times [16, 17], it was shown that for typical Er:YAG laser treatment parameters the critical temperature is above Tabl [7]. This was attributed to the uniquely short thermal exposure times, facilitated by the uniquely short optical penetration depth of the Er:YAG laser wavelength ( = 2.940 nm) of = 1-3 m in soft tissues. The resulting extremely large thermal gradient between the superficially optically heated tissue and the underlying tissue leads to a very short temperature decay time that combined with a typical Er:YAG laser pulse duration results in a total duration of the thermal exposure on the order of 1 millisecond.
It has also been suggested that the extremely short high-temperature, yet safe Er:YAG thermal pulses imposed on the epitelium involve an additional intense heat shocking regenerative mechanism that is complementary to the conventional deep thermal stimulation of fibroblasts [7]. This additional, superficial mechanism of action for regenerating epithelial and deeper lying connective tissues was proposed to be based on triggering stimulating signal transduction processes for transcription factor activation, gene expression and fibroblast growth, thus leading to new collagen and extracellular matrix formation [17-23].
While ablative laser resurfacing procedures have been found to be extremely effective, a major disadvantage is the erosion of large surfaces, which necessitates a recuperation period of 1 to 2 weeks. There are also potential risks of infections, scarring or hyper- and hypo-pigmentation [24-26]. For this reason, it has been proposed to utilize the unique superficial absorption characteristics of Er:YAG also for less invasive non-ablative treatments [3, 5, 27, 28]. As opposed to ablative procedures, the main mechanism of action of standard non-ablative procedures is based on selective deep thermal damage followed by new collagen formation. With the Er:YAG laser, however, the mechanism of action potentially involves also the proposed superficial heat shock triggering [7, 27, 28].
In non-ablative treatments, the depth of tissue’s thermal response is determined by the amount of heat which can be delivered to the tissue in a non-ablative manner. The maximal coagulation depth is thus limited by the maximal surface temperature Tmax = Tabl , and related maximal laser pulse fluence Fabl (in J/cm2), above which ablation starts [6]. The ablation threshold fluence can be increased by using longer laser pulses during which the delivered heat has more time to diffuse deeper into the tissue, effectively reducing surface temperature and thus increasing Fabl [6, 8]. However, since the existing Er:YAG laser technology limits single pulse durations to below several milliseconds, this limits the maximal coagulation depths which can be achieved with single pulses to maximally up to several tens of micrometers. For this reason, Er:YAG laser non-ablative procedures are typically carried out with longer-duration sequences of multiple laser pulses [29-31].
Many of the published studies of non-ablative, thermal treatments with Er:YAG laser pulse sequences have been made at relatively high sequence fluences, close to or slightly above the ablation threshold [29, 30-34]. This resulted in a significant damage to the epidermis, and in most cases also in subsequent removal of the damaged epidermis, making the treatments “delayed ablative”.
In this study, we explored a different modality of non-ablative Er:YAG treatments, a “smooth-surfacing”, where the cumulative fluence is set to be not only below the ablation threshold but also below the pain threshold for treatments without anesthesia, or at most with topical anesthesia. The technique involves delivering laser energy in 0.1-10 s long “SMOOTH” mode sequences, each consisting of several consecutive sub-ablative sub-millisecond Er:YAG laser pulses [3, 5, 27, 28]. During the SMOOTH mode sequence, the laser-generated heat is effectively “pumped” by means of heat diffusion away from the epithelia, several hundred microns deep into the connective tissue.
The temperature and tissue response characteristics during smooth-resurfacing were analyzed numerically for a broad range of treatment parameters. The tissue response was calculated using the two-process VHS model, in order to take into account the fast superficial temperature changes during individual laser pulses, and slower cumulative heat deposition during the over-all laser pulse sequence. Pain thresholds were measured during treatments with and without topical anesthesia, for different pulse sequence settings. Using the obtained pain threshold values, we evaluated the tissue response for a wide range of sub-surfacing laser parameters.
Smooth-Surfacing of Soft Tissues Using Thermal Er:YAG Laser Pulse Sequences