NON-ABLATIVE RESURFACING OF SOFT TISSUES

Abstract

The present disclosure relates to a method and an apparatus for generating a laser pulse sequence for application to a predetermined target tissue. Means are provided for setting a cumulative fluence F.sub.s of the laser pulse sequence. Means are provided for determining a duration of the laser pulse sequence as a function of the cumulative fluence F.sub.s such that the predetermined target tissue is heated to a final temperature that is within a predetermined range.

Claims

1. Apparatus for generating a laser pulse sequence for application to a predetermined target tissue, comprising: means for setting a cumulative fluence F.sub.s of the laser pulse sequence; means for determining a duration of the laser pulse sequence as a function of the cumulative fluence F.sub.s such that the predetermined target tissue is heated to a final temperature that is within a predetermined range.

2. Apparatus according to claim 1, wherein the means for determining is adapted to determine the duration such that the final temperature is below a predetermined pain threshold temperature specific for the predetermined target tissue.

3. Apparatus according to claim 1, wherein the means for determining is adapted to determine the duration based on a temperature model for the final temperature, wherein the temperature model is a function of the duration and the cumulative fluence F.sub.s.

4. Apparatus according to claim 1, wherein the laser pulse sequence has a predetermined number N of sequential pulses, and the means for determining is adapted to determine the duration such that the final temperature at time t.sub.s=N×Δt after application of the first pulse of the pulse sequence is within the predetermined range, wherein Δt represents an average separation time between subsequent pulses of the pulse sequence.

5. Apparatus according to claim 1, further comprising: means for setting at least one parameter indicative of the predetermined target tissue; wherein the means for determining are adapted to determine the duration based on the at least one parameter.

6. Apparatus according to claim 1, further comprising: means for generating the laser pulse sequence with the determined duration; and/or means for receiving a setting of the duration of the laser pulse sequence and means for issuing a warning if the setting of the duration deviates from the determined duration by a predetermined threshold.

7. Apparatus for generating a laser pulse sequence for application to a predetermined target tissue, comprising: means for setting a duration t.sub.s of the laser pulse sequence; means for determining a cumulative fluence F.sub.s of the laser pulse sequence as a function of the duration such that the predetermined target tissue is heated to a final temperature that is within a predetermined range.

8. Apparatus according to claim 1, further comprising one or more user interfaces for receiving: a selection input representative of the duration and/or the cumulative fluence of the laser pulse sequence; and/or a selection input representative of a tissue type, and/or a selection input representative of a pixelization of the laser pulse sequence.

9. Apparatus according to claim 1, wherein the apparatus comprises an Er:YAG laser source for generating the laser pulse sequence and wherein the means for determining is adapted to determine the duration t.sub.s and/or the cumulative fluence F.sub.s of the pulse sequence such that (1−x.sub.1)ΔT.sub.p/(A.sub.s×t.sub.s.sup.−Ks)<F.sub.s<(1+x.sub.2)*ΔT.sub.p/(A.sub.s×t.sub.s.sup.−Ks), wherein ΔT.sub.p is 22° C., x.sub.1=0.5, x.sub.2=0.5, K.sub.s=0.43; wherein A.sub.s is a predetermined parameter.

10. Apparatus according to claim 8, wherein the apparatus comprises an Er:YAG laser source for generating the laser pulse sequence and wherein the means for determining is adapted to determine the duration t.sub.s and/or the cumulative fluence Fs of the pulse sequence such that (1−x1) ΔT.sub.p/(A.sub.s×t.sub.s−Ks)<Fs<(1+x2)*ΔTp/(A.sub.s×t.sub.s−Ks), wherein ΔTp is 22° C., x1=0.5, x2=0.5, Ks=0.43; wherein As is a predetermined parameter, and wherein: A.sub.s=84 (° C. cm.sup.2)/J, if selection input according to claim 8 indicates human skin as tissue type and no pixelization of the laser pulse sequence; and/or A.sub.s=69 (° C. cm.sup.2)/J, if selection input according to claim 8 indicate human mucosa as tissue type and no pixelization of the laser pulse sequence; and/or A.sub.s=28 (° C. cm.sup.2)/J, if selection input according to claim 8 indicate human skin as tissue type and pixelization of the laser pulse sequence; and/or A.sub.s=18 (° C. cm.sup.2)/J, if selection input according to claim 8 indicate human mucosa as tissue type and pixelization of the laser pulse sequence.

11. Apparatus according to claim 1, the apparatus further comprising: at least two predetermined modes for generating the laser pulse sequence between which can be switched; pre-calibrated means for generating (control unit) the laser pulses of the laser pulse sequence with an average pulse fluence F.sub.ave=(F.sub.1+ . . . +F.sub.N)/N of the pulse sequence, such that a mean value of maximum temperature increases of the predetermined target tissue caused by each of the individual pulses of the pulse sequence is within one of the following three ranges in the first mode and within a disparate one of the following three ranges in the second mode: 5° C. to 30° C.; 30° C. to 65° C.; 65° C. to 135° C.

12. Apparatus according to claim 1, further adapted such that: the pulse sequence comprises between 6 and 40 pulses, preferably between 10 and 40, most preferably between 12 and 30; wherein the apparatus is adapted to apply the pulse sequence in a non-pixelated manner, and the cumulative fluence F.sub.s=F.sub.1+ . . . +F.sub.N of the pulse sequence is between 2.5 and 8 J/cm.sup.2, preferably between 2.5 and 6 J/cm.sup.2, most preferably between 2.5 and 5 J/cm.sup.2; or wherein the apparatus is adapted to apply the pulse sequence in a pixelated manner, and the cumulative fluence F.sub.s=F.sub.1+ . . . +F.sub.N of the pulse sequence is between 7.5 and 24 J/cm.sup.2, preferably between 7.5 and 18 J/cm.sup.2, most preferably between 7.5 and 15 J/cm.sup.2.

13. Apparatus according to claim 1, adapted to apply a first laser pulse sequence to a first position on the target tissue and a second laser pulse sequence to a second position on the target tissue; wherein the apparatus further comprises a scanner; wherein the scanner is adapted to automatically apply a first pulse of the second pulse sequence to the second position at a period of time after a first pulse of the first sequence has been applied to the first position, preferably the period of time being the inverse of a maximum pulse generation frequency, but before a last pulse of the first sequence has been applied to the first position.

14. Method for generating a laser pulse sequence for application to a predetermined target tissue, comprising: setting a cumulative fluence F.sub.s of the laser pulse sequence; determining a duration of the laser pulse sequence as a function of the cumulative fluence such that the predetermined target tissue is heated to a final temperature that is within a predetermined range.

15. Method for generating a laser pulse sequence for application to a predetermined target tissue, comprising: setting a duration t of the laser pulse sequence; determining a cumulative fluence F.sub.s of the laser pulse sequence as a function of the duration such that the predetermined target tissue is heated to a final temperature that is within a predetermined range.

Description

4. BRIEF DESCRIPTION OF THE FIGURES

[0086] FIGS. 1A, 1B: Exemplary temporal and spatial temperature profiles during and following a laser pulse sequence (with N=6 and t.sub.sep=50 ms);

[0087] FIG. 2: Shows the critical temperature as a function of the exposure time, the function representing a combined effect of two limiting Arrhenius' processes, defining cell viability at extremely long and short exposure times.

[0088] FIG. 3: Calculated dependence of the single-pulse ablation threshold fluence (F.sub.ab1) and single pulse temperature slope (η.sub.p=ΔT.sub.max/F.sub.p) on the Er:YAG single-pulse duration (t.sub.p);

[0089] FIG. 4: Measurements and simulated dependence of the sequence temperature slope (η.sub.s=ΔT.sub.s/F.sub.s) on the sequence duration t.sub.s for full-beam and patterned beam treatment of skin and mucosa;

[0090] FIGS. 5A, B: Show cumulative pain threshold fluences F.sub.pain(J/cm.sup.2) for various pulse sequence durations and patients;

[0091] FIG. 6a,b: Show the dependence of the measured pain threshold fluence F.sub.pain and pain threshold temperature increase on the sequence duration t.sub.s;

[0092] FIG. 7: Shows the calculated long-exposure coagulation depths z.sub.c (μm) as a function of the sequence duration t.sub.s (ms);

[0093] FIGS. 8A, B: Show the calculated short-exposure superficial damage (Ω.sub.s) as a function of the sequence duration t.sub.s (FIG. 10a) or number of pulses N (FIG. 10B);

[0094] FIG. 9: Shows the spatially patterned (i.e., pixelated) delivery of laser pulses

[0095] FIGS. 10A, B: Shows the evolution of pulse peak temperatures T.sub.max-i for each pulse and of the heat shock triggering amplitude A.sub.trig after each set of 6 pulses during a heat shocking triggering;

[0096] FIG. 11: Exemplary embodiment of an apparatus according to the present invention.

5. DETAILED DESCRIPTION

[0097] Ablative skin resurfacing using CO.sub.2 and Er:YAG lasers has proven to be an effective and reproducible method for treating wrinkles. Of particular interest has been resurfacing with the Er:YAG laser wavelength since it allows for the so-called “cold” ablation with minimal thermal damage below the ablation front. This unique erbium laser characteristic has been attributed to the extremely short optical penetration depth (δ) in human soft tissues of only several microns, the shortest penetration of all non-ultraviolet lasers. This is because the Er:YAG laser wavelength is positioned at the highest far-infrared water-absorption peak of λ=2940 nm. In erbium laser procedures, it is therefore the tissue's water content, not its pigment, that plays the role of an absorbing chromophore. The laser-induced temperature elevation ΔT 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 extremely short optical penetration depth.

[0098] The laser-induced temperature elevation t.sub.s accompanied by the chemical process of protein denaturation as a result of the cellular exposure to the increased temperature.

[0099] The tissue damage is typically calculated using the Arrhenius damage integral Ω calculated over the time of the thermal exposure (t.sub.exp) and the related critical temperature (T.sub.crit), representing the temperature at which the concentration of the undamaged tissue is reduced by the Euler's number e.

[0100] According to the Arrhenius damage integral, the tissue injury grows exponentially with the elevated temperature T, and linearly with the time of exposure t.sub.exp. During pulsed laser procedures, the duration of the heat shock, i.e. of the thermal exposure, t.sub.exp≈t.sub.p+t.sub.d, is not determined only by the duration of the laser pulse t.sub.p, but typically much more by the temperature decay time td required for the irradiated tissue layer to cool back down to the initial temperature. For the superficially absorbed lasers, the temperature decay time t.sub.d depends strongly on the optical penetration depth according to t.sub.d≈δ.sup.2/D, where D is the tissue's thermal diffusivity [1: Majaron B, Sustercic D, Lukac M, Skaleric U, Funduk N. Heat diffusion and debris screening in Er:YAG laser ablation of hard biological tissues. Appl Phys B 1998; 66:479-487, 7; 2: Majaron B, Plestenjak P, Lukac M. Thermo-mechanical laser ablation of soft biological tissue: modeling the micro-explosions. Appl. Phys. B 1999; 69, 71-80]. Short thermal exposures can therefore be achieved only by short-pulsed lasers with strong absorption in tissues. Here, Er:YAG is at a significant advantage due to its highest absorption in tissue water, enabling thermal exposure times below one millisecond [3: Lukac M, Lozar A, Perhavec T, Bajd F. Variable heat shock response model for medical laser procedures, Lasers Med Sci. August 2019; 34(6):1147-1158]. For example, the exposure time may be approximated by

t.sub.exp≈t.sub.ed+t.sub.d

[0101] The importance of td being very short can be better understood by considering that during tissue resurfacing, the superficial tissue is typically heated up to the ablation temperature T.sub.abl at which the tissue ablation starts as a result of micro-explosions of overheated tissue water within the elastic soft tissue. 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° C., but at a much higher temperature of T.sub.ablϰ° C. [2, 3]. According to the standard Arrhenius model of skin damage, the critical temperature would be much lower, around 55-65° C. However, for extremely short exposure times, the critical temperatures are significantly higher than what would be expected from the standard single process Arrhenius model. For thermal exposure times attainable by Er:YAG resurfacing, the critical temperature is above 250° C. This is shown in FIG. 2 which shows the critical temperature as a function of the exposure time [3] (VHS: variable heat shock).

[0102] 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. For this reason, it has been proposed to utilize the unique superficial absorption characteristics of Er:YAG also for less invasive non-ablative treatments. As opposed to ablative procedures, the main mechanism of action of non-ablative procedures is based on selective thermal damage followed by new collagen formation. The depth of the tissue's thermal response, i.e., of the tissue coagulation, is determined by the amount of heat that can be delivered to the tissue in a non-ablative manner. Since in the absence of thermal diffusion the ablation threshold fluence F.sub.abl (in J/cm.sup.2), is inversely proportional to δ, the heat energy that can be delivered into the tissue by a single Er:YAG laser pulse is relatively small. This applies especially since the existing Er:YAG laser technology limits single pulse durations to below several milliseconds, limiting the time available for conductive superficial cooling of the tissue during the laser pulse. It is for this reason that non-ablative Er:YAG treatments have been performed by repetitive stacking of sub-threshold Er:YAG laser pulses, resulting in a higher cumulatively delivered sequence fluence, F.sub.i.

[0103] Initial studies of non-ablative thermal treatments with repetitive stacking of Er:YAG laser pulses were made at cumulative fluences (F.sub.s) close to the ablation threshold. This resulted in a significant damage to the epidermis, leading to subsequent peeling of the damaged epidermis, making the treatments “delayed ablative”. Therefore, with this type of non-ablative resurfacing, the epithelium is in reality damaged, but not completely removed during the procedure, and acts as a wound dressing. For this reason, this type of sub-ablative (or minimally ablative) resurfacing modality can be called also a “sub-resurfacing” modality.

[0104] This invention describes a different resurfacing modality, further referred to also as the smooth-resurfacing modality where the applied fluences are not only below the ablation threshold but also below the patient's pain threshold. Pain threshold fluences were measured for abdominal skin with and without topical anesthesia and compared with measured superficial tissue temperature evolution during smooth-resurfacing of the skin and oral mucosa. The obtained temperature data and pain thresholds were then used to study the characteristics of the short- and long-exposure's tissue response believed to be respectively involved in the indirect and direct soft-tissue regeneration mechanisms.

[0105] With smooth-resurfacing, indirect regeneration mechanisms in addition to the direct heat injury to the deeper lying connective tissues play a role in regeneration and remodeling of the treated tissue. This indirect triggering of tissue regeneration through short-exposure intense heat shocking of epithelia is based on stimulating signal transduction processes for transcription factor activation, gene expression and fibroblast growth, leading to new collagen and extracellular matrix formation.

a) PHYSICAL MODEL OF RESURFACING

[0106] A numerical model may be applied for the physical process of sub- and smooth-resurfacing of soft tissues, such as skin and mucosa, based on that developed to study thermo-mechanical ablation with mid-IR lasers. The details of that model are described in [4: Majaron B, Plestenjak P, Lukac M. Thermo-mechanical laser ablation of soft biological tissue: modeling the micro-explosions. Appl. Phys. B 1999; 69, 71-80].

[0107] Based on the model, we assume that a single wavelength (λ) pulsed laser radiation is delivered to the surface of the treated tissue with a pulse fluence F.sub.p (in J/cm.sup.2). The tissue is modeled as a water-containing homogeneous media characterized by a single absorption coefficient of k=1/δ for the delivered laser wavelength λ. For simplicity, a square-shaped laser pulse with duration t.sub.p is assumed. If the focus is on the Er:YAG laser wavelength with a short penetration depth, effects of scattering of the laser light within the tissue can be excluded. Similarly, it may be assumed that that the laser spot size is much larger than the penetration depth δ. Therefore, the diffusion of dissipated heat may be treated in one dimension using a finite-difference scheme. In all calculations, we use the physical parameters of the irradiated media as published in [2].

[0108] To elucidate at what individual pulse fluence F.sub.p and total sequence fluence F.sub.s the laser-tissue interaction starts being ablative, the model also includes the microscopic physical model of the ablative micro-explosion process, which combines the thermodynamic behavior of tissue water with the elastic response of the solid tissue components.

[0109] The model was applied to calculate temporal and spatial temperature profiles for single pulses and as well for pulse sequences, each consisting of N consecutive Er:YAG laser pulses with individual F.sub.p and cumulative F.sub.s fluences, separated by a pulse separation time t.sub.sep. The effective duration of the modeled pulse sequences was defined by the sequence duration t.sub.s=N×t.sub.sep.

[0110] FIG. 1A shows a typical temporal profile of the tissue surface temperature, T (° C.) during a pulse sequence over the course of time in seconds, t(s). It is noted that, unless indicated otherwise, the tissue temperatures as defined herein pertain to tissue surface temperatures. The pulse sequence comprises N=6 pulses at a separation time of t.sub.sep=50 ms, and comprises a duration indicated as t.sub.s, wherein t.sub.s=N×t.sub.sep. The pulse sequence results in N high temperature peaks (designated by the reference sign T.sub.max-i) that rapidly relax deeper into the tissue by fast thermal diffusion driven by the large temperature gradient over the short optical absorption length. Due to the fast thermal diffusion from the heated ≈1-3 μm thin superficial tissue layer, the duration of the thermal exposure (t.sub.exp) to high temperature peaks T.sub.max-i is extremely short (it can be as short as t.sub.exp<1 ms but can also be longer). During the pulse sequence, the superficially laser-generated heat is thus being “pumped” by diffusion away from the epithelia, up to several hundred microns deep into the connective tissue (see FIG. 1B). The final, longer persisting surface temperature is in FIG. 1A represented by the sequence temperature T.sub.s (at t=t.sub.s=N×t.sub.sep).

[0111] FIG. 1B shows the resulting spatial profile of the tissue temperature, T (° C.) within the tissue depth, z (μm) by the end of the sequence, i.e. at time t.sub.s=N×t.sub.sep.

b) CHEMICAL MODEL OF NON-ABLATIVE RESURFACING

[0112] Typically, tissue damage response is calculated using the Arrhenius damage integral Ω calculated over the time (t.sub.exp) of the temporally square-shaped thermal exposure to elevated temperature:

Ω=A exp(−E/RT)×t.sub.exp.  (1)

[0113] Here, A is the frequency factor, i.e. the damage rate (in s.sup.−1), E is the activation energy [in J/kmol], and R is the gas constant (R=8.31 10.sup.3 J/kmol K). The damage integral defines the probability (P) for tissue damage response according to:

P=1−exp(−Ω).  (2)

[0114] Similarly, the critical (i.e., damage threshold) temperature (T.sub.crit), depends on the thermal exposure time as:

T.sub.crit=E/(R ln(A t.sub.exp)).  (3)

[0115] During Er:YAG laser pulsing, the superficial tissue's thermal exposure transitions from intense, extremely short periods of exposure to peak temperatures T.sub.max-i (see FIG. 1A), to long-duration periods of exposure to moderate temperatures (T≤T.sub.s). For this reason, a two-process response model (hereinafter referred to also as the “VHS” model) can be used in order to evaluate the tissue damage Ω from the calculated temporal and spatial temperature profiles. The VHS model assumes that the effective damage integral Ω can be calculated for any exposure from the combined effect of the damage integrals Ω.sub.long(t.sub.esp) and Ω.sub.short(t.sub.exp) belonging respectively to the long and short Arrhenius processes (See FIG. 2), as follows:

(1/Ω).sup.p=(1/Ω.sub.long).sup.p+(1/Ω.sub.short).sup.P.  (4)

where p≈0.15 is the transition coefficient that determines the transition between the two limiting biochemical processes. The details of the VHS model are described in [3].

[0116] The long and short processes are characterized by the following Arrhenius parameters: A.sub.long=4.7×10.sup.89 s.sup.−1 and E.sub.long=5.67×10.sup.7 Jkmol.sup.−1 for the long exposure process, and A.sub.short=1.45×10.sup.4 s.sup.−1 and E.sub.short=1.03×10.sup.7 Jkmol.sup.−1 for the short exposure process. Eq. (4) defines an implicit function of the critical temperature, i.e. at which Ω=1, depending on the exposure time, and is depicted in FIG. 2. For example, critical temperatures of approximately 70° C., 80° C., 90° C., 100° C., 110° C., 120° C., 160° C., 180° C., and 250° C. are obtained for thermal exposure times t.sub.exp=150 ms, 25 ms, 20 ms, 6 ms, 4.5 ms, 3.5 ms, 2 ms, 1.5 ms and 0.8 ms.

[0117] In some examples, the exposure time may be approximated by t.sub.exp=t.sub.ed+(1/D)(δ+✓(2 D t.sub.ed)).sup.2, wherein D is the thermal diffusivity of the treated tissue which may be, for the aspects disclosed herein, approximated by =0.1 mm.sup.2/s. t.sub.ed denotes the time during which the second half of the pulse energy is delivered (e.g. for temporally symmetric pulses t.sub.ed=t.sub.p/2). δ is the penetration depth in the target tissue which may depend on the wavelength of the pulse (sequence). For example, for Er:YAG lasers with a wavelength λ 2,940 nm, δ≈1 μm. For CO.sub.2 lasers with wavelength λ≈10,640 nm, δ≈15 μm. For example, for Er,Cr:YSGG lasers with λ≈2,780 nm, δ≈3 μm.

[0118] The damage integral extending deeper into the tissue, characterized predominantly by the long-pulse exposure process, was calculated by integrating the damage over temperature instead of over time, using the algorithm developed for calculating tissue damage for temporally non-square-shaped thermal exposure pulses [3]. The tissue coagulation depth (z.sub.c) was defined as the tissue depth below which the calculated cell injury is smaller than Ω=0.5.

[0119] It has been shown that the superficial tissue damage caused by a single Er:YAG laser pulse is governed mainly by the short-pulse Arrhenius process, and extends about z.sub.1/2≈10 μm deep into the tissue. For calculating the cumulative short-pulse Arrhenius process damage to this superficial tissue layer, following a series of i=1 . . . N, intense short-duration thermal exposures to T.sub.max-i, a probability-summation model was used similar to that in [5: Menendez A R, Cheney F E, Zuclich J A, Crump P. Probability-summation model of multiple laser-exposure effects. Health Phys 1993; 65:523-528]. In this model, it is assumed that the response to each pulse of a multiple-pulse exposure is independent of the response to other pulses; that is, previous pulses do not “sensitize” the tissue to subsequent pulses. The probability P.sub.i of thermal damage caused by each pulse, i=1 . . . N is then calculated from

P.sub.i=1−exp(−Ω.sub.i),  (5)

where Ω.sub.i was calculated using Eqs. 1 and 4, assuming an exposure to a constant temperature T.sub.max-i for an effective duration t.sub.eff. The effective duration was approximated by t.sub.eff≈0.25 ms, based on results of simulations in [3], assuming t.sub.p=0.3 ms, and represented the duration of an imaginary rectangular temperature pulse of a constant average temperature T.sub.maxi-, which produces approximately the same amount of damage as the actual “triangularly” shaped temperature pulse. The cumulative “sequence” probability P.sub.s (N) of inducing a thermal damage to the tissue surface during N pulses was then calculated using:

P.sub.s(N)=1−(1−P.sub.1)(1−P.sub.2) . . . (1−P.sub.N).  (6)

[0120] This gives an approximate cumulative damage integral averaged over the z.sub.1/2≈10 μm thick thermally affected superficial layer as:

Ω(N)=0.5×Ln(1/(1−P.sub.s(N)).  (7)

[0121] It is to be noted that although T.sub.max-i temperatures were calculated taking into account the gradual temperature increase during the pulse sequence, the above damage does not include the longer exposure tissue response resulting from this gradual temperature build-up.

c) HEAT-PAIN THRESHOLD MEASUREMENT

[0122] In order to determine the heat-pain threshold (HPT) for different smooth-resurfacing conditions, 15 patients participated in the experiment on a voluntary basis. Thermal stimuli were delivered using an Er:YAG laser (Dynamis SP, manufactured by Fotona d.o.o.; operating in a V-SMOOTH pulse mode). Sequences with pulse separation times t.sub.sep=25, 100, and 125 ms were applied, with pulse numbers up to N=36. The laser pulse duration was equal to t.sub.p=0.3 ms. The laser energy was delivered to the abdominal skin area using a T-Runner scanning handpiece (manufactured by Fotona d.o.o.) with a single full-beam spot size of 2r=9 mm, scanned over a scanning area consisting of 3×3 spots. Heat-pain thresholds were obtained for treatments without and with topical anesthesia (application of EMLA 25 minutes before the test).

[0123] The cumulative laser fluence was gradually increased for each successive scan in steps of 0.2 J/cm.sup.2 from a low level until the patient characterized the treatment pain as intolerable. The pain threshold can be understood as the fluence just below the value at which the patient reported the discomfort during the treatment to be unacceptable.

d) TISSUE SURFACE TEMPERATURE MEASUREMENT

[0124] Measurements of the tissue surface temperature evolution as a result of Er:YAG laser irradiation were made on the abdominal skin and on mucosa (intraorally on the cheeks) using the Dynamis SP laser system (manufactured by Fotona d.o.o.) operating in SP pulse duration mode (t.sub.p=0.3 ms). In what follows it will be assumed that the data for the intra-oral tissue represents all mucous tissues, including the vaginal tissue.

[0125] For measuring the temporal temperature profile during longer-duration pulse stacking, a thermal camera (ThermaCAM P45, manufactured by FUR Systems, USA) with a frame rate of 50 Hz was used. Two laser handpieces were used: i) a full-beam handpiece (Fotona R11) and ii) a pixelated beam handpiece (Fotona PS03), with both handpieces set to a 7 mm spotsize. The Fotona PS03 handpiece is equipped with a pixel screen, resulting in the overall laser spot having a pixelated (also referred to as patterned or dotted) internal beam structure, with the centers of the individual circular beam dots of diameter of 2r=0.85 mm being separated by approximately 2 mm center to center. The motivation behind the design of the pixelated PS03 handpiece is to make the treatment less invasive by reducing the treatment area to isolated beam islands.

e) RESULTS

aa) Single Pulse Exposure

[0126] As can be seen from FIG. 1A, each individual temperature pulse (i) consists of the temperature ramp-up heating phase during which the temperature reaches its maximal value (T.sub.max-i), and (ii) of the cooling phase during which the temperature returns back to its initial temperature (T.sub.oi). The heating phase lasts for approximately the duration of the laser intensity pulse (t.sub.p), while the cooling phase is determined predominantly by the rate of the heat flow away from the heated superficial tissue.

[0127] For fluences below the ablation threshold, the peak tissue temperature T.sub.max grows linearly with fluence at the single pulse temperature slope η.sub.p=ΔT.sub.max/F.sub.p, where ΔT.sub.max=T.sub.max−T.sub.o. The peak temperature increases with laser pulse fluence (F.sub.p) until the ablation threshold is reached, at which point the peak temperature reaches the ablation (“boiling”) temperature F.sub.p=F.sub.abl.

[0128] FIG. 2 presents the influence of the single-pulse duration t.sub.p (ms) on the temperature slope η.sub.p (° C. cm.sup.2/J) and on the ablation threshold fluence F.sub.abl (J/cm.sup.2). The initial temperature was taken to be equal to T.sub.o=35° C. The fits to the data points shown in FIG. 2 represent the dependence of the single-pulse temperature slope η.sub.p (in ° C. cm.sup.2/J), and of the single-pulse ablation threshold fluence F.sub.abl (in J/cm.sup.2) on the single full-beam laser pulse duration t.sub.p (in ms) based on the following equations:

η.sub.p≈A.sub.p×t.sub.p.sup.−Kp,  (8)

and

F.sub.abl≈A.sub.F×t.sub.p.sup.Kp,  (9)

where K.sub.p=⅓, and the full-beam values for skin are A.sub.p=173 and A.sub.F=1.28. The corresponding estimated full-beam values for mucosa (such as, for example, vaginal and oral mucous tissue) are A.sub.p=144 and A.sub.F=1.5. The corresponding values for the patterned beam are A.sub.p=81 and A.sub.F=2.7 for skin, and A.sub.p=67 and A.sub.F=3.3 for mucosa.

[0129] For fluences above F.sub.abl, the peak temperature remains fixed at T.sub.abl by means of micro-explosions, similarly to the case of boiling water that keeps its temperature at about 100° C. regardless of the heating power. Using our numerical model, the soft-tissue ablation temperature is calculated to be T.sub.abl=256±10° C., regardless of pulse duration. It is to Be noted that the relationship between the parameters A.sub.p and A.sub.F in Eqs. 8 and 9 above is then according to A.sub.F=(256° C.−35° C.)/A.sub.p.

bb) Multiple Pulse Exposure

[0130] Our simulations show that for fluences below the ablation threshold, the sequence temperature increase ΔT.sub.s=T.sub.s−T.sub.o, is linearly dependent on F.sub.s:

ΔT.sub.s≈η.sub.s×F.sub.s,  (10a)

where η.sub.s=ΔT.sub.s/F.sub.s. is the sequence temperature slope for a particular set of sequence parameters. While the dependence of the sequence temperature slope on the single-pulse duration (t.sub.p) or on the number of pulses Nis relatively small (and may be neglected), the dependence of the slope η.sub.s (in units of ° C. cm.sup.2/J) on the sequence duration t.sub.s (in msec), as obtained by fitting the numerical results to a power function, was found to be well described by:

η.sub.s=A.sub.st.sub.s.sup.Ks.  (10)

[0131] Here, the coefficients for the full beam treatments of the skin are A.sub.s=84 and K.sub.s=−0.43. Using these parameters, the numerically predicted dependence of η.sub.s on t.sub.s is in FIG. 4 represented by the upper full line. As can be seen from FIG. 4, the simulated dependence is in a good agreement with the slopes as measured for the full beam treatment of the abdominal skin (depicted by full circles).

[0132] Additionally, FIG. 4 shows the measured slopes for the full beam treatment of the mucous tissue, together with the fit of the experimental data to Eq. 10 with the coefficients K.sub.s=−0.43, and A.sub.s=69 (represented by the lower full line). Similarly, dotted lines represent fits of the experimental data to Eq. 10 for a patterned handpiece (i.e., Fotona PS03), with the coefficients K.sub.s=−0.43, and A.sub.s=28 for the patterned beam treatment of skin and A.sub.s=18 for the patterned beam treatment of mucosa. The difference between coefficients for skin and mucosa is attributed to the fact that the moist mucous tissue requires slightly more energy than the dry skin to be heated up. The lower coefficients for the patterned handpiece as compared with coefficients for the full beam handpiece are attributed to the radial heat diffusion away from the micro spots, in addition to the heat diffusion deeper into the tissue.

[0133] The obtained single pulse and sequence temperature slopes are used in further analysis to calculate the sequence ablation thresholds (F.sub.thr) as follows. The sequence ablation threshold is first reached when the maximal temperature (T.sub.max-N) of the last pulse (i=N) in the sequence with duration t.sub.s reaches the ablation temperature T.sub.abl. Assuming that all laser pulses (i=1 . . . N) in the sequence have the same fluence F.sub.p, then ablation starts when ΔT.sub.s−N−1(N−1)+η.sub.pF.sub.p=ΔT.sub.abl=256° C.−T.sub.o=221° C., where ΔT.sub.s−N−1 is the temperature elevation for the pulse sequence with N−1 pulses, and correspondingly with the pulse sequence duration equal to t.sub.s (N−1)/N. The ablation threshold fluence for any pulse sequence F.sub.thr=N×F.sub.p (in J/cm.sup.2) can then be calculated from:

F.sub.thr=NΔTabl(η.sub.P(t.sub.p)+(N−1)η.sub.s(t.sub.s(N−1)/N)).sup.−1  (11)

[0134] Analysis of Eq. 11 shows that the sequence ablation threshold depends more strongly on N and t.sub.p, and less on t.sub.s, especially for longer t.sub.s.

cc) Pain Threshold Fluences

[0135] FIG. 5A shows cumulative pain threshold fluences F.sub.pain (J/cm.sup.2), as obtained for treatments without topical anesthesia on 15 patients for four pulse sequence settings: i) N=6, t.sub.s=150 ms; ii) N=30, t.sub.s=3750 ms; iii) N=36, t.sub.s=3600 ms; and iv) N=36, t.sub.s=4500 ms. As can be seen from FIG. 6A, while pain thresholds vary from patient to patient, the pain threshold fluence is generally higher for longer sequence duration t.sub.s.

[0136] FIG. 5B shows pain threshold fluences F.sub.pain (J/cm.sup.2), as obtained for treatments with (closed symbols) and without (open symbols) topical anesthesia for sequence durations t.sub.s=150 ms and 4500 ms. Also FIG. 5B confirms the difference in pain threshold fluences for two sequence durations (150 ms and 4500 ms), and also shows their dependence on whether topical anesthesia (EMLA cream) was used or not.

[0137] Assuming that pain is associated mainly with the overall sequence temperature T.sub.s, and not with short-duration maximal temperatures T.sub.max-i, then the pain threshold temperatures (T.sub.p) can be obtained by calculating sequence temperatures at the measured pain fluences F.sub.pain, using Eq. 10.

[0138] The resulting pain threshold elevation ΔT.sub.p=T.sub.p−T.sub.o (° C.) obtained for treatments without anesthesia, for sequence times t.sub.s=150, 3600, 3750 and 4500 ms is on average ΔT.sub.p=12.7 OC. Average pain threshold temperatures for treatments without (averaged over 150, 3600, 3750 and 4500 ms data) and with topical anesthesia (averaged over 150 and 4500 ms data) were ΔT.sub.p=12.7° C., and with anesthesia, ΔT.sub.p-EMLA=16.1° C.

[0139] FIGS. 6A and 6B show the dependence of the measured pain threshold fluence F.sub.pain on the sequence duration t.sub.s, averaged over the 15 patients' data (FIG. 6A), and of the corresponding calculated pain threshold temperature elevations (FIG. 6B), with (closed symbols) and without (open symbols) topical anesthesia. The lines in FIG. 6A represent calculated pain thresholds according to Eq. 10, with ΔT.sub.p=12.7° C. without (dashed line), and ΔT.sub.p-EMLA=16.1° C. with topical anesthesia (full line). The fit to the data points is better than R.sup.2=0.98. As can be seen from FIG. 6B, the calculated pain threshold temperature represented by the sequence temperature is found to be approximately constant over a large range of sequence durations. On average, the pain threshold temperature elevation is equal to ΔT.sub.p=12.7±2.0° C. for treatments without anesthesia, and to ΔT.sub.p-EMLA=16.1±2° C. for treatments with topical anesthesia.

[0140] It is to be noted that the above pain thresholds were measured on the abdominal skin. Typically, cutaneous heat pain threshold temperatures (HPTT) for several-seconds-long exposures are for different body areas in the range of 41-52° C., in rough agreement with the pain threshold temperature for smooth-resurfacing of abdominal skin of T.sub.p 48° C., as found in our experiments. In agreement with the VHS model and the assumption that the pain threshold is related to the risk of irreversible damage, the HPTT has been observed to increase towards shorter exposures, and was reported to be equal to HPTT≈58° C. for t.sub.exp≈0.3 s and to HPTT≈75° C. for t.sub.exp≈0.05 s. Similar heat pain thresholds have also been obtained for oral mucosa, with HPTT≈48° C. for long exposures, and HPTT≈65-70° C. for approximately 0.1 s long exposures.

dd) Deep Tissue Response

[0141] FIG. 7 shows calculated coagulation depths z.sub.c (μm) as a function of the sequence duration t.sub.s (ms), for different sequence durations when smooth-resurfacing is performed, i.e. when the sequence fluence is for each sequence duration adjusted to be just below the pain threshold F.sub.pain or F.sub.pain-EMLA (abbreviated as F.sub.p or F.sub.p-EMLA in FIG. 7), for treatments without or with topical anesthesia, respectively (t.sub.sep=25 ms (diamonds), 75 ms (squares) and 125 ms (circles)).

[0142] The pain threshold fluences were calculated using Eq. 10 for ΔT.sub.p=12.7° C. (without anesthesia) and 16.1° C. (with topical anesthesia) according to F.sub.pain=ΔT.sub.p/ηs. As can be concluded from FIG. 7 (full lines), smooth-resurfacing can result in significant coagulation depths, providing that appropriately long sequence durations is are used, which allow sufficiently high yet still painless sequence fluences F.sub.s. For example, when using topical anesthesia, the pain threshold fluence at t.sub.s=10 S is equal to F.sub.pain-EMLA≈10 J/cm.sup.2. Also, based on our finding, the coagulation depths at the “smooth” fluences, i.e., at the fluences just below the pain threshold, do not depend on the number of pulses N nor on the pulse separation time t.sub.sep, but are determined solely by the sequence duration t.sub.s. Hence, methods and apparatus may be provided based on this insight, that take this into account and provide means for user-based and/or (programmed) to automatic settings accordingly, e.g. applying a sequence with a predetermined pulse sequence duration to achieve a predetermined coagulation depth. For example, a user interface may be provided such that an operator of a corresponding device may select a certain coagulation depth and/or a skin type (e.g. anesthetized or non-anesthetized skin). The device may then automatically apply a laser pulse sequence with a corresponding required cumulative fluence and corresponding pulse duration. In other examples, a user may be displayed (information on the expected) coagulation depth for a setting.

ee) Superficial Tissue Response

[0143] FIGS. 8A and 8B show the calculated short-exposure superficial damage (Ω.sub.s) for smooth-resurfacing (black and open symbols connected by full lines) applications, as a function of the sequence duration t.sub.s (FIG. 8A) or number of pulses N (FIG. 8B), as calculated using Eq. 7. Black and open symbols represent smooth-resurfacing with and without topical anesthesia, respectively. The calculated damage integrals are for pulse separation times t.sub.sep of 25 ms (diamonds), 85 ms (squares) and 125 ms (circles).

[0144] As can be seen from FIGS. 8A and 8B, using the smooth-resurfacing technique as described herein may represent a significant advantage since the damage can be limited to a maximal level defined by the number of pulses N and pulse sequence duration t.sub.s (or alternatively by the pulse separation time t.sub.sep=t.sub.s/N).

[0145] In this invention, the level of superficial heat shock triggering is evaluated by assuming that the level of thermal “needling” is related to the superficial damage resulting from the multiple short-duration exposures. The superficial heat shocking resembles the effects of the micro-needling technique, which aims not to injure keratinocytes but to stimulate them with superficial punctures and without any injury to fibroblasts. The smooth-resurfacing laser-induced thermal triggering mechanism can be viewed as non-ablative thermal “needling” (i.e., triggering) of the total treated skin surface, with the action of the spatially sharp needles being replaced by the action of temporarily “sharp” but spatially extended heat shock pulses.

f) DISCUSSION

[0146] During an Er:YAG laser pulse sequence, the laser-generated heat dynamics exhibits two phenomena: i) intense short-duration thermal pulses resulting from individual laser pulses i, with peak temperatures (T.sub.max-i) at the surface that may exceed 200° C. (see FIG. 1A), biochemically directly affecting only the approximately 10 μm deep superficial tissue layer; and ii) a slow gradual build-up of the spatial temperature distribution over the total duration of the sequence, extending several hundred microns deep into the tissue, with the surface temperatures (T.sub.s) typically below 50-70° C. (see FIG. 1B).

[0147] Similarly, based on FIGS. 7, 8A and 8B, it can be concluded that for the smooth-resurfacing as described herein, there are two optimal treatment regimes.

[0148] When a maximal heat shock triggering effect, with optimal short exposure damage Ω, and moderate coagulation depths of about 100 μm are desired, the optimal sequence durations and number of pulses are in the range up of about t.sub.s=1−3 s (See FIG. 8A) and N=12-30 pulses, or N=12-24 pulses (See FIG. 8B). An exemplary set of protocols for full beam superficial triggering of skin is depicted in Table 1 below. For example, without anesthesia, N=10-20, or 10-18 or about 12 may be preferable. For anesthesia N=10-30, or 15-25 or about 24 may be preferable.

TABLE-US-00001 TABLE 1 Without topical anesthesia Sequence duration, t.sub.s ms 150 300 900 1200 1500 Fluence, F.sub.s J/cm.sup.2 1.3 1.7 2.8 3.2 3.4 Pulse number, N / 6 6 12 12 12 ΔT.sub.s/ΔT.sub.p % 92 90 92 93 90 Coagulation depth, z.sub.c μm 64 68 71 72 66 Superficial Ω  / 0.22 0.37 0.53 0.69 0.82 With topical anesthesia Sequence duration, t.sub.s ms 150 600 1350 1800 3000 Fluence, F.sub.s J/cm.sup.2 1.6 3.0 4.2 4.8 6.0 Pulse number, N / 6 12 18 24 24 ΔT.sub.s/ΔT.sub.p-EMLA % 114 117 116 117 118 Coagulation depth, z.sub.c μm 82 109 121 133 162 Superficial Ω.sub.s / 0.35 0.64 0.84 1.07 1.29

[0149] And when deeper coagulation depths are to be achieved, then long sequence durations of about t.sub.s=5-10 s (See FIG. 7), consisting of a large number of pulses of about N=80-150, are to be used. Under these deep coagulation conditions, the heat shock triggering effect is extremely small, typically below Ω.sub.s2=0.05. An exemplary set of protocols for full beam deep coagulation of skin is depicted in Table 2 below.

TABLE-US-00002 TABLE 2 Without topical anesthesia Sequence duration, t.sub.s ms 1200 4200 8100 13800 20250 Fluence, F.sub.s J/cm.sup.2 3.2 5.6 7.2 9.2 10.8 Pulse number, N / 48 84 108 138 162 ΔT.sub.s/ΔT.sub.p % 93 95 92 94 93 Coagulation depth, z.sub.c μm 72 93 117 199 295 Superficial Ω.sub.s / 0.0025 0.0049 0.0056 0.0076 0.0087 With topical anesthesia Sequence duration, t.sub.s ms 1800 6000 12600 20400 30000 Fluence, F.sub.s J/cm.sup.2 4.8 8.0 11.2 13.6 24.0 Pulse number, N / 72 120 168 204 240 ΔT.sub.s/ΔT.sub.p-EMLA % 117 116 118 117 117 Coagulation depth, z.sub.c μm 133 236 461 674 909 Superficial Ω.sub.s / 0.0095 0.0155 0.0234 0.0270 0.0313

[0150] In some examples, the apparatus as described herein may be implemented to provide a pulse sequence with N=100 to 220 pulses and a cumulative fluence of about 7 to 15 J/cm.sup.2 (without anesthesia). In some examples, the apparatus as described herein may be implemented to provide a pulse sequence with N=150 to 250 pulses and a cumulative fluence of about 10 to 25 J/cm.sup.2 (with anesthesia). The sequence duration may be about 0.8 to 1.2 times that required to reach the pain threshold temperature as defined herein.

[0151] The apparatus as described herein may comprise a pre-calibrated mode to which a user may switch that provides one or more of the above pulse sequences, e.g., for superficial triggering and/or deep coagulation.

[0152] Therefore, when the patient's pain tolerance is used as the treatment safety criteria for selecting appropriate laser parameters, the above heat dynamics leads to two distinct treatment protocol regimens: i) the optimal superficial triggering regimen, represented by Table 1; and ii) the optimal direct stimulation and coagulation regimen, represented by Table 2. In conclusion, based on the analysis and taking in consideration also the treatment duration, the following three full beam smooth-resurfacing treatment embodiments are preferred:

aa) INTENSE smooth-resurfacing protocol for maximal superficial heat shock triggering with exemplary protocol parameters according to Table 1 for superficial Ω.sub.s=0.82 and Ω.sub.s=1.29 for treatments without and with topical anesthesia, correspondingly.

[0153] For example, a pulse sequence with 10 to 35 (or 12 to 30) pulses during a sequence duration 300 to 3000 ms (or 1000 to 2500 ms) may be used, having a cumulative fluence of 3 J/cm.sup.2 to 8 J/cm.sup.2. The apparatus as described herein may comprise a pre-calibrated mode to which a user may switch that provides such a pulse sequence, e.g. with a pulse duration depending on the skin type.

bb) HYPERSTACK smooth-resurfacing protocol for maximal deep thermal stimulation, with exemplary protocol parameters according to Table 2 for coagulation depth z.sub.c=296 μm and z.sub.c=480 μm for treatments without and with topical anesthesia, accordingly. For example, a pulse sequence with 100 to 400 (or 200 to 300 or 100 to 200) pulses during a sequence duration of 6000-40000 ms (or 8000 to 20000 ms) may be used, having a cumulative fluence of 8 J/cm.sup.2 to 14 J/cm.sup.2. The apparatus as described herein may comprise a pre-calibrated mode to which a user may switch that provides such a pulse sequence, e.g. with a pulse duration depending on the skin type.
cc) DUAL smooth-resurfacing protocol for when it is desired to achieve both, superficial triggering and deep thermal stimulation effects within a single procedure.

[0154] For example, a pulse sequence with 40 to 90 pulses during a sequence duration of 600-12000 ms may be used, having a cumulative fluence of 7 J/cm.sup.2 to 12 J/cm.sup.2. The apparatus as described herein may comprise a pre-calibrated mode to which a user may switch that provides such a pulse sequence, e.g. with a pulse duration depending on the skin type.

[0155] Our analysis also shows that patterned handpieces may be better suited for the maximal heat shock triggering effect procedures compared to full beam handpieces. Hence, it may be beneficial to use a pixelated beam. For this reason, according to one of the preferred embodiments of present invention, the energy is delivered to the tissue in a “patterned” shape, wherein the laser beam irradiates a number (M) of individual spots S within the treatment area S′. Each spot S having the size (e.g., diameter) d is separated from a neighbouring spot by the distance x (see FIG. 9). The spot size d and the distance x are chosen such that the spot size d is in the range of 0.3 mm≤d≤1.5 mm, and that the tissue coverage T.sub.C=(M×area(S))/area(S′) (in %) is in the range of 25%≤T.sub.C≤65%. Further, the size (e.g., diameter) of the treatment area S′ which comprise all the spots is in the range of 3 to 15 mm. These parameters ensure that, during the sequence duration t.sub.s the thermal diffusion in the lateral direction spreads the heat which is generated by the laser radiation away from a localized spot S towards the surroundings of the spot, thus effectively spatially homogenizing the temperature T.sub.s across the area S′, while at the same time the spotsize d is large enough to enable the pulse fluence F.sub.p to be set to a required value being below the ablation threshold F.sub.abl.

[0156] Namely, with patterned handpieces. the irradiated tissue does not cool down only by the heat diffusion deeper into the tissue but also by the heat diffusion in the radial direction away from the irradiated microspots. This cooling mechanism is more effective during the long sequence duration (t.sub.s) than during the short duration temperature peaks. For this reason, the final sequence temperature (T.sub.a) is more significantly reduced than the temperature peaks T.sub.max-i. As can be seen from FIG. 4, the temperature slope coefficient A.sub.s is for a patterned handpiece by a factor of 3-4 smaller as compared to the slope for the full beam handpiece (See Eq. 7). This means that 3-4 times higher cumulative fluences (F.sub.s) can be delivered without exceeding the pain tolerance threshold temperature. On the other hand, since the radial heat diffusion has a smaller effect on the high temperature peaks (T.sub.max-i), the higher cumulative fluence will for the same level of deep tissue coagulation result in an increased level of superficial triggering. Therefore, patterned handpieces are generally better suited for the maximal heat shock triggering effect compared to full beam handpieces.

[0157] To demonstrate the influence of radial heat diffusion for small diameter beam sizes we carried out a limited numerical analysis using a 3D cylindrical coordinate system model. The resulting temporal profiles of the skin surface temperature during a sequence with N=24 micro pulses, delivered by a full beam ((R11 with 7 mm spotsize) or by a patterned beam handpiece (PS03 with 0.85 mm micro spotsize) are shown in FIGS. 10A and 10B. FIGS. 10A and 10B show calculated temporal profile of the skin surface temperature during a sequence of four SMOOTH mode macro pulses each consisting of 6 micro pulses (resulting in N=4×6=24 micro pulses) for F.sub.s=4.9 J/cm2 with a full beam handpiece (R11 with 7 mm spotsize) (FIG. 10A); and b) F.sub.s=14.7 J/cm2 with a patterned beam handpiece (PS03 with 0.85 mm micro spotsize) (FIG. 10B).

[0158] As can be seen from FIGS. 10A and 10B, for the same final sequence temperature (T.sub.s=60° C.) (and resulting deep tissue response), the high temperature peaks (T.sub.max-i) are approximately 30% higher for the patterned handpiece. A damage integral calculation using Eqs. 5 and 6 shows that this difference results in 2 times stronger (in terms of Ω) patterned thermal “needling” of the tissue.

g) CONCLUSIONS

[0159] In conclusion, Er:YAG laser pulse stacking represents an example of complex thermal exposure dynamics during which the exposure times transition from extremely short to long durations. The tissue effects resulting from these dynamics were evaluated numerically using the VHS model, for two examples of non-ablative or minimally ablative Er:YAG laser pulse stacking treatments: i) the “sub-resurfacing” performed at or near ablation laser fluences; and ii) the “smooth-resurfacing” characterized by below-pain-threshold fluences.

[0160] Based on measurements on abdominal skin, the pain threshold temperature depends mainly on the long-exposure superficial skin temperature (T.sub.s) by the end of the pulse sequence, and not on the peak skin temperatures (T.sub.max) following individual laser pulses within the sequence.

[0161] Our simulations show that for sub-resurfacing (i.e., resurfacing with fluences at or just below the ablation threshold fluence), the parameter range where no excessive damage to the tissue can occur is very narrow. On the other hand, using pain tolerance as an indicator, the smooth-resurfacing treatments can be performed more safely and without sacrificing the treatment efficacy.

[0162] Two preferred smooth (non-ablative) resurfacing treatment modalities were identified. One involves using optimally long pulse sequence durations (≈1-3 s) with an optimal number of pulses (N≈10-30), resulting in a maximal short-exposure superficial tissue response and moderate coagulation depths. And for deeper coagulation, without significant superficial heat shocking, very long pulse sequences (>5 s) with a large number of delivered pulses are to be used (preferably in combination with topical anesthesia).

h) EXEMPLARY EMBODIMENT

[0163] FIG. 11 shows an exemplary embodiment 100 of an apparatus according to the present invention. It may comprise a user interface 110, a control unit 120 and a laser source 130, e.g. as described herein. User interface 110 may receive one or more selection inputs and/or parameters. Based thereon, selection inputs and/or parameters may be provided to control unit 120. Control unit 120 may control laser source 130 accordingly to provide one or more laser pulse sequences as described herein. Control unit 120 may include a computer and/or processor and a memory. The memory may comprise one or more computer programs stored thereon as described herein.