Sophisticated algorithms and technological advancements help realize new features in excimer refractive surgery such as shorter treatment times and distributed thermal load with optimized spot patterns.1 Ultraviolet radiation commonly used in laser refractive surgeries is regarded as “cold” because the thermal relaxation time of the molecules is usually shorter than their thermal denaturation time.2 However, even these ultraviolet laser pulses also impart a certain thermal load to the corneal tissue. This can be observed as an increase in ocular surface temperature clinically3 and in laboratory settings.4
Parameters such as ambient temperature and convection heat transfer coefficient affect the heat loss from the corneal surface.5 The uncertainty in temperature measurements in porcine corneas due to emissivity errors6 and camera angle7 has been reported. Commercially available thermography imaging systems have been used for the evaluation of mean and maximum temperature rise within an ablated area.8 The average ocular surface temperature in humans has been reported to be approximately 34 °C.9 In clinical settings, where the globe is exposed using an eyelid speculum, a decrease of approximately 4 °C is expected.10 At temperatures above 40 °C, thermal damage of the corneal tissue may occur.11 Assuming these metrics, the change in peak corneal temperature during excimer laser ablation shall be maintained at best below 6 °C, and definitely below 10 °C to avoid denaturation of the collagen proteins inducing thermal damage.12,13
Minimizing the thermal load of ablation in high-speed laser corneal refractive surgery has been analyzed theoretically and clinically in the past, in PMMA and porcine and human corneas.15 However, the impact of certain system and ablation parameters such as optical zone (OZ) size and refractive correction have not been individually and extensively analyzed. Our aim with this study was to quantify and analyze the impact of these parameters on the increase in peak corneal temperature during “fast repetition rate” excimer laser ablation in PMMA plates and porcine eyes, while using a flying spot algorithm to control the thermal load (actually the change in corneal temperature) on the cornea.
Materials and Methods
Collectively, 42 ablations were performed (21 on PMMA plates + 21 on porcine eyes). All ablations performed on ex vivo porcine eyes occurred within 12 hours of procuring the eyes. The porcine eyes were stored in buffered chilled solution (4 to 7 °C) and PMMA plates at room temperature (20 to 22 °C), during transportation and before ablation (controlled to a relative humidity below 50%). A B&L Teneo 317 Model 2 laser system with the software version 1.3 (Bausch & Lomb) was used for performing the ablations. As for the technical aspects, the system was used in its standard configuration, which means 265 mJ/cm2 radiant exposure, with a beam diameter of 1 mm FWHM, approximately 1.6 mJ energy per pulse with approximately 8 ns pulse duration, and 500-Hz repetition rate, leading to approximately 0.54 µm and approximately 360 pl ablation per pulse, and approximately 2.18x cornea-to-PMMA ratio. A pulse sorting algorithm is implemented to maximize the entropy of the pulse sequencing (a special sort of pseudo-random application). The ablation specifications for each test setting are presented in Table 1. For all porcine surface ablations, the eyes were deepithelialized manually prior to ablation (using an Amoils brush) ( https://www.dm.de/dontodent-elektrische-zahnbuerste-active-p4058172230332.html). Care was taken to use only eyes with pristine corneas in our experiments; any corneal cloudiness or incompletely deepithelialized corneas were discarded for further processing.
Ablation Parameters Including Overall Results
The PMMA plates had a size of 60 mm × 50 mm × 2 mm divided into four central 14 mm × 14 mm ablation “fields” (arranged as a 2 × 2 array), each of them including an 8-mm circular aperture to simulate the pupil.
For all ablations, nonwavefront-guided aspheric ablation profiles were calculated for a vertex distance of 12.5 mm and keratometric readings of 44.00 D. The first four ablations (ID 1–4) are PRK spherocylindrical ablations and serve as a repeatability metric to determine how reliable the temperature recording is for the same settings. The second four ablations (ID 5–8) are transepithelial ablations and serve as a linearity metric to determine how much the temperature recording is affected by ablation depth/time/location of maximum ablation (central in myopic ablations, ring-shaped in hyperopic ablations). The next four ablations (ID 9–12) are laser in situ keratomileusis (LASIK) ablations, and serve as a linearity metric to determine how much the ablation depth/time/location affects the temperature recording. By comparing the second to the third group of four ablations, we can qualify and quantify the effect of the transepithelial component of the ablations on the recorded temperature. By comparing hyperopic ablations (ID 5, 6, 11, 12) with myopic ablations (ID 7–10), we can qualify and quantify the effect of the location of the maximum ablation on the recorded temperature. The next three ablations (ID 13–15) are ablations of the same power but different OZs and serve as a metric to determine how much the ablation diameter affects the temperature recording. The last two groups of three ablations each (ID 19–21; ID 22–24) are ablations of different powers and OZ but resulting in the same central/maximum ablation depth or at the same ablation volume and serve as a metric to determine how much the pulse density affects the temperature recording.
Every eye was mounted on a self-made holder (self-made device also used for other publications) and was recorded with a thermal camera during ablation. An infrared thermal camera (FLIR-A 615, https://www.flir.de/products/a615/) was positioned at 22 cm distance from the ablation focus and was tilted horizontally to an angle of approximately 60 degrees (meaning more vertical than horizontal) (Figure A, available in the online version of this article).
Sketch and real image of the laser-camera set-up. Every polymethylmethacrylate plate or porcine eye was mounted on a holder and recorded with a thermal camera during ablation. An infrared thermal camera (FLIR-A 615) was positioned at 22 cm distance from the ablation focus and was tilted horizontally to an angle of approximately 60 degrees.
Every ablation was recorded at 200 frames per second (fps) in windowed mode (approximately 183 µm/pixel) with a centered elliptical window (region of interest) approximately 10 mm in size (at the major axis). The region of interest was specified manually and selected such that it covered the cornea completely and tightly enough that the variability between the porcine eyes could be accounted for. Therefore, some intact cornea around the ablation zone may have also been analyzed. The camera angle was manually adjusted before the experiment to completely cover the eyeball in the windowed mode. The region of interest and the camera angle were neither shifted nor adjusted throughout the experiment. Using the camera software, the measurements were internally adjusted for the recording distance, and emissivity was set at 0.86 in agreement with the literature.16 With our test setup, we can expect a measurement error of ±0.5 °C due to the thermal camera.16 The pixel with the highest recorded temperature within the region of interest was analyzed. The baseline maximum temperature before starting the ablation and the peak of the maximum temperature achieved during the entire ablation were analyzed manually from the graphical representation of the maximum temperature achieved during the video recordings. The change in peak corneal temperature was calculated from the difference of the two recorded values, the peak of the maximum temperature achieved during the entire ablation, and the baseline maximum temperature before the ablation.
With our setting we could evaluate the following aspects: correlation between measurements on PMMA versus porcine eyes; repeatability of the change in peak corneal temperature; the effects of refractive correction, laser-assisted epithelial removal in transepithelial PRK, myopic versus hyperopic corrections, and optical zone on the change in peak corneal temperature. We have also generated two hypersurface models.
The user only enters the refractive power and the optical zone into the laser. Total ablation zone, depth, volume, duration, number of pulses, pulse density, and slope are results of the user input. So, we performed a fit of the form:
The model can be refined adding the cross-product. So, we performed a second fit of the form:
Through these models, and setting a maximum permissible change in peak corneal temperature, one can obtain the refractive correction-OZ isothermal lines.
Data analysis was performed using Microsoft Excel (Microsoft Corporation) and a P value less than .05 was considered statistically significant. The t tests were assessed between groups, together with the P value of the correlations (ie, whether or not the slope is different from zero).
Ooi et al17 presented a three-dimensional radially symmetric boundary element model of the human eye for simulating changes in corneal temperature during laser thermokeratoplasty. They calculated the temperature field for heating by both a pulsed laser and a continuous wave laser. Ocular surface temperature has been shown to vary topographically.18 These differences in full three-dimensional analysis should be properly accounted for in two-dimensional thermal image analysis. Another group previously analyzed this issue in a case study on the porcine eye two-dimensional temperature distributions with a lateral resolution of 170 µm and line scans with a temporal resolution of 13 µs.16
Ablation procedures are always associated with a certain level of thermal load, so corneas undergoing excimer laser refractive surgery also suffer from this thermal load. Historically,15 a maximum corneal temperature of 40 °C is considered safe. Some groups consider approximately 45 °C as a better limit value for the cornea.19 The ocular surface temperature averages 34 °C, but after 1 minute without blinking (eg, due to an eyelid speculum), ocular surface temperature cools down to approximately 30 °C at room temperature.20 This means that for excimer laser refractive surgery, a thermal load shall be confined to 6 °C to be thermally safe and will become unsafe beyond 15 °C. The corridor of 6 to 15 °C is an increasing risk region.
Betney et al21 found that mean central ocular surface temperature after epithelial debridement was 29.2 ± 0.4 °C. Mean peak ocular surface temperature during PRK was 37.8 ± 0.7 °C, with most of the temperature increase occurring during the first 15 seconds.
In our study, we could directly examine the effects of refractive correction, OZ, and laser-assisted epithelial removal in transepithelial PRK on the change in peak corneal temperature of the ablation onto the cornea. Figure D shows an excellent correlation between the change in peak corneal temperature of equivalent ablations in PMMA and porcine eyes. This indicates that the measurements can be trusted (and are not a matter of chance), and that the uncertainties (noise) are smaller than the differences in signal.
The eyes were stored in a chilled solution and were taken out of the solution minutes before they were prepared and processed by the laser, so that they were not chilled anymore but also did not reach ambient temperature. Due to the long duration of performing all tests belonging to the experiment, we preferred not to store them at ambient temperature. There had been control measurements showing that the increase in temperature toward ambient temperature occurs much slower than the duration of the ablation process, so that the temperature drift toward ambient temperature shall not act as an important confounding factor.
As in previous studies, we recorded the baseline maximum temperature before starting the ablation for a few seconds (hundreds of frames) and could observe for all tests that this temperature was stable before ablation. Then the peak of the maximum temperature (irrespective of its location within the region of interest, ie, the spatial maximum for a given time point) was continuously recorded, and the maximum of these maxima (irrespective of its time point during the ablation process, ie, the spatial maximum for the entire ablation) was selected. The change in peak corneal temperature was calculated from the difference of the two recorded values, the peak of the maximum temperature achieved during the entire ablation and the baseline maximum temperature before the ablation. Per definition, there is no added error from the fact that the two values may correspond to different spatial locations, because it still represents the change in peak corneal temperature occurring during the ablation process. We did not intend to report the actual maximum temperature rise at a certain location. That would be feasible by calculating the differences in a point-by-point fashion throughout the ablation (ie, the baseline would be a matrix of temperatures prior to the ablation, and not a single number) and then analyzing the maxima of each matrix difference to determine the maximum temperature rise. Our aim was rather to document the change in peak corneal temperature. However, for myopic ablations, both points correspond well because the maximum ablation is at the center and the maximum temperature was always close to the center. In contrast, for hyperopic ablations the differences may be larger because the ablation is spread over a pericentral annulus (as discussed in previous publications).
Myopic corrections produced higher changes than hyperopic treatments, but on a per-diopter basis, the opposite is true. The per-diopter basis is the more accurate way to compare, with a notation that myopic treatments are often higher in dioptric correction than hyperopic ones and thus may have higher temperature changes in individual cases.
Of all analyzed metrics, the refractive correction together with the maximum slope of the treatment were showing the best coefficient of determination. Both showed that higher dioptric powers and steeper maximum slopes are directly related to higher change in peak corneal temperature.
The analysis of our results revealed the thermal response of the cornea to varying test parameters. The range of tested refractive correction (approximately −18.00 to +5.50 D) represents a wide range, which is beyond the standard limits in refractive surgery. The OZ of 8.5 mm may be considered as the upper boundary value in refractive surgery. The lower boundary of the 4.5-mm OZ was selected to increase the radiation power density much higher than the clinically used values. The change in peak corneal temperature decreased linearly (R2 = 0.47) with increasing ablation diameter, peaking at 4.5 mm OZ (16 °C). These results correspond well with the theoretical predictions, because the pulse density decreases for a wider OZ and there is a higher random chance to hit a cooler corneal region.
Interestingly, all correlations showed a clear nonzero intercept (ranging from 6.4 to 7.7 °C), meaning that ocular surface temperature increased by at least 7 °C even for low corrections in wide optical zones, and further increased for higher corrections or smaller OZs. This is compatible with the minimum change in peak corneal temperature measured in this series being 5 °C in a hyperopic ablation (+1.50 D in 6.7-mm OZ with 8.8-mm total ablation zone leading to 41 µm maximum depth with 6 µm central depth, 1.1 µL ablation volume, taking 6 seconds and 3,220 pulses).
In our setting, transepithelial PRK ablations led to a higher change in peak corneal temperature than LASIK ablations, likely due to the extra ablation depth/volume/pulses required to remove the epithelium (phototherapeutic keratectomy-like ablation). Our results correspond well with the expected 1 to 2 °C higher transepithelial thermal response as predicted by theoretical models.22 It is expected that the role of epithelium as a protective outermost layer of the cornea (besides the tear film) results in a higher response to the implied laser energy, in comparison to the deeper layers of the cornea. In this regard, Figure BB and Figure CB clearly show a break of approximately 3 seconds between the epithelial and stromal stages of the transepithelial PRK ablations, leading to a partial recovery of the ocular surface temperature followed by a second (stronger) increase of ocular surface temperature during the stromal stage.
The ablation parameters used for the transepithelial PRK tests aimed to cover the range of potential transepithelial PRK applications. Myopic errors up to −10.00 D have been reported in the literature as being effectively corrected using transepithelial PRK; −4.00 D is a common setting in refractive surgery (probably close to the mean/median correction used overall). Hyperopia of +1.75 D is common in hyperopic ablations (probably close to the mean/median correction used overall), and +4.75 D has been reported in the literature as being effectively corrected using transepithelial PRK for high hyperopia.23,24
With the parameters that are under user control (refractive power and OZ), we developed two multilinear models (with and without the cross-product). The fit was good (coefficient of determination 79%). In the first model (without cross-product), square root of the diopters is a metric of the strength of the treatment and OZ^4 relates to the overall volume, whereas in the second model (with cross product), D relates to the strength of the treatment, OZ^2 relates to the ablation depth, and D*OZ relates to the maximum slope.
The models allow for the calculation of the isothermal lines. According to the 6 to 15 °C corridor and considering that 6.4 °C was the smallest intercept in our data, the 8, 10, and 12 °C isothermal lines are depicted in Figure F). It can be seen that to keep the change in peak corneal temperature below 8 °C, only 3.00 to 4.00 D can be treated (in OZ of 6.5 mm or larger); for 10 °C, a maximum of 7.00 D (in OZ of 7 mm or larger) can be treated; and for 12 °C, 10.00 D can be treated in at least 6 mm OZ.
Overall, the change in peak corneal temperature is a function of the laser energy and its distribution. For a constant pulse energy, more pulses (as required for higher correction powers) accumulate more energy, and lead to a higher temperature rise. Therefore, the spatial and temporal distribution of the laser energy is key to controlling the change in peak corneal temperature. As demonstrated by controlling the local heating and cooling phases and the heat propagation, one can optimize the pulse distribution to keep the temperature rise below a predetermined limit.22,25 Intuitively, larger ablation zones (OZs or total ablation zones) provide longer times for each individual corneal position to cool down, reducing the overall peak change in peak corneal temperature.
Because the energy per pulse is ultimately the building block of the cumulative change in peak corneal temperature, one could hypothesize that reducing pulse energy would be a protective measure against excessive peak temperature rise. However, as published before,4 the temperature rise is a function of the local laser power (pulse energy × local repetition rate at each corneal position). Furthermore, there seems to be an optimum between energy of the individual pulses and total number of pulses, creating a delicate balance.26 It may be that the 1.6 mJ per pulse examined here is already beyond the optimum, and that 1.6 mJ per pulse combined with a 500-Hz repetition rate produces a higher than optimal local laser power.
One can argue that the limits of 4015 to 45 °C21 are historical and somewhat arbitrary. Protein denaturation occurs in vitro somewhere between these two values, but this may not reflect the clinical experience. One reason for that may be that the potential denaturated region may only be a few micrometers in depth. This thin layer of damaged tissue may then be removed by the next pulses so that at the end only a thin layer on the stromal surface may remain affected. Kymionis et al27 used another laser platform with two different repetition rates (200 vs 400 Hz) and observed a significantly higher incidence of haze with the faster repetition rate.
However, other studies undermine the impact of ablation frequency regarding their clinical side effects. Khoramnia et al28 evaluated the effect of three excimer laser ablation frequencies (200, 500, and 1,000 Hz) on the cornea using a 1,000-Hz scanning-spot excimer laser and found no specific side effects associated with the high repetition rates based on structural and ultra-structural evaluation of corneas. Similar findings were reported by Shanyfelt et al29 for corneal ablation profiles created at 60 and 400 Hz.
Another option is to extend the safe corridor by cooling down the cornea (before, during, and/or immediately after ablation) by applying a chilled balanced salt solution.30 This method was discussed at the time PRK dominated the market, but may have lost some of its popularity now.31–34 It has been shown30 that by instilling chilled balanced salt solution, ocular surface temperature may drop by 10 to 14 °C (from 30 to 34 °C to 20 °C). However, this transient cooling is counterbalanced by the heat propagation from deeper corneal layers (with a thermal relaxation time of approximately 7 seconds) and by the heat propagation from the environment (with a longer relaxation time in the order of minutes), so that after approximately 40 seconds the ocular surface temperature stabilizes at approximately 30 °C. We also recorded this effect with the non-invasive thermal camera in a few real treatments on patients' eyes (data not shown), and observed that the cornea cooled down to approximately 22 °C (matching well the reported approximately 20 °C30). So, a feasible approach would be to apply chilled balanced salt solution “just before” commencing the laser ablation and also immediately thereafter.
The thermal load problem in refractive surgery in relation to the laser pulse frequency has been explored by many researchers and commercial platforms in the past.35 Our results suggest that high-refractive corrections (eg, −8.00 D) or especially transepithelial ablations (eg, −4.00 D) may lead to a change in peak corneal temperature beyond 10 °C, leading to potential thermal damage.12 Furthermore, the difference between the epithelial and stromal thermal response to ablation (higher for epithelium) shines light on the thermal safety control of transepithelial and stromal refractive procedures. The role of the epithelium in the corneal thermal response must be considered while planning refractive surgeries using transepithelial ablations, in addition to the considerations of epithelial thickness profile versus the applied epithelial ablation profiles.36 Our experimental results may be extended to clinical applications considering the similarities between human and porcine eyes.8 Because the tested parameters are interrelated for refractive surgery planning and laser system specifications, our recommendation covers many of these parameters together. The impact of each individual parameter on the change in peak corneal temperature beyond a safe limit is reflected in Figures 1–4 and Figure E
Mrochen et al4 performed a similar study in bovine eyes investigating the influence of temporal and spatial spot sequences on the ocular surface temperature increase during corneal laser surgery. They recorded the increase in the corneal temperature for various refractive corrections and temporal distribution of ablation profiles (four spot sequences: line, circumferential, random, and an optimized scan algorithm) using an infrared camera. They found that the highest maximum temperature increases were observed with the line and circumferential scans and the lowest with the random and optimized scans. However, we report here a lower change in peak corneal temperature for the same refractive correction compared to their optimized scan algorithm, allowing a spot positioning with a maximum local frequency of 200 Hz at each point of the treatment zone.
Although we tested the extremities of each parameter technically feasible with our set-up, extrapolation of our results beyond these limits must be done with caution. Furthermore, the measurement was updated at every second to third laser pulse due to the acquisition rate of the camera. This means that the laser repetition rate of 500 Hz was higher than the frame rate (200 Hz) of the camera. We were able to choose between different frame rates (in windowed mode) with the thermal camera used in our test set-up. The frame rate affected the size of the acquisition window, with a higher frame rate allowing a smaller acquisition window. Our choice was based on the largest acquisition window compatible with the fastest frame rate. Then the working distance was selected to offer a window large enough to comfortably cover the limbus in the porcine eyes (with some room for variability in the eyes). Finally, the observation angle was selected as the most vertical line that did not interfere with the laser contour.
Due to the vastly different time scales (nanoseconds for the laser pulses, milliseconds for the camera frame rate), the resulting thermal images contain information that is integrated or averaged in time. However, our interest was to measure the temperature of the remaining cornea as opposed to the dynamics of the ejected and ablated tissue. Hence the “slow time domain” (averaging/integrating) helped us to get closer to the biological effects on the residual stroma.
The dynamics of the temperature change during the ablation process could also be relevant and were not analyzed in this study. The extrapolation of our results based on postmortem porcine eyes to human eyes could be affected by the condition of the porcine eyes. However, care was taken to use only eyes with pristine corneas in our experiments.
The use of a literature-based emissivity factor may be considered a limitation of this study.
The presented results reiterate the dilemma in excimer laser development: how to maintain a safe thermal load at the cornea while speeding up treatment times. Either aspect can eventually affect the hydration of the corneal surface during the treatment.37 The corneal hydration reduces with increasing treatment times, affecting the ablation properties of the cornea, favoring a faster treatment. However, faster treatments may increase the risk of corneal tissue denaturation. In addition to these factors, the benefits of transepithelial refractive procedures must be evaluated against the additional increase in corneal temperature compared to purely stromal refractive procedures. A delicate balance needs to be maintained to achieve optimum ablation efficiency in coherence with safe thermal controls.
The effect of different parameters on change in peak corneal temperature during corneal laser ablation could be quantitatively evaluated. The proposed models adequately represent the measured thermal load. Without cooling of the cornea, the laser system used seems to be safe in terms of thermal load for ablations up to 3.00 to 4.00 D combined with OZs of 6.5 mm or larger, and it may remain safe for up to 7.00 D if the OZs are enlarged to 7 mm or more. These results support the need for cooling the cornea to avoid approaching the denaturation temperature. This is particularly important for transepithelial ablations, which further increase the temperature due to the extra phototherapeutic keratectomy–like ablation, and the more sensitive thermal response of the epithelium.27