Analysis of the change in peak corneal temperature during excimer laser ablation in porcine eyes

Samuel Arba Mosquera; Shwetabh Verma

doi:10.1117/1.JBO.20.7.078001

2 July 2015 Analysis of the change in peak corneal temperature during excimer laser ablation in porcine eyes

Samuel Arba Mosquera, Shwetabh Verma

Author Affiliations +

Journal of Biomedical Optics, Vol. 20, Issue 7, 078001 (July 2015). https://doi.org/10.1117/1.JBO.20.7.078001

Abstract

The objective is to characterize the impact of different ablation parameters on the thermal load during corneal refractive surgery by means of excimer laser ablation on porcine eyes. One hundred eleven ablations were performed in 105 porcine eyes. Each ablation was recorded using infrared thermography and analyzed mainly based on the two tested local frequencies (40 Hz, clinical local frequency; 1000 Hz, no local frequency). The change in peak corneal temperature was analyzed with respect to varying ablation parameters [local frequency, system repetition rate, pulse energy, optical zone (OZ) size, and refractive correction]. Transepithelial ablations were also compared to intrastromal ablations. The average of the baseline temperature across all eyes was 20.5°C±1.1 (17.7°C to 22.2°C). Average of the change in peak corneal temperature for all clinical local frequency ablations was 5.8°C±0.8 (p=3.3E−53 to baseline), whereas the average was 9.0°C±1.5 for all no local frequency ablations (p=1.8E−35 to baseline, 1.6E−16 to clinical local frequency ablations). A logarithmic relationship was observed between the changes in peak corneal temperature with increasing local frequency. For clinical local frequency, change in peak corneal temperature was comparatively flat (r²=0.68 with a range of 1.5°C) with increasing system repetition rate and increased linearly with increasing OZ size (r²=0.95 with a range of 2.4°C). Local frequency controls help maintain safe corneal temperature increase during excimer laser ablations. Transepithelial ablations induce higher thermal load compared to intrastromal ablations, indicating a need for stronger thermal controls in transepithelial refractive procedures.

1. Introduction

Laser-based refractive surgery techniques are constantly being developed with a common aim of optimizing the surgery outcomes in terms of visual acuity, contrast sensitivity and night vision. Sophisticated algorithms and technological advancements help realize new features in refractive surgery like shorter treatment times, distributed thermal load with optimized spot patterns,¹^,² calibration systems,¹^,³ precise eye tracking, and optimized ablation efficiency at non-normal incidence.⁴^,⁵ Spatial distribution of laser pulses on the cornea is controlled such that each pulse sequentially ablates a small amount of corneal tissue, collectively etching a very refined pattern calculated and designed to counterbalance the aberrations. Ultraviolet (UV) radiation commonly used in laser refractive surgeries is regarded as “cold” radiation. The reason for this consideration is that the thermal relaxation time of the molecules is usually shorter than the thermal denaturation time.⁶ However, even these UV laser pulses also impart a certain thermal load to the cornea tissue seen as an increase in ocular surface temperature (OST), observed clinically⁷^–⁹ as well as in laboratory settings.¹⁰^,¹¹ Furthermore, smaller laser spots are used along with the higher system repetition rates controlling the treatment time in most modern refractive surgery systems; each of these factors could impact the thermal response of the corneal tissue individually and in combination with other factors.

Heat that is gained from blood is conducted across the ocular media and is transferred into the environment from the corneal surface via convection and radiation. Parameters such as ambient temperature and convection heat transfer coefficient affect the heat loss from the corneal surface.¹² Contact thermometry is nowadays replaced by noninvasive infrared (IR) thermography. Thermography, however, does not measure temperature directly, but measures radiation intensity from which temperature is calculated. This means that parameters such as spatial, temporal and temperature resolution, field of view, angle of observation, the calibration of the apparatus, and the shielding against stray-radiation may influence the accuracy of the result. The emissivity of the emitting surface is highly important. In the absence of a tear film, the determined temperature is considered from the cornea.¹³

The use of IR thermography in ophthalmology was pioneered by Mapstone through a series of experiments published in 1968.¹⁴^–¹⁷ The uncertainty in temperature measurements in porcine cornea due to emissivity errors¹⁸ and camera angle¹⁹ has been previously reported. Commercially available thermography imaging systems featuring much greater system time constants, have been used for the evaluation of mean and maximum temperature rise within an ablated area.²⁰ Kim et al.²¹ studied the expression patterns of heat shock proteins following eyeball heating or cooling in relation to corneal wound healing and intraocular complications after excimer laser treatment. They were able to demonstrate that heat shock proteins were induced by thermal preconditioning and appeared to be a major factor in corneal protection against serious thermal damage.

Ishihara et al.⁸ showed that the transient surface temperature of the cornea during ablation was much higher than previously reported, with thermal radiation from the surface of porcine corneas during ArF excimer laser irradiation showing a temperature over 100°C at a fluence of $80 mJ / {cm}^{2}$ and of 240°C at a fluence of $180 mJ / {cm}^{2}$ .

Maldonado-Codina et al.²² investigated the temperature changes occurring during photorefractive keratectomy (PRK) in 19 bovine corneas when performed at different ablation depths using noncontact, color-coded ocular thermography. They observed an average temperature rise of 8°C at the corneal surface with a maximum rise of 9°C. They found a positive correlation between the refractive correction and the peak temperature rise. De Ortueta et al.²³ evaluated the thermal load of ablation in high-speed laser corneal refractive surgery with the AMARIS excimer laser (SCHWIND eye-tech-solutions). They found that, overall, the maximum temperature change of the ocular surface induced by the refractive ablations was $\leq 4 ° C$ . The increase in the peak temperature of the ocular surface never exceeded 35°C in any case. This low-thermal load was independent of the amount of correction the eye achieved. Vetrugno et al.²⁴ reported that “high-repetition-rate excimer laser systems require spot sequences with optimized temporal and spatial spot distribution to minimize the increase in OST.” Wernli et al.²⁵ evaluated the effects of initial surface temperature of poly-methyl-methacrylate (PMMA) plates used for daily laser calibration. They found that the ablation depth increased linearly from 73.9 to $96.3 μ m$ within a temperature increase from 10.1°C to 75.7°C (increase rate of $0.3192 μ m / K$ ). The linear correlation was found to be significant ( $P < 0.05$ ) with a coefficient of determination of $R^{2} = 0.95$ .

The average OST in humans has been reported to be approximately 34°C.²⁶^–²⁹ In clinical settings, where the globe is exposed using a lid speculum, a decrease of 3 to 4°C is expected in the corneal surface temperature.²³ At temperatures above 40°C, thermal damage of the corneal tissue may occur.³⁰ Assuming these metrics, the change in peak corneal temperature shall be maintained at best below 6°C, and definitely below 10°C, during the excimer laser ablation. If this thermal load is not controlled, it may lead to denaturation of the collagen proteins inducing thermal damage.¹¹^,³⁰^,³¹

Minimizing the thermal load of ablation in high-speed laser corneal refractive surgery has been analyzed theoretically and clinically in the past, in PMMA, porcine as well as human cornea, generating several models¹¹^,²⁰^,³²^–³⁴ for temperature increase during laser ablation. However, the real impact of the thermal load in laser corneal refractive surgery is still controversially discussed and warrants further in-depth exploration, particularly for high-repetition-rate scanning-spot excimer laser applications. The impacts of certain system and ablation parameters such as local frequency, system repetition rate, pulse energy, optical zone (OZ) size, and refractive correction, have not been individually and extensively analyzed in the light of an increase in optical surface temperature. Our aim with this work is to quantify and analyze the impact of these parameters on the increase in peak corneal temperature during “fast repetition rate” excimer laser ablation in porcine eyes, while using a flying spot algorithm to control the thermal load on the cornea. To these tested parameters, we analyzed the thermal impact of transepithelial ablations in comparison to stromal ablations. We used noninvasive IR thermography with a fast acquisition rate to record each ablation.

2. Methods

All the ablations were performed on ex vivo porcine eyes within five hours of procuring the eyes. The porcine eyes were stored in glucose solution (1L saline solution $+ 19 ml$ of Dulbecco's Modified Eagle Medium low glucose solution from PAA Laboratories GmbH) at room temperature ( $\sim 23 ° C$ ), during transportation and before ablation. Every eye was mounted on a holder and was recorded with a thermal camera during ablation. A customized AMARIS laser system (SCHWIND eye-tech-solutions, Kleinostheim, Germany) allowing for manipulation of ablation parameters was used for performing the ablations with only a single fluence. The treatment plane was calibrated to the corneal vertex with the help of a slit lamp from the laser system. The impact of system and ablation parameters such as local frequency, system repetition rate, pulse energy, OZ, and refractive correction were analyzed, with only one parameter varied at a time. The system and ablation specifications for each test setting are presented in Table 1. For most of the ablations, the eyes were de-epithelized manually prior to ablation (using an Amoils brush). However, to analyze the role of epithelium in temperature control, items 32 and 34 (in Table 1) were each ablated on porcine eyes with natural epithelium. Each item in the Table 1 was repeated in three different eyes to analyze repeatability; additionally, $a$ and $b$ in items 33 and 35 represent consecutive ablations performed in the same eye (one on top of the other after a time gap of $\sim 90 s$ ) to analyze the thermal response of the outer and comparatively deeper stromal layers. The ablations represented with $a$ and $b$ had identical ablation parameters. The items 33 ( $a + b$ ) and 35 ( $a + b$ ) were also repeated in three different eyes. Therefore, items 32 and 34 analyzed the corneal thermal response to transepithelial ablations, items 33 and 35 ( $a$ ) analyzed the corneal thermal response to first stromal ablations (closer to the corneal surface) while items 33 and 35 ( $b$ ) analyzed the corneal thermal response to second stromal ablations (starting at deeper layers within the same eye). The second stromal ablation was performed $\sim 128 μ m$ (nominal depth of the treated correction of $- 9 D$ with an OZ 6.3 mm) deeper than the first stromal layer. It must be noted that with the first stromal ablation ( $33 a$ and $35 a$ ), the nominal ablation depth of $\sim 128 μ m$ was achieved only in the center. In order to bring the periphery in level with the center (within the OZ) and achieve a flat stromal bed, before the second stromal ablation, the cornea was first ablated with a $+ 8 D$ hyperopic profile (with no ablation at the center and $\sim 128 - μ m$ ablation at the periphery). Subsequently, the second stromal ablation was performed (represented by $33 b$ and $35 b$ in Table 1) on a relatively flat stromal bed.

Table 1

System and ablation parameters. Each item was repeated in three different porcine eyes. Further, a and b in items 33 and 35 represent consecutive ablations in the same eye. Items 32 and 34 were each ablated on three different porcine eyes with natural epithelium. The rest of the ablations were performed on manually deepithilialized eyes (Amoils brush). For all the ablations, nonwavefront guided aspheric profiles were used, calculated for a vertex distance of 12 mm and keratometric readings (K1 and K2) of 44 D.

Item No.	System repetition rate (Hz)	Refractive correction (D)	Ablation depth (μm)	Optical zone (mm)	Local frequency (Hz)	Pulse energy (mJ)
Local frequency
1	1050	$- 9.0$	128.2	6.3	40	1.33
2	1050	$- 9.0$	128.2	6.3	1000	1.33
3	1050	$- 9.0$	128.2	6.3	10	1.33
System repetition rate
4	1050	$- 9.0$	128.2	6.3	40	1.33
5	1050	$- 9.0$	128.2	6.3	1000	1.33
6	750	$- 9.0$	128.2	6.3	40	1.33
7	750	$- 9.0$	128.2	6.3	1000	1.33
8	500	$- 9.0$	128.2	6.3	40	1.33
9	500	$- 9.0$	128.2	6.3	1000	1.33
Optical zone size
10	1050	$- 9.0$	53.0	4.0	40	1.33
11	1050	$- 9.0$	53.0	4.0	1000	1.33
12	1050	$- 9.0$	128.2	6.3	40	1.33
13	1050	$- 9.0$	128.2	6.3	1000	1.33
14	1050	$- 9.0$	166.0	7.5	40	1.33
15	1050	$- 9.0$	166.0	7.5	1000	1.33
Pulse energy
16	1050	$- 9.0$	128.2	6.3	40	1.33
17	1050	$- 9.0$	128.2	6.3	1000	1.33
18	1050	$- 9.0$	128.2	6.3	80	0.67
19	1050	$- 9.0$	128.2	6.3	1000	0.67
Refractive correction
20	1050	$- 9.0$	136.3	6.5	40	1.33
21	1050	$- 6.0$	90.6	6.5	40	1.33
22	1050	$- 3.0$	45.0	6.5	40	1.33
23	1050	$+ 2.0$	32.0	6.5	40	1.33
24	1050	$+ 4.0$	69.6	6.5	40	1.33
25	1050	$+ 6.0$	111.6	6.5	40	1.33
26	1050	$- 9.0$	136.3	6.5	1000	1.33
27	1050	$- 6.0$	90.6	6.5	1000	1.33
28	1050	$- 3.0$	45.0	6.5	1000	1.33
29	1050	$+ 2.0$	32.0	6.5	1000	1.33
30	1050	$+ 4.0$	69.6	6.5	1000	1.33
31	1050	$+ 6.0$	111.6	6.5	1000	1.33
Effect of epithelium
32	1050	$- 9.0$	128.2	6.3	40	1.33
$33 a$	1050	$- 9.0$	128.2	6.3	40	1.33
$33 b$	1050	$- 9.0$	128.2	6.3	40	1.33
34	1050	$- 9.0$	128.2	6.3	1000	1.33
$35 a$	1050	$- 9.0$	128.2	6.3	1000	1.33
$35 b$	1050	$- 9.0$	128.2	6.3	1000	1.33

The customized AMARIS laser system used in our tests allowed for manipulation of ablation parameters like system repetition rates with or without the local ablation frequency controls. The system repetition rate of the laser system was set at 1050 Hz for most of the tests. The local frequency of 40 Hz in Table 1 represents the system repetition rate limited to the “intelligent thermal effect control” frequency³³ (normal iTEC frequency) of the AMARIS platform (i.e., system repetition rate limited to 40 Hz local ablation frequency), whereas the 1000-Hz local frequency represents essentially no local frequency for the 1050 Hz (and lower) system repetition rate, with almost every pulse (9 out of 10 pulses) delivered with 1050-Hz frequency and minimal pulses (1 out of 10 pulses) delivered with 525-Hz frequency. These local frequency settings are represented here with the terms “clinical local frequency” (for 40 Hz) and “no local frequency” (for 1000 Hz), respectively. The “no local frequency” of 1000 Hz (in Table 1) should be considered as the maximum allowed frequency of the laser system. In addition to the system repetition rate of 1050 Hz, we also tested the impact of two additional system repetition rates (500 Hz and 750 Hz) with and without the local frequency controls (lines 4 to 9 in Table 1). In these cases (cases 7 and 9 in Table 1), “no local frequency,” although set at 1000 Hz, actually means the highest achievable system repetition rate of the laser system (500 Hz and 750 Hz, respectively).

The change in local frequency did not affect the pulse energy in our test laser system. However, as a feature of our test software, under the influence of local frequency control, a change in laser pulse energy manipulated the local frequency in order to maintain a constant fluence received by the eye per unit of time. This control was deactivated for “no local frequency.” Therefore, in cases 16 and 18 (Table 1), manipulation of pulse energy resulted in a change in local frequency. However, in cases 17 and 19 “no local frequency” was maintained (i.e., maximum allowed system frequency) even though the pulse energy was manipulated. We discuss the implication of these features in later sections.

Collectively, 105 eyes received 111 ablations in total. For all the ablations, nonwavefront guided aspheric ablation profiles were calculated for a vertex distance of 12 mm and keratometric readings (K1 and K2) of 44 D using the SCHWIND custom ablation manager (SCHWIND eye-tech-solutions, Kleinostheim, Germany). All the ablations were based on a pseudorandom algorithm with the local frequency controls implemented as described elsewhere.³⁴

An IR thermal camera (FLIR-A 615) with a macro-objective was positioned at $10$ cm distance from the ablation focus and was tilted horizontally to an angle of $\sim 60 \deg$ . Every ablation was recorded at 200 fps in windowed mode ( $\sim 83 μ m / pixel$ ) 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 was also analyzed around the OZ. 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 0.9 in agreement to the literature.²⁰ With our test setup, we can expect a measurement error of $\pm 0.5 ° C$ due to the thermal camera.²⁰ 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 (Fig. 1) achieved during the video recordings. Two such example video recordings are presented here (Fig. 2 with clinical local frequency, Fig. 3 with no local frequency). 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. Figure 1 presents an example with clinical local frequency, with a baseline maximum temperature of 20.8°C, peak of the maximum temperature 25.8°C and change in peak corneal temperature 5°C.

Fig. 1

A screenshot from the video recording with the thermal camera for the first ablation performed with the clinical local frequency for a refractive correction of $- 9 D$ with 6.3-mm OZ. The baseline and peak corneal temperature were recorded at 20.7°C and 25.8°C, respectively.

Fig. 2

An example ablation recording with clinical local frequency (Video 1, MOV, 5.2 MB) [URL: http://dx.doi.org/10.1117/1.JBO.20.7.078001.1].

Fig. 3

An example ablation recording with no local frequency (Video 2, MOV, 2.93 MB) [URL: http://dx.doi.org/10.1117/1.JBO.20.7.078001.2].

Data analysis was performed using Microsoft Excel (Microsoft Corporation). Correlation and paired single tailed $t$ -tests were performed to analyze the differences among datasets with a 0.05 level of statistical significance. For curve fitting, we took an exploratory approach of maximizing the coefficient of determination ( $r^{2}$ values).

3. Results

The maximum baseline temperature across all the ablated eyes was on an average $20.5 ° C \pm 1.1$ (17.7°C to 22.2°C). The change in peak corneal temperature for the ablations performed with clinical local frequency was on an average $5.8 ° C \pm 0.8$ ( $p = 3.3 E - 53$ to baseline); whereas $9.0 ° C \pm 1.5$ was the average for the ablations performed with no local frequency ( $p = 1.8 E - 35$ to baseline, $1.6 E - 16$ to clinical local frequency ablations). We analyzed the impact of each of the varying parameters on the change in peak corneal temperature.

3.1.

Local Frequency

Three local frequency controls (10, 40, and 1000 Hz) were tested with a system repetition rate of 1050 Hz. The change in peak corneal temperature progressed logarithmically with respect to the increasing local frequencies (Fig. 4), with statistically significant differences among all the three tested local frequencies maximizing between clinical and no local frequency ( $p = 0.005$ ) (Table 2). For the clinical local frequency, the highest change in peak corneal temperature of 5.8°C was observed versus 11°C for no local frequency and 2.7°C for 10-Hz local frequency.

Table 2

Thermal response to changing local frequency; for system repetition rate 1050-Hz, pulse energy 1.33 mJ, and optical zone (OZ) size 6.3 mm.

Local frequency (Hz)	Average change in peak temperature (°C)
10	$2.0 \pm 0.6$
40	$5.6 \pm 0.3$
1000	$10.2 \pm 0.7$
$p$ -value	0.005

Fig. 4

The peak corneal temperature increased logarithmically with the increasing local frequency. However, for the clinical local frequency, the maximum change in peak corneal temperature was 5.8°C.

3.2.

Repetition Rate

Three different system repetition rates (500, 750, and 1050 Hz) were tested with clinical and no local frequency controls. For the clinical local frequency, the measurements at 500 and 750 Hz system repetition rate were similar ( $p = 0.3$ ) in comparison to 1050 Hz ( $p < 0.02$ ) (Table 3). However, for no local frequency, measurements at all the tested repetition rates differed significantly ( $p < 0.04$ ), increasing rapidly with the repetition rate (Fig. 5). The differences between the two local frequencies became more significant with increasing repetition rate ( $p = 0.003$ for 1050-Hz system frequency). For the clinical local frequency, the maximum change in peak corneal temperature was limited to 6.8°C for the entire range of system repetition rates, however, reaching as high as 11°C for no local frequency.

Table 3

Thermal response to changing repetition rate. For both local frequencies, significant differences were observed between repetition rates, compared to 1050-Hz repetition rate. For all repetition rates, measurements at clinical local frequency were significantly lower versus no local frequency, with differences becoming more significant with increasing repetition rate.

Repetition rate (Hz)	Clinical local frequency	No local frequency	p-value
500	$6.5 \pm 0.2$	$6.7 \pm 0.3$	0.2
750	$6.4 \pm 0.4$	$8.8 \pm 0.2$	0.003
1050	$5.6 \pm 0.3$	$10.2 \pm 0.7$	0.003
$p$ -value	$< 0.02$ (wrt 1050 Hz)	0.04	–

Fig. 5

The change in peak corneal temperature with respect to the system repetition rate. The differences between the two local frequencies became more significant with increasing repetition rate ( $p = 0.003$ for 1050-Hz repetition rate), peaking up to 6.8°C for the clinical local frequency.

3.3.

Optical Zone

Three different OZs (4, 6.3, and 7.5 mm) were tested with clinical and no local frequency controls. For both clinical and no local frequencies, the differences between all the three OZs were statistically significant ( $p < 0.02$ ) (Table 4). The paired $t$ -test shows higher significant differences between the two local frequencies for 4 mm ( $p = 0.0001$ ) and 7.5 mm ( $p = 0.0002$ ) OZ compared to 6.3 mm (0.003). However, the progression with respect to the OZ shows a converging trend between the two local frequencies (Fig. 6) for wider OZs, with the change in peak corneal temperature increasing with the OZ for clinical local frequency while it decreased for no local frequency. For the clinical local frequency, the highest change in peak corneal temperature was limited to 6.6°C for the entire range of OZ, however, it reached as high as 12.1°C for no local frequency.

Table 4

Thermal response to changing OZ size. For both local frequencies, significant differences were observed between the three OZs. The measurements at clinical local frequency were significantly lower compared to no local frequency for the entire range of tested OZ sizes.

OZ (mm)	Clinical local frequency	No local frequency	p-value
4	$4.4 \pm 0.1$	$12.0 \pm 0.1$	0.0001
6.3	$5.6 \pm 0.3$	$10.2 \pm 0.7$	0.003
7.5	$6.4 \pm 0.2$	$8.6 \pm 0.3$	0.0002
$p$ -value	0.02	0.02	–

Fig. 6

The change in peak corneal temperature with respect to the optical zone (OZ) sizes. A converging trend is observed between the two local frequencies with increasing OZ sizes. The maximum change in peak corneal temperature (6.6°C) with clinical local frequency was obtained for a 7.5-mm OZ.

3.4.

Pulse Energy

Two single pulse energy levels were tested (0.67 and 1.33 mJ). The clinical local frequency of 40 Hz was maintained for 1.33 mJ pulse energy, but for half of the pulse energy, the local frequency was increased to 80 Hz in accordance with the pulse energy control of the local frequency. Therefore, the two tested local frequencies are represented as controlled (80 Hz) and no local frequencies (1000 Hz) in Fig. 7. For both the local frequencies, the differences between the two pulse energy levels were statistically significant ( $p = 0.01$ for controlled local frequency and $p = 0.03$ for no local frequency, Table 5). The differences between the two local frequencies increased with increasing pulse energy (Fig. 7), with the change in peak corneal temperature decreasing for higher pulse energy with the controlled local frequency, and increasing for no local frequency. For the controlled local frequency, the highest change in peak corneal temperature was limited to 6.85°C for both the tested pulse energy settings, however, it reached as high as 11°C for no local frequency.

Table 5

Thermal response to changing pulse energy. The pulse energy levels showed significant differences for the two local frequencies. Between the two local frequencies the differences became more significant with increasing pulse energy.

Energy (mJ)	Clinical local frequency	No local frequency	p-value
0.67	$6.5 \pm 0.3$	$8.7 \pm 0.9$	0.003
1.33	$5.6 \pm 0.3$	$10.2 \pm 0.7$	0.01
$p$ -value	0.01	0.03	–

Fig. 7

The change in peak corneal temperature with respect to the pulse energy. The clinical local frequency of 40 Hz was maintained for 1.33-mJ pulse energy, but for half the pulse energy the local frequency was increased to 80 Hz. However, with the controlled local frequency of 80 Hz at 0.67-mJ pulse energy, the maximum change in peak corneal temperature reached up to 6.85°C.

3.5.

Refractive Correction

Myopic ( $- 9$ , $- 6$ , and $- 3 D$ ) and hyperopic ablations ( $+ 2$ , $+ 4$ , and $+ 6 D$ ) were performed with clinical and no local frequency each. The change in peak corneal temperature was significantly higher for myopic corrections in comparison to hyperopic corrections for both the local frequencies ( $p = 0.001$ and $p = 0.000009$ for clinical and no local frequency, respectively) (Table 6). The differences between the local frequencies were stable and significant for all the refractive corrections, favoring the clinical local frequency peaking at 6.1°C but reaching up to 9.4°C for no local frequency (Fig. 8).

Table 6

Thermal response to changing refractive correction. More significant differences showed between myopes and hyperopes for the no local frequency versus the clinical local frequency. For all refractive corrections, the clinical local frequency was favorable in terms of thermal control.

Refraction (D)	Ablation depth (μm)	Clinical local frequency	No local frequency	p-value
$- 9$	136.3	$5.8 \pm 0.2$	$9.2 \pm 0.3$	0.00008
$- 6$	90.6	$6.0 \pm 0.1$	$8.7 \pm 0.4$	0.001
$- 3$	45.0	$5.2 \pm 0.1$	$8.2 \pm 0.5$	0.004
$+ 2$	32.0	$4.4 \pm 0.3$	$7.3 \pm 0.6$	0.002
$+ 4$	69.6	$5.0 \pm 0.2$	$7.3 \pm 0.8$	0.01
$+ 6$	111.6	$5.2 \pm 0.4$	$6.7 \pm 0.5$	0.03
$p$ -value myopes versus hyperopes	–	0.001	0.000009	–

Fig. 8

The change in peak corneal temperature with respect to the ablation depth. Here M and H represent myopic and hyperopic corrections, respectively. Also the ablation depth of 136.3, 90.6, 45.0, 32.0, 69.6, and 111.6 (in $μ m$ ) correspond to refractive correction of $- 9$ , $- 6$ , $- 3$ , $+ 2$ , $+ 4$ , and $+ 6$ diopters, respectively. A significantly higher peak corneal temperature was observed for myopic corrections compared to hyperopic corrections, for both the local frequencies, peaking to 6.1°C for clinical local frequency.

3.6.

Effect of Epithelium

Transepithelial and stromal ablations were performed with both clinical and no local frequency. For both local frequencies, the change in peak corneal temperature was significantly higher for transepithelial ablations compared to stromal ablations (mean 0.9°C higher, $p < 0.007$ for clinical local frequency and mean 1.3°C higher, $p < 0.1$ for no local frequency). However, the change in peak corneal temperature was similar for the two stromal ablations ( $p = 0.3$ for clinical and $p = 0.1$ for no local frequency). Between the two local frequencies, the differences were significant for all corneal ablations ( $p < 0.003$ ) (Table 7).

Table 7

Role of epithelium in thermal control. For both local frequencies, the average change in peak corneal temperature was significantly higher for transepithelial ablations versus intrastromal ablations. The clinical local frequency was again favorable in terms of thermal control, in comparison to no local frequency, with the differences slightly increasing with deeper corneal ablations.

Ablation	Clinical local frequency	No local frequency	p-value
Transepithelial	$7.0 \pm 0.2$	$10.8 \pm 0.4$	0.003
1st stromal	$6.1 \pm 0.3$	$9.7 \pm 0.2$	0.002
2nd stromal	$6.0 \pm 0.1$	$9.3 \pm 0.2$	0.001
$p$ -value versus transepithelial	$< 0.007$	$< 0.01$	–

4. Discussion

The analysis of our results revealed the thermal response of the cornea to varying test parameters with and without the local frequency controls. In our experiments, we set −9 D refractive correction as our basis since nowadays it may be considered as a rather high correction. The clinical local frequency of 40 Hz was specified based on the available literature,³⁰ and is also being used as a standard value in the AMARIS platform. The no local frequency of 1000 Hz was specified since for a 1050-Hz system repetition rate, this represents no local frequency controls with almost every pulse (9 out of 10 pulses) delivered with 1050-Hz frequency and minimal pulses (1 out of 10 pulses) delivered with 525-Hz frequency. The system repetition rate of 1050 Hz is the highest available repetition rate among excimer laser systems, to the best of our knowledge. The OZ size of 6.3 mm may be considered as the standard clinical value in refractive surgery. A pulse energy level of 1.33 mJ was specified as a standard value for AMARIS high fluence. Paired single tailed $t$ -tests were preferred compared to other statistical tests, such that for each varying parameter, each tested value could be individually compared to all the other tested values. The specific varying parameter and its corresponding impact are discussed as follows: