Introduction and Background
Retinal laser photocoagulation is a widely used treatment for peripheral ischemic retinal diseases such as diabetic retinopathy, ischemia due to vein occlusion and other conditions,12.3.–4 as well as for treating diabetic macular edema,5 either on its own or in combination with pharmacological agents. Since the introduction of antivascular endothelial growth factor (anti-VEGF) therapies, laser photocoagulation has been challenged as the standard treatment, especially for diabetic edema, but still remains a relevant alternative for this indication.
So far, the treating physician has to choose the exposure time and laser beam diameter based on the treatment indication and the laser power according to the ophthalmoscopic evaluation of the appearance of the lesions after treatment. The method of treatment as introduced by Meyer-Schwickerath6 has become a standard procedure with the development and application of -lasers around two decades later.7 The laser light, typically in the green spectral range, is absorbed mainly in the retinal pigment epithelium (RPE) and the choroid. Due to heat diffusion, the adjacent retinal layers are thermally denatured, and the result becomes apparent as whitish lesions. Recent trends in the development of this modality have focused on reducing the treatment time. It has been shown that short irradiation times of about 20 ms can lead to the reduction of axial tissue damage8 and reduced pain for the patient during treatment.9 This and the desire for a more efficient treatment has led to the development of automated pattern photocoagulation, where each individual lesion is created within a short timeframe,10 typically within 10 to 20 ms. However, the likelihood for overtreatment increases with decreasing treatment duration, due to the reduction of the therapeutic window,11 which describes the range of laser powers that lead to visible lesions without causing rupture. This reduction of the therapeutic window can be problematic due to local pigmentation variations up to a factor of three in the absorbing layers12,13 as well as age-related variations of the light transmittance through the ocular media,14 treatments with the same laser parameters can have a wide variety of outcomes, in extreme cases unnecessarily strong lesions, rupture, and bleeding.
Depending on the indication, the treatment endpoint ranges from small and gentle lesions for diabetic macula edema, typically with lesion sizes around 50 to to more pronounced lesions for diabetic retinopathy with lesion diameters of up to . However, the mechanism of the therapeutic effect of the treatment is still not fully understood. Hence, the required minimal lesion strength for a positive therapeutic outcome is unknown. This is why researchers have begun studying the dosage threshold of irradiations for clinical viability15,16 and started investigating the efficacy of low-dose treatments.1718.–19
Reducing the risk of overly damaging treatment outcomes and reducing pain are major goals in improving the reliability of photocoagulation therapy. With shorter irradiation times, however, this increases the need for an objective measurement of the denaturation effect and a real-time laser exposure control.20 Several ideas for reliably creating subvisible or mild treatment effects have been engineered and tested over the last decade.21,22 One approach (“endpoint management”) uses titration lesions to determine the laser parameters that lead to a barely visible lesion and uses a mathematical model based on the Arrhenius-theory to adjust the therapy laser power and time for the application of the nondamaging retinal therapy. A clinical trial on patients with central serous retinopathy using this method has shown the clinical efficacy of subvisible laser exposures.23
Another approach uses the consecutive application of laser pulse trains in the regime. The treating physician chooses the appropriate laser parameters for the “micropulse” treatment, either from experience or after creating a test burn outside the treatment area.24
Both of these mentioned methods rely on predicting the extent of lesions due to a titration process. However, without direct feedback from the tissue, these methods are subject to variations in treatment outcome, due to the variation of the absorption and the associated laser-induced temperature rise variability.
In order to utilize site-individual feedback and automatic dosimetry for the treatment, a temperature-based method using the optoacoustic effect has been introduced.25 Optoacoustics provides a noninvasive real-time monitoring of tissue temperature during photocoagulation. Originally developed for use during transpupillary thermotherapy,26 the feasibility of the temperature rise measurement was demonstrated on porcine fundus explants,27 and recently the first temperature monitoring on patients during standard photocoagulation was showcased.28 The validity of the temperature measurement was shown using fundus images and optical coherence tomography (OCT).29 Using the optoacoustic temperature measurement, a feedback algorithm was developed. The lesion strength depends on the temperature rise and the treatment time, so the method monitors both during the irradiation. Once the threshold temperature for the desired lesion strength at the elapsed time is reached, the laser is switched off automatically. This dosimetry has been successfully demonstrated in vivo in rabbits.25 It was achieved by monitoring the temperature increase of the retinal tissue during the photocoagulation at a rate of 1 kHz and automatically stopping the treatment according to a characteristic curve that determines the relation between the temperature and irradiation time. It was shown that the developed algorithm was able to create ophthalmoscopically uniform lesion sizes independent of the laser power, transmission and degree of pigmentation, and showing via histology that the ganglion and nerve fiber layers were reliably unaffected.30 The creation of lesions of different strengths was also demonstrated on rabbits in vivo.31,32
In this work, we present an alternative approach for a temperature-based feedback control, which adjusts the therapy laser power during the irradiation, in order to achieve a preset temperature rise within a fixed time frame.
Optoacoustic Temperature Measurement
The temperature measurement is based on the optoacoustic effect. The temperature-dependent pressure generation is described thoroughly by Brinkmann et al.28 If a laser pulse with pulse duration is applied to an absorber, the absorbed energy is converted to heat and the temperature of the absorber rises by33,34 33 For retinal irradiations, the pressure wave propagates through the eye globe and its amplitude can be detected at the cornea. Due to absorption and acoustic impedance mismatches of the eye constituents the amplitude is attenuated, however, the detected pressure wave amplitude is proportional to . Due to the temperature dependency of , the detected amplitude of the pressure wave contains information on the base temperature of the irradiated tissue . We can express using a second-order polynomial approximation 26
and are the tissue specific parameters that were determined for rabbits by a calibration measurement of the optoacoustical signal for known temperatures26 as and .
The proportionality constant is unique for each irradiation site and contains information of the pigmentation at each irradiated area, but also on the light and sound transmission through the eye as well as the detector characteristics. It is, therefore, required to determine for each individual irradiation site prior to the application of the therapy laser radiation. If one measures the pressure at a known temperature , in this case the body temperature during in vivo experiments, Eq. (5) can be solved for . This step allows us to take the site-individual variations in signal into account.
Power-Controlled Retinal Photocoagulation
In order to perform irradiations within a set time frame with a fixed desired end temperature , the laser power during the photocoagulation is adjusted depending on the site-individual temperature information. This is shown in Fig. 1.
Each irradiation consists of three phases. In the calibration phase of duration , calibration pulses are applied to the unheated fundus tissue in order to determine the irradiation site specific parameter . This calibrates the feedback signal to the base temperature of the absorber before the start of the treatment (here, the body temperature of the rabbit for the in vivo experiments). After the calibration, an irradiation at a low power is performed for an irradiation time , during which the temperature rise is monitored. The knowledge of and can be used to calculate the required power to reach the desired temperature rise . The tissue is irradiated with the calculated treatment laser power and the irradiation time , and the temperature reaches in order to create the desired lesion strength at the end of the irradiation.
Temperature Profile During Retinal Photocoagulation
When heating the rabbit fundus, the radiation is absorbed in the monolayer melanin granules of the RPE and the pigmented part and blood of the choroid. As the rabbit is a merangiotic animal, the blood vessels are confined to these regions and the absorption in the inner retina can be neglected. A model for rabbit eyes with absorption coefficients of and with thicknesses 4 (absorbing part of the RPE) and (absorbing part of the choroid), respectively,35 was used for the calculations. According to the heat diffusion equation28 with the area of irradiation.
However, is not a suitable parameter to describe the coagulation process. The peak RPE surface temperature of the irradiated volume , which is located at the center of the irradiated area close to the surface of the RPE, provides a better description, as it determines the onset of the retinal coagulation. is correlated to via a time-dependent conversion function
The conversion function and its properties are described in detail by Schlott et al.25 and the radial and axial temperature profiles are more thoroughly investigated by Brinkmann et al.28 Hence, only the adjustments that are necessary for this method are described here.
For the power-controlled irradiation, the conversion function is not just one continuous function but has to be determined for the probe phase and the treatment phase independently. We used a mathematical model to derive for the two irradiation phases at the rabbit fundus. The commercially available software COMSOL Multiphysics 4.2 is a finite-element numerical simulation software, which was used here to solve the heat diffusion equation [Eq. (6)]. The light–tissue interaction is not modeled directly but rather treated as a heat source , which describes the laser energy deposition per time and volume in and can be described within the irradiated volume during the irradiation as34 and the absorption coefficients according to Hammer35 as shown in Table 1.
Layer thicknesses and absorption coefficients of the rabbit fundus as used in the numerical model.
|Layer||Thickness (μm)||Absorption coefficient (cm−1)|
|Pigmented part of RPE||4||1204|
|Unpigmented part of RPE/choroid||37||0|
|Absorbing part of choroid||20||260|
Light scattering is not incorporated in the model. The geometry is more thoroughly discussed by Schlott et al.25 and shown in Fig. 2(a). The origin is at the center of the square irradiated area at the surface of the RPE. The temperature rise is shown for the end of the treatment at with a calibration power (which corresponds to a pulse energy of 3 µJ applied with a repetition rate of 1 MHz), a probe power and a treatment laser power in a vertical -slice through the fundus layers, where is one of the radial coordinates and is the axial coordinate with the upper edge of the RPE located at . The axial temperature profile development over time (b) as well as the profile in -direction (c) is also shown.
In the model, the remaining tissue parameters are assumed to be that of water [; ; ]. The temporal temperature development of and as calculated from the temperature profiles is shown in Fig. 3.
For the experiments, we used a fiber with a square core, so the integration in Eq. (7) was carried out in carthesian coordinates. This also necessitates the simulation being run in three-dimensions, because there is no radial symmetry. The simulation parameters were chosen to be which corresponds to a pulse energy of 3 µJ applied with a repetition rate of 1 MHz, , and . By taking the ratio of and according to Eq. (9), one can determine the conversion function . We split the conversion function into two parts, and , corresponding to the probe and treatment phase, respectively.
From Fig. 1, the ratio of powers can be correlated to the ratio of temperatures . If the tissue was allowed to cool down to prior to applying , then , because the temperature rise is proportional to the introduced heat.36 However, if the temperature is not completely cooled down when the treatment phase starts, becomes a function of the spot size used, the time of the probe phase and the time gap to the start of the treatment phase. Therefore, we used the mathematical model to calculate the relationship between the ratios. The temperature rise from the two successive phases was simulated by varying the laser powers and for different beam diameters and calculating the temperatures. By plotting over , we found a linear correlation that could be fitted according to , where is the slope of the curve and is the intersection with the temperature axis. We found for all diameters, but the intersection was radius-dependent, as is shown in Fig. 4. The simulated data were fitted with a parabola.
The powers and consist of the laser power of the therapy beam and the mean power introduced by the measurement laser pulses (, ). remains the same for the successive irradiations (). Using this and solving for yields
is the power that is required during the treatment phase in order to reach at the end of the treatment phase. Of particular interest is the conversion function for different ratios of laser powers . This is depicted in Fig. 5. After calculating , the temperature is monitored using the corresponding conversion function based on .
Materials and Methods
The experimental setup is shown in Fig. 6. The therapy laser is a frequency doubled Nd:YAG (Visulas 532 s, Carl Zeiss Meditec AG, , cw). After starting the irradiation via a footswitch, the PC opens the shutter that blocks the measurement laser radiation (Uniblitz LS6, aperture 6 mm, time to open: 0.7 ms). The temperature measurement is performed using light from a frequency doubled Q-switched Nd:YLF laser (QC-523-1000, CrystaLaser, , , ), which emits one pulse every millisecond. The radiation is coupled via a microscope objective into a multimode fiber (square core with edge length , ). The fiber is connected to a laser slit lamp (LSL 532s, Carl Zeiss Meditec AG), which guides the light through a contact lens. The lens is a modified commercial lens (Mainster Focal Grid, OMRA-S, Oculus Instruments) with a laser spot magnification of 1.05 into which an annular piezoceramic ultrasonic transducer with a center frequency of and a bandwidth of 410 kHz was embedded to detect the incident pressure waves. The light is then focused onto the rabbit fundus for the irradiation. The measurement light induces the optoacoustic ultrasound waves from the retinal layers, which travel through the ocular media and reach the contact lens with the embedded transducer, where they are converted into electrical signals that are amplified and recorded via a digitizer card (CompuScope 8347, Gage Applied Technologies). The data processing and feedback control are performed by a custom made LabVIEW routine. The measured pressure provides the temperature information according to Eq. (5).
The light from the therapy laser after opening the shutter (Uniblitz LS6) is superimposed with the light from the measurement laser and both are coupled into the fiber. This way the measurement irradiation is guaranteed to irradiate the same volume as the therapy irradiation. The laser power is controlled using an input/output card (NI PCIe-6341, National Instruments) to perform the treatment during the treatment phase.
The experiments on chinchilla gray rabbits in vivo were carried out according to the German law for protection of animals, approved by the Ministry of Agriculture, the Environment and Rural Areas of Schleswig-Holstein, Kiel, Germany. Prior to the irradiation, the rabbits were anesthetized using ketamine (10%, 0.5 to ) and xylazine (2%, 0.2 to ).
After the animal was narcotized, it was fixed in a custom-made holding device in front of the slit lamp. The modified contact lens was placed on the rabbit’s eye and mechanically fixed. In order to adapt the acoustical and optical properties of the contact lens and cornea, methylcellulose gel (2%) was used as a contact gel. The body temperature was measured rectally. The beam edge length was set to on the slit lamp, which results in a beam edge length of on the rabbit fundus. The calibration phase time was chosen to be and the probe and treatment phase times were both set to , bringing the resulting total irradiation time per site to 50 ms. During the calibration phase, the shutter for the therapy laser remains closed and was opened at the beginning of the probe phase. The therapy laser power during the probe phase was chosen so that at the end of the probe phase was . This ensured that the temperature rise during the probe phase would not result in irreversible coagulation effects of the tissue, while at the same time ensuring that the temperature rise is above noise level and can be fitted well. The therapy laser radiation heats the fundus tissue and the resulting temperature curve during the probe phase is fitted with a curve as described by Schlott et al.25 to determine , which is then used to determine the required laser power during the treatment phase . The calculation of and the adjustment of the therapy laser can be carried out within during the irradiation. The aim temperatures were chosen as 50°C, 55°C, 60°C, 65°C, 70°C, and 75°C. For the first five rabbits (group 1), only the aim temperatures of 50°C, 55°C, 65°C, and 75°C were used in order to find the ED50 value for ophthalmoscopic visibility. Two additional rabbits (group 2) were irradiated with the aim temperatures 60°C, 65°C, 70°C, and 75°C to resolve the ED50 value more thoroughly. The measurement pulse energy was adjusted between 3 and , depending on the quality of the pressure signal. The measurement pulse energy has to be high enough in order to create a strong photoacoustic signal, but at the same time low enough to prevent it from noticeably heating the tissue during the calibration phase. While the exposure to the measurement pulses in this case is above the threshold values for the maximum permissible exposure, the applied energy does not lead to a denaturation of the tissue that would influence the measurement of the temperature. Regardless, the mean power of the pulses has to be considered for the following therapy as shown in Eq. (12).
In order to evaluate the visibility of the lesions, fundus images were taken 1 h after the treatment using a fundus camera (VISUCAM®, Carl Zeiss Meditec AG). The visibility of the lesion was judged from these fundus images by four independent observers. The lesion was classified as “visible” (1), if three of four observers agreed on seeing the lesion and as “invisible” (0) otherwise.
Accuracy of Aim Temperatures
The distribution of the final treatment temperatures and the corresponding irradiation powers are shown in Table 2 and displayed with their standard deviations and , respectively, in Fig. 7, as well as the standard error for . The end temperatures of irradiations could be evaluated. To better compare the standard deviations of the temperature and the laser powers, the relative deviation for both is given in %. For the temperature, one has to consider the error on the temperature rise in accordance to Eq. (12), so the relative error is, therefore, given as with . There is no such offset for the power, so is simply . The dashed line in Fig. 7 represents the bisecting line through the origin for the temperature axis, that is, the line on which the measured temperatures should fall.
Aim temperatures, measured treatment temperatures, corresponding treatment powers, and number of irradiations for the power-controlled coagulations.
|Taim (°C)||Ttreatment±σT (°C)||σrel,T (%)||Ptreatment,th±σP (mW)||σrel,P (%)||N||Standard error (°C)|
The mean measured temperatures (dots) agree very well with the target values (dashed line). The standard deviation of the temperatures increases with higher aim temperatures. For all temperatures, the corresponding treatment laser powers (triangles) show a larger standard deviation than the temperatures (about twice the relative deviation), indicating that the algorithm used laser powers over a wide range to reach the desired end temperature.
Visibility Threshold of Irradiations
Figure 8 shows a fundus image of the power-controlled irradiation sites for the different aim temperatures taken 1 h after the end of the irradiation for one of the rabbits from group 1, Fig. 9 shows the same for one of the rabbits from group 2. The images show the appearance of the irradiation sites for the different aim temperatures.
Visible lesions appear as white areas due to the increased scattering of the coagulated tissue. The bright white lesions marked with an “M” are marker lesions that are used for orientation. Within a square that is formed by four marker lesion areas, nine study irradiations with the same were performed.
The irradiation sites for and 55°C almost exclusively show no visible coagulation effects and the irradiation sites for 60°C and 65°C are in the transition regime between invisible and visible effects, where the 60°C mostly remain invisible and the majority of the 65°C lesions become visible as mild coagulations. The irradiation sites for 70°C and 75°C are almost exclusively visible and are more pronounced than the lesions for 60°C and 65°C. The intensity of those lesions is compatible with the current clinical treatment recommendation for panretinal photocoagulation treatment.
The visibility for each is presented in Table 3. The number of irradiation areas where the visibility could be evaluated is compared to the number of visible irradiation outcomes for each aim temperature . The fraction of these two values is given in %.
Number of evaluable irradiation areas Nirr and the number of visible lesions among them for each Taim.
|% visible lesions||0||1.5||25.7||42.5||89.4||97.1|
The evaluation of the visibility of the lesions from the fundus images and the corresponding temperatures can be used to find the ED50 value for ophthalmoscopical visibility, which is the temperature at which there is a 50% probability that the lesion becomes visible. The value is determined by performing a probit analysis of the binary data (visible and invisible) of the lesions. The result can be seen in Fig. 10, the visible lesions are indicated as black crosses in the top row, and the invisible lesions are at the bottom of the graph.
A total of 920 lesions in 11 eyes of 7 rabbits was evaluated. The number is lower than the number of irradiated sites, as not all areas could be imaged with the camera. Additionally, we omitted regions where image quality was too poor to give reliable results. The probit curve is displayed as the solid line, the ED50 value was calculated to be 65.3°C with a 95% confidence interval (indicated by the dashed lines). As the total number of visible and invisible lesions is difficult to assess from the graph, as the crosses conceal each other, the density of the visibility is shown for distinct temperature intervals (bars). The bar graph shows the normalized visibility per temperature interval, i.e., the number of visible lesions divided by the total number of irradiations in a temperature interval , where is the temperature the bar is centered on, with a temperature range of . For example, the bar centered at represents the fraction of visible lesions among the total irradiations, where was measured between 57.5°C and 62.5°C. The numerical values for these visibility fractions for the temperature intervals of 50°C to 80°C are shown in Table 4.
Total number and fraction of visible lesions in a temperature interval of 5°C centered on Tc.
|% visible lesions||0||0||21.3||42.4||86.7||96.4||100.0|
Accuracy of the Power Controlled Temperature Dosimetry
Figures 8 and 9 show the fundus image of the irradiations for different aim temperatures. Most of the irradiations with the aim temperatures of 50°C and 55°C remained subvisible (261 out of 264), despite a wide range of treatment powers . This shows that the site-individual absorption was accounted for and provides a reliable option to perform treatments with subvisible outcome. The breakdown into each category is shown in Table 3. In the transition regime with and 65°C, about 40% of the lesions become visible and are visibly less pronounced than the lesions created with and 75°C, which are visible in of the cases.
The mean value for the end temperature for the irradiations with , 57°C, and 65°C is very close to the desired aim temperature, but the standard deviation increases.
For all temperatures, the corresponding treatment laser powers show a larger standard deviation (about twice the relative deviation) than the temperatures, so the site-individual temperature feedback is crucial for a reliable dosimetry.
The subvisible irradiations achieved the highest temperature accuracy regarding the absolute temperature values. Looking at the temperature distribution of the different power-controlled lesions as shown in Fig. 7, the subvisible irradiations showed the least deviation in achieved temperature with a standard deviation of 2.1°C () and 3.8°C (). The standard deviation increases with the aim temperature , though it does not exceed 5°C. The increase is probably mostly caused by the propagation of errors during the determination of the probe phase temperature . The broader the gap between and , the more an error in the determination of the temperature is exacerbated for the treatment irradiation. Additionally, the temperature determination is most accurate close to the point of reference that is determined during the calibration phase.
One thing to consider specifically for higher temperatures already causing tissue denaturation is the increased light scattering. In this case, less probe and treatment laser light reaches the RPE. Furthermore, the tissue parameters and thus the Grüneisen-parameter [Eq. (3)] change. As a consequence, the temperature measurement loses its accuracy and deviations need to be taken into account, which become stronger the more the denaturation zone moves toward the inner retina.
The time controlled dosimetry introduced by Schlott et al.25 and Koinzer et al.30 was shown to be effective in creating uniform lesions across a wide range of irradiation laser powers. However, with their approach, the irradiation time varies depending on the laser power and the degree of pigmentation from a few milliseconds to well over 500 ms in some cases. In rare occurrences, the temperature threshold for ophthalmoscopic visibility could not be reached at all, regardless of the irradiation time. This is due to the fundus tissue reaching a thermal equilibrium, where the outflux of the heat is equal to the heat induced by the absorption of the laser light. Irradiations that are too short lead to an inaccurate determination of the temperature due to the increased influence of noise in combination with the limited number of temperature data points. Also, there is a technical limitation to the response time of the algorithm at about 4 ms. On the other hand, if the irradiation time is too long, the temperature measurement is more likely to be subject to errors due to eye movement. The variation of the irradiation time would make this approach unsuitable for automated pattern photocoagulation, where the individual irradiation has to be completed within 10 to 50 ms. Also it was shown that while the lesions are of the same diameter, they can differ in appearance for different irradiation times.30
With the demonstrated approach here, there was no instance of not being able to reach the aim temperature up to 75°C, improving the reliability of the dosimetry when compared to the time-variable approach. Additionally, as the treatment time is fixed at 50 ms, the appearance of the lesions is less prone to variations. The reliable time frame also makes the power-controlled dosimetry approach suitable to combine with some of the pattern coagulator systems operating in this time regime. The efficacy for shorter durations would have to be examined for application in pattern coagulators that operate in the 20-ms regime.
We have found that the most critical phase during the irradiation is the probe phase, where only a slight increase in temperature is desired. It is imperative that the temperature rise above the start temperature is high enough to be fitted reliably. Several steps were taken to increase the accuracy of the temperature determination of the probe phase.
The transducer signal carries not only the pressure data, but also noise and disturbances influence the pressure value determination. It is, therefore, important to match the transducer properties to the expected signal. In order to minimize the effect, the transient curves are filtered using a bandpass filter. Based on simulations realizing the photoacoustic signal generation for an irradiation with a diameter of and a pulse duration of 75 ns, we expect a central frequency of the transient pressure wave of 1.2 MHz. The transducers used in this work, therefore, have a resonance frequency at with a bandwidth of 410 kHZ. According to these considerations, the bandpass filter’s upper and lower bounds were chosen to be 400 kHz and 1.5 MHz, respectively.
Additionally, one has to consider the irradiation time. We found that 20 ms is a suitable irradiation time for . This results in 20 data points to fit the temperature curves, which proved to be sufficient to achieve a reliable temperature assessment in order to calculate the power needed for the treatment phase . However, probe repetition rate can be increased for higher accuracy.
The presented results are derived using a model for the rabbit fundus. When adapting the technique for the human fundus, some changes to the model are necessary. The different structure of the rabbit and the human eyes requires some alteration of the layer structure used for the determination of the conversion function. Additionally, the holangiotic nature of the retina requires the consideration of the absorption of laser light in the inner retina.
The first visible lesions appear at a temperature of about 55°C, albeit very rarely. For our treatment time of 20 ms, the ED50 temperature value of creating a lesion is at 65.3°C. This is in good accordance with results published earlier,25 where an ED50 value of 61.7°C for 20 ms can be extrapolated. However, there is no data for irradiation times of 20 ms to directly compare it to, as the previous experiments usually had longer irradiation times. Additionally, the fit curve derived by Schlott et al. seems to underestimate the temperatures for lower irradiation times, indicating that the ED50 value for this particular irradiation time agrees very well with the data presented here. Denton et al. have determined the threshold temperature for RPE damage as 53°C in the time range of 100 ms to 1 s, which is considerably lower than the ED50 value found here. However, their experiments were performed on RPE cell culture and with longer irradiation times.37 According to the Arrhenius theory, one would expect the threshold temperature for shorter irradiation times to be higher than for longer ones. Additionally, the threshold for RPE damage is lower than for ophthalmoscopically visible effects, because the heat has to be conducted into to retinal layers in order to denature the tissue.
The comparison between Tables 3 and 4 shows that the fraction of visible lesions in a temperature regime is very similar to the fraction of visible lesions for a given aim temperature. This shows that the desired coagulation effect is achieved by the method.
The ED50 value found here also compares quite well with the damage threshold temperature of 63°C found by Sramek et al.11 for the peak temperature for 20-ms irradiations in rabbits. Those temperatures were not measured directly, but inferred via a mathematical model based on the Arrhenius damage model.
When looking at the visibility of the lesions in Fig. 10, there is an overlap in visible and invisible irradiations between 60°C and 70°C. The reason for this might be uncertainties in the determination of the temperature. The conversion function to determine the peak temperature from the measured mean temperature assumes the ratio of the absorption coefficients of the RPE and the choroid to be fixed. However, it was shown that the conversion function, and therefore, the laser-induced peak temperature can vary up to 9% if the ratio of absorption coefficient changes by a factor of two.38 Also there are different tissue effects, for example, the creation of edema, which lead to a whitening of the fundus area, which cannot be distinguished from thermal denaturation on the fundus image. Thus, ophthalmoscopic visibility after 1 h might not be the best criterion to judge the treatment outcome. We are currently investigating the visibility thresholds of alternate modalities, such as fluorescein angiography and OCT, based on a classification system published by Koinzer et al.39 Finally, it is also conceivable that there is a variance of the biological reaction to identical heat stimuli caused, for example, by a locally varying blood vessel distribution in the choroid.
It was shown that it is possible to reproducibly apply irradiations with subvisible and visible effects within the total time frame of 50 ms by automatic adjustment of the laser power once during the course of each irradiation. The precision was 3.5°C (21.2%) in the subvisible regime. An irradiation scheme was devised, with 10 system calibration pulses per spot, followed by 20 ms of low laser power irradiation with a moderate temperature increase for absorption balancing, which is used to determine the therapy laser power for the last 20 ms in order to reach the desired final temperature increase. This method with a 1-kHz optoacoustic temperature probing was tested on rabbits in vivo for aim temperatures ranging from 50°C to 75°C. The irradiations at 50°C remained ophthalmoscopically subvisible after 1 h, and the standard deviation for the achieved temperatures ranged from 2.1°C for an aim temperature of 50°C to 5°C for 70°C. A comparison with the used laser powers shows that the same treatment temperatures can be achieved with a wide range of irradiation powers, indicating that irradiation-site-individual absorption levels are taken into account. The ED50 value for ophthalmoscopical visibility was found to be 65.3°C after 1 h, which is well supported by the literature.
This method can be well combined with automated pattern coagulators to improve the accuracy of these approaches and also to study the effects of subvisible irradiations in more detail. The next step for the presented method will be clinical trials to test the efficacy in practice and develop it toward a tool allowing for an objective and standardized examination of subvisible treatments.
R. Brinkmann is an inventor of patents related to the method of temperature measurements, which are owned by Medizinisches Laserzentrum Lübeck GmbH. The data reported in this study were acquired and processed by coauthors unaffiliated with any commercial entity.
This work was funded by the German Federal Ministry of Education and Research (No. 13GW0043B).
Alexander Baade studied physics at Johannes Gutenberg University in Mainz, Germany, with an emphasis on quantum optics and computer physics. He joined the Medical Laser Center Lübeck in 2009 with a research focus on optoacoustic temperature measurements during ophthalmologic procedures. He is currently working for the FARO Scanner Production GmbH in Stuttgart.
Claus von der Burchard studied medicine in Kiel and Göttingen and graduated in 2014. Since then, he has worked as a resident and research fellow at the Department of Ophthalmology, University of Kiel. His research focusses on retinal diagnostics and subthreshold laser treatment of retinal disorders.
Meike Lawin studied optronics majoring photonics at the Technical University of Ilmenau, Germany, and graduated in 2014. Afterwards she has been researching in the field of optoacoustics at the Medical Laser Center of Lübeck until 2016. She is currently working in the field of medical engineering at the CeramOptec GmbH in Bonn.
Stefan Koinzer graduated with an MD degree from the University of Kiel, Germany, in 2004. He did his internship in Offenburg and Kiel and completed his specialty training in ophthalmology in 2009. Since then he has been working as a research fellow for the Department of Ophthalmology at the University of Kiel. His research has focused on RPE, retinal laser photocoagulation and ocular histopathology. He works now in private practice.
Benedikt Schmarbeck is an engineer at the Medical Laser Center in Lübeck. He has held a BSc degree in medical engineering science from the Universität zu Lübeck since 2013. In 2015, he received his MSc degree in medical engineering science from the Universität zu Lübeck. His current fields of work are biomedical imaging, interferometry, photoacoustic measurement techniques, retinal photocoagulation, and tissue optics.
Kerstin Schlott studied experimental physics at the University of Hamburg, Germany, and graduated in 2007. Since then she has been working on optoacoustic temperature measurement and automatically controlled retinal laser coagulation at the Institute of Biomedical Optics in Lübeck. Since 2014 she has been researching optoacoustic detection of laser-induced microbubbles in the pigment epithelium at the Medical Laser Center Lübeck.
Yoko Miura studied medicine in Osaka, Japan, and received a medical license in 1997 and her PhD in 2002. After working as an assistant professor in the Department of Ophthalmology at Osaka City University till 2006, she undertook basic research at the university eye clinic in Kiel. In 2009, she moved to Lübeck, received a German medical license in 2012, and today is a faculty member at the Institute of Biomedical Optics, University of Luebeck.
Johann Roider, MD, is a professor of ophthalmology. He studied mathematics, physics, and medicine at Munich universities, Germany. He did a research fellowship at the Wellman Laboratories of Photomedicine, Boston. His clinical training was obtained at the universities of Lübeck and Regensburg. Since 2002 he has been the chairman of the Department of Ophthalmology, University of Kiel. His research focuses are laser–tissue interactions and vitreo-retinal disorders.
Reginald Birngruber studied electrical engineering and physics and graduated in physics (PhD) and in medical biophysics (MD habilis) in Germany. After his retirement from CEO and CRO of the Medical Laser Center Lübeck and the chair of the Institute of Biomedical Optics, he now makes his expertise and contacts available to colleagues in Lübeck and Boston. Current activities include mechanisms of laser tissue effects, optical tissue diagnostics, laser therapy control and new therapeutic techniques.
Ralf Brinkmann studied physics at the University of Hannover, Germany, with a focus on quantum optics and lasers. After a 5-year industrial interim period, he joined the Medical Laser Center in Lübeck, Germany, in 1993, and received his PhD from the University of Lübeck. Since 2005 he has been holding a permanent position as a faculty member at the University Institute of Biomedical Optics, and is also leading the Medical Laser Center as CEO.