## 1 Introduction  
In the early 1950s, light coagulation was introduced as an outstanding optical therapeutic tool for remote tissue manipulation to prevent retinal detachment after retinal break formation [1]. The resulting technique is known today as laser photocoagulation (LPC) and is widely used as a treatment modality. In addition to its application for retinal detachment, LPC was found to be beneficial for ablative treatment of diabetic retinopathy [2],diabetic macular edema [3], central serous retinopathy [4] and other retinal pathologies. In LPC, millisecond laser-pulsed irradiation is absorbed by tissue pigments and converted to heat, causing photothermal denaturation. The delicate multi-layer retinal structure, consisting of the retinal pigment epithelium (RPE), Bruch's membrane, choriocapillaris and the otherwise healthy overlying neuroretina, including the sensitive photoreceptor cells, are collaterally damaged by laser pulses in the millisecond range, due to the relatively slow heating process and the relatively fast heat dissipation into the directly unheated neighboring tissue. This indirect heating resulting from excessive power in turn leads to scotoma, reduced night vision and disruption of the retinal anatomy through scarring [5; 6]. Therefore, it is not possible to use LPC near the fovea because of the high risk of permanent loss of central vision [7; 8]. These disadvantages have led to the development of new laser treatment modalities, which aim to produce specific therapeutic results without damaging photoreceptors and the neural retina by selectively acting on target structures. Among these treatment techniques, selective retina therapy (SRT), which targets only the prismatic retinal pigment epithelial cell monolayer, is particularly suitable [9]. The treatment is intended to selectively destroy a portion of defective RPE cells and utilize the regenerative abilities of the epithelium to recover its original structure. The targeted apoptotic RPE cells are cleared by phagocytosis, while extracellular debris is cleared by macrophages. Migration and proliferation of healthy neighboring RPE cells follow [10; 11]. The stimulation of RPE cell migration is hypothesized to finally lead to improved metabolism at the target sites [9], which has been reported to have beneficial effects for patients suffering from RPE-related retinal diseases such as central serous chorioretinopathy [12]. The basis for this selective RPE damage is provided by the intracellular melanosomes, which despite the <10 \(\upmu\)m thickness of the RPE monolayer [13], absorb about 50% of the incident light in the green spectral range. Within the thermal confinement of SRT, peak temperatures initiate microsecond-lifetime microbubble formation (MBF) on the melanosome's surface [14]. The rapid mechanical expansion and collapse of these microbubbles causes RPE cell-wall disruption, followed by immediate or delayed cell death [15]. Nevertheless, in practice, achieving selective RPE cell damage without damaging the surrounding tissue via heat diffusion is challenging due to the RPE's strongly varying melanin concentration. A successful SRT application is essentially based on the following two factors: first, the applied laser pulse duration, and second, accurate real-time energy dosing that depends on radiant exposure and local RPE absorptivity. Regarding pulse durations, different experiments have shown that RPE cell death coincides with MBF for laser pulse durations up to 50 \(\upmu\)s [16; 17]. [PERSON] et al. were able to show that, in the case of 1, 5 and 10 \(\upmu\)s laser pulse duration, more than 95% of dead cells were associated with MBF, whereas the ratio decreased to 65% and 45% for longer 20 and 40 \(\upmu\)s pulses, respectively [18]. This suggests that pulse durations longer than 20 \(\upmu\)s delineate the limit of thermal confinement for SRT; beyond this, the technique appears impractical even with the use of real-time dosimetry.