Laser photocoagulation of the retina has been practiced for more than 30 years. Its therapeutic success is undisputed, being founded on many studies for different retinal diseases. The uses of the laser are both therapeutic (for example diabetes, thromboses of the eye, age-related macula degeneration (AMD)) and preventive (for example retinopexy). For photocoagulation, according to the type of disease, between a few single expositions in the macula and up to 3000 foci are applied in the case of pan-retinal photocoagulation. According to the spot size, 100-500 mW laser power is applied within 50-300 ms per focus. Mainly lasers in the green spectral range (Ar-ion lasers with 514 nm or frequency-doubled Nd-lasers at 532 nm) are used, but also lasers and laser diodes in the near IR-range.
The only dosage of laser photocoagulation employed so far consists in subsequently checking the ophthalmologic appearance of the coagulation site at the fundus. That the retina turns grey or white in the process, shows an irreversible thermal necrosis of the neuronal retina that can reach, depending on the intensity and extent of the coagulation, the entire retina from the retinal pigment epithelium (RPE) via the photoreceptors to the nerve fiber layer and can also include necroses of all uvulas. The large spatial extent of the coagulation effects results from the thermal conduction from the melanine-containing absorbing layers into the neighboring tissue layers.
The pan-retinal laser coagulation of the retina is the most common use of the laser in ophthalmology. Here the inner layers of the peripheral retina are to be destroyed thermally in 3-10 sessions with up to 3000 laser foci of different size so as to prevent the unchecked vessel growth and the blindness connected thereto at a later point. With most patients the laser treatment is extremely painful. Only a retrobulbar injection can avoid pain. This however entails that the mobility of the bulbus that is necessary for carrying out the photocoagulation in the periphery is switched off: in the case of over-coagulations and in particular in the case of repeat treatments, additional thermal damage of ganglion cell layers is to be feared that can lead to extensive defects of the field of vision.
Since the absorbing granula vary considerably in terms of their local and spatial density it is not surprising that the histological results after laser coagulation can vary considerably even in the case of identical exposition parameters. The extent of the damage essentially is a function of the extent of the laser-induced rise in temperature. In practice, it cannot be predicted due to different pigmentation and thus absorption of the retina both inter- and also intraindividual.
Automatic, temperature-controlled online dosimetry for laser treatment with minimal invasive damage is a goal to be desired that cannot be achieved through the application method that is conventional at the moment for ophthalmoscopic examination.
Physics offers different methods for measuring temperatures, however almost all of them are practically unsuitable for measuring the ocular fundus.
Invasive measurement methods as for example thermal probes or dyes that fluoresce as a function of the temperature are too annoying—among others on account of side effects—and/or too imprecise. Due to the absorption of the thermal radiation in the eye, thermal imaging cameras cannot be used.
Methods that are based on auto-fluorescence seem to be suitable, as is for example taught by DE 102 40 109 A1, but a uniform distribution of the chromophores that does not exist in practice is a precondition here.
The analysis of temperature-dependent, thermo-mechanical expansion of an absorber and the pressure wave emitted therewith after the application of a short laser pulse has been described in Sigrist M. W., “Laser Generation of Acoustic Waves in Liquids and Gases”, Journal of Applied Physics 60(7):R83-R121, 1986. On this basis the optoacoustic temperature measurement on the retina was developed as it is illustrated in DE 101 35 944 C2. Additional, repetitive irradiation with short laser pulses produces pressure transients whose amplitude can be recorded with an ultrasound sensor (for example piezo element) that is integrated into the contact lens required for laser treatment anyway. The instantaneous increase in temperature can be determined from the amplitude. In the process, the dependency of the temperature on the choroid perfusion and the light absorption and thus the absolute necessity of an online dosimetry based on the temperature could be illustrated.
The method of DE 101 35 944 C2 was previously used for temperature measurements in Transpupillary Thermotherapy (TTT) and the Selective Retina Therapy (SRT). In the case of SRT, the treatment pulses themselves can be used for determining the temperature. WO 2005/007002 A1 further describes strategies that use the optoacoustic signal for controlling the therapeutic laser. WO 2005/007002 A1 however assumes that microscopic bubbles will form sooner or later due to the laser impact. Such bubbles significantly change the behavior of the pressure transients and thus serve to identify the damage threshold in the vicinity of which the laser should operate. This is then realized by a suitable feedback.
In laser photocoagulation, only maximum temperatures of 40-80° C. are realized for treatment of the ocular fundus. Bubble formation cannot set in below 100° C., so that this cannot be any option for laser control.
The aim of photocoagulation is the thermal denaturation of proteins and tissue. It is in particular the dependency of the extent and depth of damage of coagulations of the retina that has been well researched experimentally and theoretically using different laser parameters (e.g. Birngruber R, Hillenkamp F, Gabel V P., “Experimental studies of laser thermal retinal injury”, Health Phys 44(5):519-531, 1983 or Birngruber R, Hillenkamp F, Gabel V P., “Theoretical investigations of laser thermal retinal injury”, Health Phys 48(6):781-796, 1985). The findings are that the damage to the tissue is both a function of the duration of the laser irradiation and also directly—and particularly critically—of the temperature increase caused during this period. Here the damage integral Ω describes a certain change that is a function of damage criteria and tissue and that is influenced by the temperature curve T(t) over the total duration of the temperature increase ts.
      Ω    ⁡          (              t        s            )        =      A    ·                  ∫        0        ts            ⁢                          ⁢                        ⅆ          t                ·                  T          ⁡                      (            t            )                          ·                  ⅇ                      -                                          Δ                ⁢                                                                  ⁢                E                                            k                ·                                  T                  ⁡                                      (                    t                    )                                                                                          
The activation energy ΔE and the frequency factor A can be determined experimentally, in that the threshold for thermal damage for different temperature increases and exposition times are determined, k designating the Boltzmann constant. The constants differ for many tissues. Ω is influenced exponentially by the temperature and approximately linearly by the time. This means that the effects of an excessive temperature can be far more serious than that of an irradiation period that is too long. To describe a stronger denaturation, Ω>>1 is selected (e.g. Ω=100), if the value clearly stays below Ω=1, Ω<<1, no thermal changes are to be expected.