A light measurement system for localizing light within a light scattering medium.

Abstract

An optical measurement system in a subject, comprising an acoustic transducer, a laser that emits light within an infrared (IR), visible or ultraviolet wavelength range, an optical filter, a photodetector and an optical reflector.

Description

[Technical Field]

[0001] The present invention relates to an apparatus and method for selectively obtaining light from deep within a light scattering medium, such as biological tissue. The present invention relates to an apparatus and method for achieving more localized measurements than conventional methods by using light from deep within a light scattering medium. In particular, the present invention relates to the use of acoustically frequency-shifted laser light and optical components to achieve localized optical measurements in the depths of a light scattering medium. [Background technology]

[0002] The ability to irradiate light into a cloudy, opaque light-scattering medium—that is, into a light-scattering medium that scatters light—is important in many settings, including biomedical applications. Optical imaging provides high-sensitivity molecular contrast easily and non-invasively. Conventional optical microscopy methods rely on light passing through shallow layers of about tens of micrometers. Confocal microscopy can extend the penetration depth to about 0.1 mm, while optical coherence tomography (OCT) can reach a depth of about 0.5 mm. This is nearly the limit of high-resolution optical imaging modalities, which rely on the suppression of diffuse scattering light. At greater depths, diffuse light scattering has a significant impact on the interaction between the light field and the light-scattering medium. The attenuation of light rays at a particular depth is largely determined by the scattering of light by the light scattering medium. This fundamental aspect means that, even at the most favorable wavelength, the maximum penetration depth to which light can be detected is limited to about 10 cm in typical biological tissues.

[0003] Other optical analysis and imaging techniques are being used in the diffuse scattering region, ranging from approximately 1 millimeter to several centimeters. One example is pulse oximetry, used to determine blood oxygen saturation. One example of optical imaging is diffuse optical tomography, which can be performed, for example, using endogenous tissue contrast, administered fluorescent contrast agents, or administered fluorescent agents. Generally, due to diffuse scattering, spatial positioning using these methods is limited to a range of a few millimeters to one centimeter. Furthermore, because the detected light has a very low intensity, detection methods are often inefficient, cumbersome, and require advanced and expensive equipment.

[0004] To improve the spatial positioning when using light to probe opaque, turbid light-scattering media, photoacoustic tomography techniques have been devised. In this photoacoustic tomography technique, sound waves emitted locally due to slight heating of biological tissue caused by the absorption of laser pulses are detected by an ultrasonic transducer. Photoacoustic tomography utilizes the fact that ultrasound is scattered several orders of magnitude less in biological tissue than light waves. In this way, the spatial location of the sound wave source can be determined to a degree of approximately 1 millimeter or less.

[0005] Another example of using acoustics to improve the localization of photo-interactions in biological tissues is the use of acoustic photon tagging, as described in Non-Patent Document 1. By insonizing biological tissue and directing laser light into this tissue, only the laser light passing through the region occupied by the sound field will experience frequency shifts due to the acoustic frequency. By optically detecting only frequency-shifted light, it is possible to understand the interaction between the sound field and light, and to make it occupy a small, localized volume compared to the light.

[0006] Another example of the use of sound is described in Non-Patent Document 2. Further examples of the overlap between ultrasound and optical spectroscopy are described in Patent Document 1. Using the output from an ultrasonic transducer, ultrasound is focused into a selected object at various modulation frequencies, and the object is physically modulated (vibrated) by the radiation pressure of the ultrasound. However, while photoacoustic methods can offer better spatial resolution than purely optical methods, they have a limitation in that light does not easily penetrate deep into the light-scattering medium due to strong attenuation caused by scattering.

[0007] As mentioned above, known methods for performing optical measurements within light-scattering media have limitations in penetration depth and / or low spatial resolution. Therefore, newly improved apparatus and methods for deep light penetration into light scattering media are advantageous. [Prior art documents] [Patent Documents]

[0008] [Patent Document 1] U.S. Patent Application No. 20060253007 [Non-patent literature]

[0009] [Non-Patent Document 1] J. Gunther & S. Anderson-Engels (October 2017), Review of current methods of acousto-optical tomography for biomedical applications, Frontiers of Optoelectronics, 10(3), 211-238. [Non-Patent Document 2] M. Kempe et al, Acousto-optic tomography with multiply scattered light, Vol. 14, No. 5 / May 1997 / J. Opt. Soc. Am. A. [Non-Patent Document 3] H. Zang et al (March 2012), Slow light for deep tissue imaging with ultrasound modulation, Appl. Phys. Lett. 100(13), 131102. [Summary of the Invention] [Problems to be Solved by the Invention]

[0010] Therefore, embodiments of the present invention, either alone or in any combination, provide the optical measurement system described in the claims to selectively achieve high light intensity deep in the light-scattering medium, for example, to mitigate, alleviate or eliminate one or more deficiencies, disadvantages or problems in the art as described above. [Means for Solving the Problems]

[0011] The optical measurement system 100 of the present invention includes an optical unit and an acoustic unit. The optical measurement system 100 in FIG. 1 includes an optical unit having a laser 12 and an optical filter 14. Also, the acoustic unit of this optical measurement system 100 includes an ultrasonic device 13. Then, this ultrasonic device 13 emits a sound field into the light scattering medium 11. The position of the sound field is 15. The laser 12 may be directed into the light scattering medium 11 through one or more input apertures 19, and the light 18 is diffusely scattered in the light scattering medium 11. And a part of the scattered light 18c passes through the position 15 which is the sound field of the ultrasonic wave, and a part of the light can be frequency-shifted through the interaction with the sound field of the ultrasonic wave. Since the optical frequency of the laser light is much larger in order than the frequency of the ultrasonic wave, the frequency shift corresponds to a small wavelength difference of the laser light. The frequency-shifted light may diffuse from the position 15 and propagate through the light scattering medium 11, and a part of it may exit the light scattering medium 11 through one or more output apertures 20 and reach the optical filter 14. And by this optical filter 14 removing most of the original laser frequency, substantially frequency-shifted light can be obtained and detected. The frequency-shifted light carries information only from the position 15, and with such a setting, spatial information within the light scattering medium 11 can be provided. By scanning this position 15, a map or image of the optical contrast within the light scattering medium 11, for example, information regarding oxygen saturation in a biological tissue, etc. can be obtained. An image obtained by the optical measurement system 100 of this embodiment is shown in FIG. 2.

[0012] Around the laser injection point, a light reflecting member 16 is provided, and this light reflecting member 16 may reflect most of any light leaking from the light scattering medium 11. Similarly, the light reflecting member 17 can be provided around the point where the frequency-shifted light exits the light scattering medium toward the optical filter.

[0013] The light-reflecting member 16 has the function of preventing some of the laser light from leaking out of the light-scattering medium 11 before the laser light interacts with the ultrasonic sound field 15. In this way, the light-reflecting member 16 can play a role in increasing the total amount of light in the biological tissue including point 15, and can increase the frequency-shifted amount of light. As a result, the signal can be increased.

[0014] Similarly, the light-reflecting member 17 can have the function of preventing some of the frequency-shifted light from flowing out of the light-scattering medium 11 through a channel other than the output aperture 20, thereby increasing the amount of light passing through the optical filter 14. Furthermore, as a result, the signal can be reduced.

[0015] Here, the optical filter 14 can be composed of a slow-light filter, as shown in Figures 3 and 4. Furthermore, the optical filter 14 may be composed of a host crystal 21 doped with ions 22 that have strong absorption at the original frequency of the laser 23. Furthermore, when this ion 22 replaces other atoms in the host crystal 21, the ion 22 may slightly distort the crystal lattice 21. The individual absorption of these ions 22 is strong and narrow in frequency, but a different crystal field is observed in the entire collection of doped ions, and therefore the frequencies at which they absorb may change slightly in relation to each other. Figure 3 shows three different ion classes 22, each exhibiting different energy separation between the ground state and the excited state, as seen in Figure 3b). The numerous different ion classes allow us to obtain the total absorption profile of the doped ion aggregate of the host crystal, and its frequency range can be several orders of magnitude wider than any frequency shift caused by the sound field.

[0016] This absorption profile allows for the construction of different persistent spectral structures using optical pumping techniques. This optical pumping can be performed with or without applying an electric and / or magnetic field to split the energy levels of ions. Figure 4 shows an example of the resulting filter and crystal absorption profile. In Figure 4, the absorption of light 23 at the original laser frequency may be unaffected or even high, but the absorption of frequency-shifted light 24 can be excluded, and the frequency-shifted light 24 can travel to the detector. A side effect of this modified absorption profile is that the group velocity of frequency-shifted light may decrease by several orders of magnitude. Therefore, time gating allows the signal 24 to be further distinguished from the original laser light 23. It should be emphasized that the term "light" as used herein is not limited to light visible to the human eye, but also includes wavelengths in the ultraviolet and infrared regions.

[0017] As used herein, the term “equips” shall be interpreted as indicating the presence of a described feature, integer, step, or component, and shall not preclude the presence or addition of one or more other features, integers, steps, components, or groups thereof. [Brief explanation of the drawing]

[0018] These and other embodiments and features of the present invention will become apparent and clear from the following description of the embodiments of the present invention with reference to the drawings.

[0019] [Figure 1] This shows an example of image generation settings using acoustic photon tagging. [Figure 2] The image shown is of a photoabsorbent inclusion captured using the embodiment of the optical measurement system in Figure 1. [Figure 3]A schematic crystal containing three different classes of doped ions is shown. Slow-light filters can be fabricated using such crystals. [Figure 4] A schematic slow-light filter is shown. By manipulating the absorption of the crystal's ionic class, frequency-shifted light (solid line) is allowed to pass through, while the original laser frequency (dashed line) is absorbed. [Figure 5] A schematic setup for a transmission mode experiment using a homogeneous medium is shown. [Figure 6] The frequency-shifted signal intensity is shown for cases with and without the use of light-reflecting materials at various depths. [Figure 7] This shows the frequency-shifted signal intensity for a single signal, with and without the use of a light-reflecting material at a given depth. [Figure 8] This shows the effect of the size of the light-reflecting material on the acoustic photon tagging signal intensity. [Figure 9] This shows the experimental setup for reflection mode imaging of light-absorbing inclusions. [Figure 10] The images show light-absorbing inclusions taken with and without a reflective material covering the input side (a). [Figure 11] This shows the effect of light-reflective materials on the ability to identify oxygenation levels, with and without their use. [Figure 12] This demonstrates the existence of an optimal design configuration for imaging tissue oxygenation using acoustic photon tagging. [Figure 13] This diagram schematically illustrates how a hollow core optical guide improves the amount of light directed to the crystal within a cryostatic device. [Modes for carrying out the invention]

[0020] The following inventions focus on embodiments of the present invention that are applicable to improving the localization of optical interactions in light scattering media. The present invention may be applicable to the measurement of subjects, for example, humans or animals. The present invention may be applicable to measuring deep oxygen saturation in biological tissues, such as the brain, heart, female breasts, and muscle tissue. On the other hand, this description is not limited to this application and can be applied to a number of other systems where localizing optical interactions in a light scattering medium is useful.

[0021] The present invention generally comprises an optical unit and an acoustic unit. In Figure 1, the optical measurement system 100 includes an optical section of the system having a light source 12, for example, a laser, and an optical filter 14. The acoustic section of the optical measurement system 100 includes an ultrasonic device 13. The ultrasonic device 13 emits a sound field into the light scattering medium 11 so as to occupy point 15. The light ray 18a emitted by the laser 12 is directed into the light scattering medium 11, where the light ray 18a is diffusely scattered to become diffuse light 18b. A portion of the diffused light 18b can pass through the area occupied by the ultrasonic sound field 15 and can be frequency-shifted through interaction with the ultrasonic sound field 15. The frequency of ultrasound is 10 6 It is sufficient for it to be around Hz, and the optical frequency is 10 14 It can be around Hz. Since the optical frequency of laser light is on a much larger order of magnitude than the frequency of ultrasound, the frequency shift corresponds to a small wavelength difference in the laser light, and this wavelength difference can be on the order of a few femtometers. The frequency-shifted light may diffuse from the position occupied by the ultrasonic sound field 15 and propagate through the light scattering medium, and a portion of the frequency-shifted light 18c can exit the light scattering medium and reach the optical filter 14. The optical filter 14 removes most of the original laser frequency, thereby obtaining substantially frequency-shifted light, which can then be detected. The frequency-shifted light, 6 If the separation is only at around Hz, the optical filter 14 is approximately 10 14It needs to have special capabilities to suppress the carrier frequency in Hz. Furthermore, the optical filter 14 needs to be largely independent of the angle of the diffusing incident light. This performance can be implemented as a so-called slow light filter, as described in Non-Patent Document 3. The frequency-shifted light carries information only from position 15, and therefore, this setup provides spatial information within the light scattering medium 11. By scanning the position 15, which represents the volume occupied by the ultrasound, a map or image of the optical contrast within the light scattering medium 11 can be obtained, such as information regarding oxygen saturation in the tissue. Lateral scanning can be performed either by mechanically moving the ultrasonic source or by electronically moving the ultrasonic sound field by controlling the individual elements of the ultrasonic source. Vertical scanning is performed by timing the optical pulses to match the positions of the ultrasonic pulses at various depths. The experimentally obtained image generated in this way can be seen in Figure 2, where position 15 is scanned over a high absorption region with a wavelength of less than 800 nm, i.e., a low oxygen saturation region. To extract information regarding oxygen saturation, it is preferable to perform measurements at at least two different wavelengths. For example, it is preferable to have one wavelength of approximately 800 nm and one wavelength of 600 nm to 770 nm. The spectral absorption profiles of oxygen-saturated hemoglobin and unoxygen-saturated hemoglobin are similar in terms of absorbance at approximately 800 nm, but differ in their composition at wavelengths below or above 800 nm. Alternatively, one wavelength can be selected as approximately 800 nm, and the other wavelength as being in the range of 820 nm to 1200 nm. By comparing signals at two wavelengths, the ratio of oxygen-saturated hemoglobin to non-oxygen-saturated hemoglobin can be estimated. Furthermore, this method is not limited to just two wavelengths. For example, it may be preferable to use three or more wavelengths to improve the accuracy when determining oxygen saturation.

[0022] To sufficiently improve the signal level, it is advantageous to add light-reflecting elements around the laser injection point and / or the optical filter exit point. Referring to Figure 1, a first light reflecting member 16 is provided around the laser injection point to reflect most of the light leaking from the light scattering medium 11. Similarly, the second light reflecting member 17 can be provided around the point where the frequency-shifted light exits the light scattering medium toward the optical filter 14.

[0023] The first light reflecting member 16 has the function of preventing a portion of the laser light from leaking out of the light scattering medium 11 before the laser light interacts with the ultrasonic sound field at point 15. In this way, the first light reflecting member 16 can increase the total amount of light at point 15, thereby playing a role in increasing the amount of frequency-shifted light. As a result, the signal can be increased.

[0024] Similarly, the second light reflecting member 17 has the function of preventing some of the frequency-shifted light from flowing out of the light scattering medium 11 before it is emitted and collected by the optical filter 14. In this way, the second light reflecting member 17 can play a role in increasing the amount of frequency-shifted light collected by the optical filter 14. Furthermore, as a result, the signal can be reduced.

[0025] As mentioned above, adding a light-reflecting material to the surface of the light-scattering medium is advantageous in the case of photoacoustic methods because, in this method, it is not necessary to spatially localize the original laser light. This is, for example, in contrast to diffuse optical tomography, in which adding a light reflecting member to the surface of the light scattering medium increases the light fluence within the light scattering medium while reducing the spatial position specificity of the light, which is an important performance of diffuse optical tomography.

[0026] Also, the light reflecting member may be formed from any suitable material having a high reflectivity at the relevant wavelength. This includes glass or metal mirrors, metal surfaces, synthetic resins having a high reflectivity, surfaces coated with a reflective paint, and the like. In this context, a surface having a reflectivity of more than about 80% is considered to have a high reflectivity, but even a low reflectivity value, adding any reflective surface is advantageous.

[0027] Experiments were conducted to demonstrate the advantages of using a light reflecting member on the surface of the light scattering medium. The experimental setup is shown in FIG. 5. Central frequency f US An ultrasonic pulse having was generated from the ultrasonic transducer 13, and this ultrasonic pulse propagated within a phantom tissue slab 11 having known optical characteristics. When the ultrasonic sound field reached point 15, a short optical pulse from a narrow frequency laser 12 having a frequency f C was emitted into the slab through the point-sized input aperture 19 towards the slab. This initial light can be referred to as the carrier light. The carrier light diffused into the ultrasonic pulse, and a part of it was frequency-shifted by the acousto-optic effect to a frequency f S = f C + f US and became. This ultrasonic-shifted light can also be referred to as tagged light. Next, both light fields were collected at one output aperture 20 with a diameter of 1 cm on the side opposite to the input side. The tagged light was separated from the carrier light using a slow light filter and detected with a photomultiplier tube. Slow light filters and photomultiplier tubes are collectively shown as 14. A light-reflecting film with reflectivity R = 0.98 was used as the light-reflecting member 16 and placed on the input side, which was provided with a 2.6 mm diameter hole for transmitting the carrier light, and the signal light was measured again. For slabs of different thicknesses, the signal light was measured with and without the film.

[0028] The number of detected photons is shown in Figure 6 as a function of the phantom thickness. Figure 7 shows the signal over 1000 shots averaged for a phantom with a thickness of 6.7 cm. Figures 6 and 7 both show that the signal intensity is improved by 2 to 3 times, and that the effect becomes greater as the imaging depth increases.

[0029] To further demonstrate the advantages of using a light-reflecting material on the surface of the light-scattering medium 11, we conducted simulations. It is well known in the art to use the diffusion equation to calculate the optical fluence F in the light scattering medium 11. The diffusion equation can be written as follows:

[0030]

number

[0031] In the above formula, D = 1 / (3(a + c)) is the diffusion constant, a is the absorption coefficient, c is the decay scattering coefficient, and S represents the light source. The light that a photodetector can detect is indicated by the photon flux J that is imminent on the surface of the photodetector. The flux according to Fick's Law is as follows:

[0032]

number

[0033] The boundary conditions for the light diffusion equation are as follows:

[0034]

number

[0035] In the above formula, A is a constant given by the average reflectance of light at the interface of the normal vector n between the biological tissue and the surrounding light scattering medium (e.g., air or a reflective material). The diffusion equation can be solved, for example, by the finite element method.

[0036] To evaluate the effect of light reflection means on the surface of a light scattering medium, it is useful to define the contrast-to-noise ratio (CNR) when measuring the oxygen saturation of biological tissues. In short, CNR is a measure of how well changes in tagged light can be detected when position 15 is moved between two nearby points that have locally different optical properties. These optical properties depend on the oxygenation of the blood in the biological tissue. The CNR per detection position for a number of laser pulse shots N is defined as follows:

[0037]

number

[0038] In the above formula, ST1 and ST2 are shifted light detected from two different regions during scanning of position 15. SC is the detected signal light, and TF is the filter transmittance of the carrier light. For example, with 80dB of filter suppression, TF = 10 -8 That is the case.

[0039] The simulation was performed using COMSOL, a commercially available finite element method solver. Slab shapes with varying thicknesses were simulated with and without a reflective boundary on one side. This reflection boundary has a 2.6 mm diameter hole within it (just like in the experimental setup). The detected transport light was simulated using the flux on the opposite side into the optical guide, generated using a point light source located at a depth of 1 / c below the film hole. To verify the simulation, the carrier optical signal was compared with experimental results, and the data was found to be in agreement in both cases.

[0040] It is beneficial to evaluate the effect of the size of the area covered by the light-reflecting film 16 around the light injection point. In practice, it is desirable to configure the optical reflection boundary to cover the smallest possible area. Figure 8 shows the effect of changes in the size of the light-reflecting film 16 on the signal of the carrier light. These results indicate that the signal intensity increases as the size of the light-reflective film 16 increases. However, if the size of the light-reflecting film 16 is increased beyond 2cm to 3cm, the advantages diminish. This means that the area covered by the light-reflective film 16 does not need to be very large.

[0041] Furthermore, it is beneficial to evaluate the effect of the size of the holes in the light-reflecting film 16, which are necessary for injecting the transport light. Similar to the evaluation described above, signal strength increases as the size of the hole decreases. However, the advantages diminish if the size of the hole is reduced to less than 3mm to 4mm. Therefore, it is not necessary to make the hole extremely small.

[0042] In Figure 5, the light source 12 and the photodetector are arranged in transmission mode, that is, on opposite sides of the light scattering medium 11. However, the present invention is not limited to transmission. The relative positions of the light source 12, the photodetector, and the ultrasonic transducer 13 may be arbitrary. In some cases, it is practical to place all components on the same side and perform measurements in reflection mode. This was done in the second experiment, which is shown in Figure 9. In this experiment, absorbance inclusions 26 were embedded within the phantom tissue 11. The position 15 of the ultrasonic pulse was scanned in the lateral and vertical directions with a resolution of 0.75 mm in a plane perpendicular to the phantom surface. At the location of each ultrasonic pulse, an optical pulse irradiated the phantom at the input aperture 19 to generate frequency-shifted signal pulses, which were then resolved downstream of the output aperture 20 using a slow-light filter.

[0043] Images of the inclusions were obtained for each case in which the light-reflective film, which functions as both the light-reflective member 16 and the light-reflective member 17, is used or not used. These images can be seen in Figures 10a) and 10b), respectively. The light-reflective film covered the portion of the input surface not occupied by the light input, ultrasonic source, and 1 cm diameter light guide.

[0044] In the additional simulations, the simulations were performed in reflection mode. To simulate the signal light, a two-step calculation was performed. First, the carrier light was simulated in the same manner as shown in Figure 8. Next, the tagged light is powered F US *The simulation was performed with the ultrasonic pulse emitting K as the central point source, but here, F US is the carrier fluence of the ultrasonic pulse, and K is an experimentally determined empirical tagging coefficient. Subsequently, the scattering and absorption properties of the light scattering medium were selected for 85% oxygenated muscle tissue at a wavelength of 690 nm. The ultrasonic pulse was modeled as a sphere with a radius of 2 mm. Then, ST1 was assumed to be the tagged optical signal when the local oxygenation around the ultrasonic pulse was set to 40%, and ST2 was assumed to be the tagged optical signal when the local oxygenation was 85%, and the CNR was calculated.

[0045] In Figure 11, N = 200, and CNR is shown as a function of the ultrasound pulse depth for 40% oxygenation in the ultrasound pulse and for a comparison with an 85% baseline. Compared to the case without a light-reflective material, the CNR clearly increases when a light-reflective material is applied. The increase in CNR in the simulation results is in good agreement with the increase in CNR observed in the experimental results. The CNR for the image obtained using the light-reflecting material in Figure 10 is 1.54, while the CNR for the image obtained without using the light-reflecting material is 1.02.

[0046] It is beneficial to evaluate the distance between the laser injection site and the photodetection site and determine whether an optimal distance exists in terms of CNR. Since optical filtering of the carrier light is not perfect, it is sometimes optimal to have a certain distance between the light source and the photodetector. Figure 12 shows the CNR as a function of the separation distance between the light source and the photodetector for an ultrasonic pulse at a depth of 5 cm and 80 dB of filter suppression. The optimal spacing in this case is 20 mm. Generally, the optimal spacing depends on the scattering and absorption of the light scattering medium, the depth of the ultrasonic pulse, the tagging coefficient K, and the filter suppression. Therefore, in reflection mode, the optical measurement geometry can be customized based on these factors.

[0047] The advantages of using light-reflecting material near or adjacent to the light source injection site are not limited to acoustic-photon tagging, but are also advantageous when performing photoacoustic imaging or tomography. In this case, the photoacoustic signal increases due to the increase in the amount of light in the light scattering medium 11.

[0048] As mentioned above, in order to ensure a sufficient signal level in the device, it is important to minimize optical loss between the light collection area and the photodetector. In some cases, light is guided to the slow light filter using an optical guide. However, slow light filters need to be confined in a cryogenic device and stored at low temperatures in order to function as intended.

[0049] This invention is advantageous when used in cases where measuring deep tissue oxygenation is medically beneficial. For example, the present invention can be used in the management of ischemic stroke. In emergency medical situations, it is crucial to quickly diagnose the presence of ischemic areas in the patient's brain. Most patients with ischemic stroke undergo thrombectomy to remove blood clots. In severe cases, patients may not wake up during thrombectomy, making it difficult for surgeons to immediately determine whether the treatment was successful. In such cases, monitoring the oxygenation status of the brain is advantageous as a means of providing feedback for treatment procedures.

[0050] Another medical application in which the present invention is advantageous is monitoring the oxygenation status of the myocardium for emergency treatment in cases where myocardial infarction is suspected.

[0051] Another medical application in which the present invention is advantageous is in sports medicine, where it is used to evaluate oxygen uptake and metabolism in muscles.

[0052] Another medical application in which the present invention is advantageous is in the field of oncology, where oxygen saturation levels may be altered in tumors compared to healthy tissue. For example, tumors may contain necrotic areas with low oxygen saturation.

[0053] Slow-light filters rely on the use of optical transitions of dopant ions within crystalline materials. In order to fabricate a frequency filter narrow enough to separate frequency-shifted light from light of its original frequency, the linewidth of this transition needs to be narrow so that the ends of the absorption Lorentz distribution curve do not overlap between the original laser frequency and the acoustically shifted frequency. The width of these transition lines is limited by the reciprocal of the transition's coherence time. Furthermore, if this coherence time has a limited lifetime, the shorter the lifetime of the high-level state, the wider the line becomes. In the case of ions in a crystal lattice, the reason for the shortened lifetime is the interaction between the ion and phonons, and phonons are vibrational quanta that propagate within the crystal structure. If the higher state of an ion in an optical transition is close to another state, the ion can transition to this state by absorbing or emitting one or more phonons, and the probability of an event involving more phonons becomes increasingly less likely. Furthermore, since the maximum phonon energy is a characteristic of the host crystal, using a "softer" crystal with a lower maximum phonon energy can extend the lifetime of the higher state. In this case, phonon transitions in these crystals require interactions with a larger number of phonons, resulting in a decrease in transition rates and a longer lifetime for higher states. By using soft crystals, more optical transitions of specific ions can be preserved for use in filtration. This allows for the use of a wider range of wavelengths to evaluate the oxygenation of biological tissues using the same crystal.

[0054] Certain dopant ions, such as lanthanum, can be used only to fabricate filters at wavelengths specific to that ion. Therefore, by using multiple different dopant ions, such as lanthanum and europium, it becomes possible to customize where the filter can be fabricated. Multiple ions can be made accessible to the same signal light by doping them into the same host crystal, i.e., by "simultaneous doping," or by combining different crystals. This combination can be achieved by either stacking or layering crystals, i.e., "sandwiching" them, or by melting different crystals together, i.e., "splicing" them.

[0055] The structure fabricated by the absorption profile may be one or more notch filters that select each individual frequency-shifted sideband of the original light. Alternatively, a bandstop filter may be used where the only absorption not excluded is at the original frequency. This is beneficial because the frequency-shifted light in the primary sideband (±1 ultrasonic frequency) accounts for only half of the light that is removed from the original frequency for a given ultrasonic pulse. However, this bandstop filter does not produce as strong a slow-light effect on the light that passes through it as a notch filter.

[0056] Depending on the application and the ultrasonic pulse, both filters can be beneficial.

[0057] For any spectral structure to have the maximum width, the frequencies of one or more other ground states that the ion photo-pumps must be sufficiently far apart. This may be an inherent property of the ion and the crystal itself, where the ground state spontaneously splits into multiple hyperfine states. In some cases, this partitioning can be achieved by applying electric and / or magnetic fields to partition a generally degenerate state. By applying an electric field and / or a magnetic field, the ground state lifetime can be extended, and consequently, the filter lifetime can be extended. This extended filter lifetime reduces the time during the imaging sequence that needs to be allocated to filter preparation. In this way, extending the filter lifetime directly results in faster imaging speeds. Creating a slow light filter has the added effect of simultaneously creating one or more so-called "anti-holes." Such antiholes are locations on the crystalline absorption profile where absorption is actually increased. At least one anti-hole is located at a frequency distance equal to the ground state, which is divided from the center frequency of the slow light filter. This occurs due to optical pumping, and at that time, more ions occupy this state. In this way, by applying the correct electric and / or magnetic fields, the efficiency of the slow light filter can be improved by providing anti-hole protection at the original laser frequency, in addition to extending the Hall lifetime.

[0058] Inside a cryogenic apparatus, a magnetic field can be generated using permanent magnets. This has the advantage of not requiring power supply or cooling as in the case of an external electromagnet, and not generating further thermal load through lead wires as in the case of using superconducting magnet coils or capacitors. Furthermore, placing magnets inside the cryogenic device allows for the generation of a stronger magnetic field with smaller magnets compared to placing them outside. In this way, using permanent magnets makes it possible to reduce both the size and the cooling power consumption of the cryostat device being used.

[0059] Furthermore, the shape of the ultrasonic pulse also has a significant impact on the spatial resolution of the image and the intensity of the frequency-shifted light. Furthermore, this ultrasonic pulse can be chirpened, which, along with a slow light filter, allows for the probe of multiple spatial points with the same ultrasonic pulse. This effectively reduces acquisition time, enabling faster image reconstruction.

[0060] A chirp ultrasonic pulse is a pulse that has a gradient of center frequency along its temporal and spatial profile. When short light pulses are used, the frequency-shifted light is also chirpped in this way, and each frequency component is associated with a different spatial region. Because filtered light of different frequencies propagates at different speeds, these frequency components can be separated in time using a slow light filter. In this case, the different readout times correspond to different probe positions in the biological tissue.

[0061] In some cases, multiple ultrasonic pulses are generated simultaneously at different frequencies, and each pulse frequency may produce the same effect. Furthermore, if each of these pulse frequencies is suited to a slow-light filter with a different time delay, multiple points can be imaged with the same light pulse.

[0062] To optimize the detection signal, it is important to collect the signal light from the light scattering medium 11 and guide it to the filter as efficiently as possible. To achieve this, it is preferable to use an optical guide device, such as a single optical guide or a bundle of optical fibers. Furthermore, when this optical guide device is coupled to a cryogenic chamber that holds the filter, it is important to minimize light loss and maintain the insulation of the cryogenic chamber. An embodiment that achieves this is shown in Figure 13. The hollow optical guide 27, which has a reflective wall, guides light from an external optical guide device to the inside and outside of the crystal 21 located inside the cryogenic apparatus. In this way, the light rays 28 that do not collide with the filter and then diverge so greatly that they cannot be imaged on the photodetector are contained by the hollow optical guide 27. This device allows a vacuum to be maintained within the hollow optical guide 27, thereby maintaining thermal integrity. At the same time, the hollow light guide 27 brings the light-receiving surface on which the filter acts closer to the outside of the cryogenic device. Since there is no need to store the image within the optical system, the hollow light guide 27 makes it possible to minimize light loss both towards and from the filter.

[0063] When presenting oxygen saturation data to the user, it is preferable to simultaneously display the conventional ultrasound scan (e.g., B-mode scanning) and the image showing oxygen saturation, using either an absolute or relative scale. Oxygen saturation can be represented as a 2D luminance monochrome image or a 2D false color image. A preferred representation is to overlay a false-color oxygen saturation image onto a conventional ultrasound image. In a preferred method, oxygen saturation images can be switched on or off using software functions, or gradually blended onto ultrasound images.

[0064] The present invention has been described above based on embodiments. However, other embodiments are similarly possible within the scope of the present invention. Different method steps can be provided within the scope of the present invention. The various features and steps of the present invention can be combined in combinations other than those described above. The scope of the present invention is limited only by the appended claims.

[0065] As used herein and in the claims, the indefinite articles "a" and "an" mean "at least one" unless otherwise specified. As used herein and in the claims, the phrase "and / or" means elements that are linked, such as "either one or both," that is, elements that are linked in some cases and unlinked in others.

Claims

1. In a light measurement system for a test subject, It comprises an ultrasonic transducer, a laser that emits light within the wavelength range of infrared (IR), visible light, or ultraviolet light, an optical filter, a photodetector, and a light reflecting member. The ultrasonic transducer is configured to direct an ultrasonic sound field having an ultrasonic frequency to at least one location within the tissue area of ​​the subject. The laser is configured to direct light having an optical frequency into the tissue area of ​​the subject, The optical filter is positioned upstream of the photodetector so as to suppress the optical frequency while transmitting light having an optical frequency shifted by the frequency of the ultrasonic wave. The photodetector detects the light whose frequency has been shifted by the frequency of the ultrasonic wave. A light measurement system configured such that the light-reflecting member is positioned on the surface of the tissue portion of the subject.

2. The light measurement system according to claim 1, wherein the light-reflecting member is arranged on the surface of the tissue portion to increase the amount of light detected by the photodetector.

3. The optical measurement system according to claim 1, wherein the light-reflecting member is configured to be arranged around the injection point of the laser.

4. The light measurement system according to claim 1, wherein the light reflecting member is arranged around the point from which the frequency-shifted light exits the tissue area toward the optical filter.

5. The light-reflecting member has holes for light to enter the tissue area and / or for the frequency-shifted light to exit, The optical measurement system according to claim 1, wherein the hole has a diameter of 3 mm to 4 mm.

6. At least one location within the tissue region is occupied by at least one ultrasonic sound field. The optical measurement system according to claim 1, wherein the laser is directed so that at least a portion of the emitted light passes through the position.

7. The optical measurement system according to claim 1, wherein the frequency-shifted light carries information from at least one location within the tissue site.

8. The optical measurement system according to claim 1, wherein the laser and ultrasonic sources are configured to scan at least one location to obtain a map or image of optical contrast within the tissue site, for example, information regarding oxygen saturation in the tissue site.

9. The light measuring system according to claim 1, wherein the light reflective member is composed of at least one of glass, a metal mirror, a metal surface, a light-reflective synthetic resin, or a surface covered with a reflective paint.

10. The light measurement system according to any one of claims 1 to 9, wherein the optical filter has a slow light filter structure.

11. The light measurement system according to claim 10, wherein the slow light filter structure comprises two different absorption lines and constitutes a filter at two wavelengths.

12. The light measurement system according to claim 10, wherein the slow light filter structure comprises a host crystal.

13. The optical measurement system according to claim 12, wherein the host crystal is composed of a single ion, is simultaneously doped with multiple ions, or is composed of multiple different crystals that are spliced ​​or sandwiched together.

14. The light measurement system according to claim 10, wherein the slow light filter structure includes a notch filter configuration that blocks all light other than the frequency-shifted light.

15. The light measurement system according to claim 10, wherein the slow light filter structure comprises a bandstop filter configuration that blocks only light of the original frequency.

16. The optical measurement system according to claim 10, wherein elements that generate an electric field and / or a magnetic field are configured to affect the lifespan of the slow light filter structure.

17. The optical measurement system according to claim 16, wherein the electric field and / or magnetic field are positioned at the original laser frequency of the laser, with the antihole of the slow light filter structure.

18. The optical measurement system according to claim 16, wherein the element that generates the electric field and / or magnetic field comprises a permanent magnet.

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