Light measurement system for localizing light in a light scattering medium - Patents.com

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 deep within a light scattering medium, such as biological tissue. The present invention relates to an apparatus and method for using light deep within a light scattering medium to achieve more localized measurements than conventional methods. In particular, the present invention relates to the use of acoustic frequency shifted laser light and optics to achieve localized optical measurements deep within a light scattering medium. [Background technology]

[0002] The ability to project light into turbid, opaque, light-scattering media, i.e., media that scatters light, is important in many settings, including biomedical applications. Optical imaging provides sensitive molecular contrast simply and non-invasively. Conventional optical microscopy relies on the transmission of light through shallow layers, on the order of tens of micrometers. Confocal microscopy extends the penetration depth to around 0.1 mm, while optical coherence tomography can reach depths of around 0.5 mm. This is approximately the limit of high-resolution optical imaging modalities that rely on the suppression of diffusely scattered light. At greater depths, diffuse light scattering has a significant effect on the interaction of the light field with the light scattering medium. The attenuation of a light ray at a particular depth is determined largely by the scattering of the light-scattering medium. This fundamental aspect means that the maximum penetration depth at which light can be detected, even at the most favorable wavelengths, is limited to around 10 cm in typical biological tissue.

[0003] Other optical analysis and imaging techniques have been developed in the diffuse scattering region, from about one millimeter up to several centimeters. One example is pulse oximetry for determining oxygen saturation in the blood. One example of optical imaging is diffuse optical tomography, which can be performed, for example, using intrinsic tissue contrast, using an administered fluorescent contrast agent, or using an administered fluorescent agent. Typically, diffuse scattering limits spatial localization using these techniques to within a few millimeters to a centimeter. Additionally, the light intensity of the detected light is very low, and detection techniques are often inefficient, cumbersome, and require sophisticated and expensive equipment.

[0004] Photoacoustic tomography techniques have been devised to improve spatial localization when using light to probe turbid, opaque, light-scattering media. In photoacoustic tomography, acoustic waves are locally emitted due to slight heating of living tissue caused by absorption of a laser pulse and are detected by an ultrasound transducer. Photoacoustic tomography exploits the fact that ultrasound waves scatter several orders of magnitude less than light waves in biological tissue. In this way, the spatial localization of the source of the acoustic wave can be on the order of one millimeter or less.

[0005] Another example of using acoustics to improve the localization of optical interactions in biological tissue is the use of acoustic photon tagging, as described, for example, in "Acoustic Photon Tagging" by Friedrich Schmidt, "Analyzing the Localization of Optical Interactions in Living Tissues," Proc. By insonifying tissue and directing laser light within the tissue, only the laser light passing through the area occupied by the acoustic field will be frequency shifted by the acoustic frequency. By optically detecting only the frequency-shifted light, it is possible to know that the sound field has interacted with the light, which can occupy a small, localized volume compared to the light.

[0006] Another example using acoustics is described in Non-Patent Document 2. A further example of the overlap of ultrasound with optical spectroscopy is described in US Pat. No. 5,399,633. The output from an ultrasonic transducer is used to focus ultrasonic waves at various modulation frequencies into a selected object, physically modulating (vibrating) the object with the ultrasonic radiation pressure. However, while photoacoustic techniques can provide better spatial resolution than purely optical techniques, they have the limitation that light does not easily penetrate deep into light-scattering media due to strong attenuation due to scattering.

[0007] As discussed above, known methods for making optical measurements in light scattering media suffer from limited penetration depth and / or poor spatial resolution. Therefore, new and improved apparatus and methods for deep light penetration into light scattering media would be 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 (2017, October), 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 (2012, March), Slow light for deep tissue imaging with ultrasound modulation, Appl. Phys. Lett. 100(13), 131102. Summary of the Invention [Problem to be solved by the invention]

[0010] Accordingly, embodiments of the present invention, alone or in any combination, seek to mitigate, alleviate or eliminate one or more deficiencies, disadvantages or problems in the art, e.g. as described above, by providing a light measurement system as claimed in the appended claims for selectively achieving high light intensity deep within a light scattering medium. [Means for solving the problem]

[0011] The optical measurement system 100 of the present invention includes an optical section and an acoustic section. The optical measurement system 100 of FIG. 1 includes an optical section having a laser 12 and an optical filter 14 . The acoustic part of the optical measurement system 100 also comprises an ultrasonic device 13 . The ultrasonic device 13 then emits an acoustic field into the light scattering medium 11 . The sound field position is 15. A laser 12 may be directed through one or more input apertures 19 into the light scattering medium 11 where light 18 is diffusely scattered. A portion of the diffuse light 18c then passes through position 15, which is an ultrasonic sound field, and a portion of the light can be frequency shifted through interaction with the ultrasonic sound field. Since the optical frequency of the laser light is an order of magnitude larger than the frequency of the ultrasound, the frequency shift corresponds to a small wavelength difference of the laser light. The frequency-shifted light may diffuse from position 15 and propagate through the light scattering medium 11, some of which may exit the light scattering medium 11 through one or more output apertures 20 and reach the optical filter 14. The optical filter 14 then filters out most of the original laser frequency, resulting in a substantially frequency-shifted light that can be detected. The frequency shifted light carries information from position 15 only and such a setup can provide spatial information within the light scattering medium 11 . By scanning this location 15, a map or image of the optical contrast within the light scattering medium 11 can be obtained, such as information regarding oxygen saturation in living tissue. An image obtained by the light measurement system 100 of this embodiment is shown in FIG.

[0012] Around the laser injection point is a light reflecting member 16 which may reflect a large portion of any light that escapes from the light scattering medium 11 . Similarly, a light reflecting member 17 may be provided around the point where the frequency shifted light exits the light scattering medium towards the optical filter.

[0013] The light reflecting member 16 has the function of preventing a portion of the laser light from escaping from 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 serve to increase the total amount of light in the biological tissue containing the point 15, and can increase the amount of frequency-shifted light. As a result, the signal can be increased.

[0014] Similarly, the light reflecting member 17 can function to prevent some of the frequency-shifted light from exiting the light scattering medium 11 through anything other than the output opening 20, thus increasing the amount of light passing through the optical filter 14. This also results in an increased signal.

[0015] Here, the optical filter 14 can be composed of a slow-light filter, as shown in FIGS. This optical filter 14 may also consist of a host crystal 21 doped with ions 22 that have strong absorption at the original frequency of the laser 23 . And when the ions 22 replace other atoms in the host crystal 21, the ions 22 can distort the crystal lattice 21 slightly. Although the absorption of these ions 22 individually is strong and narrow in frequency, the entire collection of doped ions will see different crystal fields and therefore may absorb at slightly different frequencies relative to each other. In figure 3, three different ion classes 22 are shown, which have different energy separations between the ground and excited states, as can be seen in figure 3b). The large number of different ion classes gives rise to a total absorption profile of the ensemble of doped ions in the host crystal, the frequency of which can be several orders of magnitude wider than any frequency shift caused by the acoustic field.

[0016] In this absorption profile, it is possible to create distinct and persistent spectral structures using optical pumping techniques. This optical pumping can be performed with or without the application of electric and / or magnetic fields to split the energy levels of the ions. An example of the resulting filter and crystal absorption profile can be seen in FIG. In FIG. 4, the absorption of the original laser frequency light 23 may be unaffected or high, but the absorption of the frequency shifted light 24 may be eliminated and the frequency shifted light 24 may proceed to the detector. A side effect of this modified absorption profile can be that the group velocity of the frequency-shifted light is reduced by several orders of magnitude. Therefore, the signal 24 can be further distinguished from the original laser light 23 by time gating. 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] The term "comprising" as used in this specification is to be interpreted as indicating the presence of stated features, integers, steps or components, and does not exclude the presence or addition of one or more other features, integers, steps, components or groups thereof. [Brief description of the drawings]

[0018] These and other aspects and features by which embodiments of the present invention may be practiced will become apparent and elucidated from the following detailed description of the embodiments of the invention, taken in conjunction with the drawings.

[0019] [Figure 1] 1 shows an example imaging setup using acoustic-photon tagging. [Diagram 2] 2 shows an image of a light-absorbing inclusion taken using the embodiment of the light measurement system of FIG. 1; [Diagram 3]FIG. 1 shows a schematic crystal with three different classes of doped ions that can be used to make slow light filters. [Figure 4] A schematic slow light filter is shown, where the absorption of an ionic class of crystals is manipulated to allow frequency-shifted light (solid line) to pass, while the original laser frequency (dashed line) is absorbed. [Diagram 5] 1 shows a schematic transmission mode experimental setup with homogeneous media. [Figure 6] 4 shows frequency-shifted signal intensity with and without a light reflecting member at various depths. [Figure 7] 4 shows the frequency-shifted signal strength for a single signal at a given depth with and without the use of a light reflecting member. [Figure 8] 13 illustrates the effect of the size of the light reflecting member on the acoustic-photon tagging signal strength. [Figure 9] 1 shows a reflectance mode experimental setup for imaging light absorbing inclusions. [Figure 10] The images of the light-absorbing inclusions taken with (a) a reflective member covering the input side and without (b) are shown. [Figure 11] 13 shows the effect of the light reflecting member on the ability to discern oxygenation levels with and without the light reflecting member. [Figure 12] We show that an optimal design configuration exists for imaging tissue oxygenation using acoustic photon tagging. [Figure 13] 1 illustrates diagrammatically how a hollow core light guide improves the amount of light guided to a crystal in a cryostatic device. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0020] The following invention focuses on embodiments of the invention that are applicable to improving the localization of light interactions in light scattering media. The invention may be applicable to measurements on subjects, for example humans or animals. The invention may be applicable to measuring oxygen saturation deep in biological tissue, such as the brain, heart, female breast, muscle tissue, etc. However, the present description is not limited to this application and can be applied to many other systems in which it is useful to localize light interactions in a light-scattering medium.

[0021] The present invention generally comprises an optical portion and an acoustic portion. In FIG. 1, a light measurement system 100 includes a light source 12 , eg, a laser, and the optical portion of the system includes an optical filter 14 . The acoustic part of the optical measurement system 100 comprises an ultrasound device 13 . An ultrasonic device 13 launches an acoustic field into the light scattering medium 11 to occupy a point 15 . A light beam 18a emitted by the laser 12 is directed into the light scattering medium 11, where the light beam 18a is diffusely scattered to become a diffuse light beam 18b. A portion of the diffuse light 18b may pass through the location occupied by the ultrasonic field 15 and may be frequency shifted through interaction with the ultrasonic field 15. The frequency of the ultrasound is 10 6 Hz, and the optical frequency may be on the order of 10 14 It may be on the order of Hz. Because the optical frequency of the laser light is an order of magnitude larger than the frequency of the ultrasound, the frequency shift corresponds to a small wavelength difference in the laser light, which may be on the order of a few femtometers. The frequency-shifted light may diffuse from the location occupied by the ultrasonic sound field 15 and propagate through the light-scattering medium, and a portion of the frequency-shifted light 18c may exit the light-scattering medium and reach the optical filter 14. An optical filter 14 removes most of the original laser frequency, resulting in substantially frequency-shifted light which can be detected. The frequency-shifted light is 6 If the frequencies are only separated by about 10 Hz, the optical filter 14 will 14It is necessary to have special capabilities to suppress the carrier frequency of 100 Hz. Additionally, the optical filter 14 must be largely independent of the angle of the incident light that it scatters. This capability can be implemented as a so-called slow light filter, as described in Non-Patent Document 3. The frequency shifted light carries information from position 15 only and therefore this setup provides spatial information within the light scattering medium 11 . By scanning a location 15 representing the volume occupied by ultrasound waves, a map or image of the optical contrast within the light scattering medium 11 can be obtained, such as information regarding oxygen saturation in tissue. Lateral scanning can be achieved either by mechanically moving the ultrasound source or by controlling individual elements of the ultrasound source to electronically move the ultrasound field. Longitudinal scanning is accomplished by timing the light pulse to coincide with the various depth positions of the ultrasound pulse. An experimentally obtained image produced in this manner can be seen in FIG. 2, where location 15 is scanned across a region of high absorption below 800 nm wavelength, ie, low oxygen saturation. In order to extract information regarding oxygen saturation, measurements are preferably taken at least at two different wavelengths. For example, it is preferable to have one wavelength of about 800 nm and one wavelength of 600 nm to 770 nm. The spectral absorption profiles of oxygen-saturated and non-oxygen-saturated hemoglobin are similar in absorbance at about 800 nm, but are composed differently at wavelengths below and above 800 nm. Alternatively, one wavelength can be selected to be approximately 800 nm and the other wavelength in the interval from 820 nm to 1200 nm. By comparing the signals at the two wavelengths, the ratio of oxygen-saturated to non-oxygen-saturated hemoglobin can be estimated. However, this technique is not limited to two wavelengths. For example, it may be preferable to use more than two wavelengths to improve accuracy in determining oxygen saturation.

[0022] To improve the signal level sufficiently, it may be advantageous to add light reflecting members around the laser injection location and / or the optical filter exit location. Referring to FIG. 1, a first light reflecting member 16 that reflects most of the light leaking from the light scattering medium 11 is provided around the laser injection point. Similarly, a second light reflecting member 17 may be provided around the point where the frequency shifted light exits the light scattering medium towards the optical filter 14 .

[0023] The first light reflecting member 16 has the function of preventing a portion of the laser light from escaping from 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 serve to increase the total amount of light at point 15, thereby increasing the amount of frequency-shifted light. As a result, the signal can be increased.

[0024] Similarly, the second light reflecting member 17 serves to prevent some of the frequency-shifted light from escaping the light scattering medium 11 before it leaves and is collected by the optical filter 14. In this manner, the second light reflecting member 17 can serve to increase the amount of frequency-shifted light collected by the optical filter 14 . This also results in an increased signal.

[0025] As mentioned above, the addition of a light reflecting member to the surface of the light scattering medium is advantageous for the photoacoustic technique because the photoacoustic technique does not require spatial localization of the original laser light. This is in contrast to, for example, diffuse optical tomography, where adding light-reflecting elements to the surface of a light-scattering medium can increase the light fluence within the light-scattering medium but can also reduce the spatial localization of light, a key performance feature of diffuse optical tomography.

[0026] Additionally, the light reflective member may be formed from any suitable material that has high reflectivity at the relevant wavelengths. This includes glass or metal mirrors, metal surfaces, highly reflective synthetic resins, surfaces coated with reflective paint, and the like. In this context, a surface having a reflectivity of greater than about 80% is considered to have high reflectivity, although adding any reflective surface, even at lower reflectivity values, can be advantageous.

[0027] An experiment was carried out to prove the advantage of using a light reflecting member on the surface of a light scattering medium. The experimental set-up is shown in Figure 5. center frequency f US An ultrasound pulse having a magnitude of 100 nm was generated from an ultrasound transducer 13 and propagated into a phantom tissue slab 11 with known optical properties. When the ultrasonic field reaches point 15, the frequency f C A short light pulse from a narrow frequency laser 12 having a wavelength of 1000 nm was launched into a dot-sized input aperture 19 towards the slab. This initial light can be called a carrier light. The carrier light is diffused into the ultrasonic pulse, and a part of it is frequency shifted by the acousto-optic effect to a frequency f S = f C + f US It became. This ultrasonically shifted light can also be referred to as tagged light. Both light fields were then collected at a single output aperture 20 with a diameter of 1 cm, opposite the input side. The tagged light was separated from the carrier light using a slow light filter and detected by a photomultiplier tube. The slow light filters and photomultiplier tubes are collectively designated 14. A light reflecting film with a reflectance R = 0.98 was used as the light reflecting member 16, and was placed on the input side provided with a hole with a diameter of 2.6 mm for transmitting the carrier light, and the signal light was measured again. For different slab thicknesses, the signal light was measured with and without the film.

[0028] The amount of detected photons as a function of phantom thickness is shown in FIG. FIG. 7 shows the signal over time for a 1000 shot average for a phantom with a thickness of 6.7 cm. Both Figures 6 and 7 show that the signal strength is improved by a factor of 2-3, with the effect being greater at deeper imaging depths.

[0029] To further demonstrate the advantages of using a light reflecting member on the surface of the light scattering medium 11, a simulation was performed. The use of the diffusion equation to calculate the light fluence F in the light scattering medium 11 is well known in the art. 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 attenuation scattering coefficient, and S represents the light source. The light that a photodetector can detect is represented by the photon flux J impinging on the plane of the photodetector. The flux according to Fick's law is:

[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 normal vector n between the biological tissue and its surrounding light scattering medium (eg, air or a reflective material). The diffusion equation can be solved, for example, by the finite element method.

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

[0037]

number

[0038] In the above formula, ST1 and ST2 are the 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, for a filter suppression of 80 dB, TF = 10 -8 It is.

[0039] The simulations were performed using COMSOL, a commercially available finite element solver. Slab geometries with varying thickness were simulated with and without a reflective boundary on one side. This reflecting boundary has a 2.6 mm diameter hole in it (as in the experimental setup). The detected transport light was simulated by a counter flux into the light guide, generated with a point source at a depth 1 / c below the film hole. To validate the simulation, the optical carrier signal was compared with the experimental results, and the data agreed in the two cases.

[0040] It is instructive 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 light reflecting boundaries to cover as small an area as possible. FIG. 8 shows the effect of changing the size of the light reflecting film 16 on the optical signal carrier. The results show that the signal strength increases as the size of the light reflective film 16 increases. However, when the size of the light reflecting film 16 is enlarged beyond 2 cm to 3 cm, the advantages are diminished. This means that the area covered by the light reflecting film 16 does not need to be very large.

[0041] It is also beneficial to evaluate the effect of the size of the holes in the light reflecting film 16 required to inject the transport light. Similar to the evaluation above, as the hole size decreases, the signal strength increases. However, as the size of the hole decreases below 3-4 mm, the advantages decrease. Therefore, the holes do not need to be very small.

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

[0043] Images of the inclusions were taken for each of the cases where a light reflecting film functioning as both the light reflecting member 16 and the light reflecting member 17 was used and where it was not used. These images can be seen in Figure 10a) and Figure 10b), respectively. The light reflective film covered the portion of the input surface not occupied by the light input, the ultrasound source, and the 1 cm diameter light guide.

[0044] In an additional simulation, the simulation was performed in reflection mode. To simulate the signal light, a two-stage calculation was performed. First, the carrier light was simulated similarly to the simulation shown in FIG. Then, the tagged light is converted to power F US *K, where F US is the delivery fluence of the ultrasonic pulse, and K is an empirical tagging factor determined experimentally. The scattering and absorption properties of the light scattering medium were then selected for 85% oxygenated muscle tissue at a wavelength of 690 nm. The ultrasound pulse was modeled as a sphere with a radius of 2 mm. The CNR was then calculated assuming that ST1 is the tagged optical signal when the local oxygenation around the ultrasound pulse is set to 40% and ST2 is the tagged optical signal when the local oxygenation is 85%.

[0045] In FIG. 11, N=200, the CNR is shown as a function of ultrasound pulse depth for 40% oxygenation in the ultrasound pulse compared to an 85% baseline. It can be seen that the CNR increases obviously when the light reflecting member is applied compared to the case without the light reflecting member. The increase in CNR in the simulation results is in good agreement with the increase in CNR seen in the experimental results, with the CNR for the image acquired when the light reflecting member of FIG. 10 was used being 1.54, and the CNR for the image acquired when the light reflecting member was not used being 1.02.

[0046] It is beneficial to evaluate the distance between the laser injection site and the light detection site to assess whether an optimal distance exists in terms of CNR. Since optical filter suppression of the carrier light is not perfect, it may be optimal to have some distance between the light source and the photodetector. FIG. 12 shows the CNR as a function of the separation distance between the source and the photodetector for an ultrasound pulse at a depth of 5 cm and a filter suppression of 80 dB. The optimum separation distance in this case is 20 mm. In general, the optimum spacing depends on the scattering and absorption of the light scattering medium, the depth of the ultrasound pulse, the tagging factor 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 a light reflecting member near or adjacent to the light source injection site are not limited to acoustic photon tagging, but are also advantageous when performing optoacoustic imaging or tomography. In this case, the photoacoustic signal increases due to an increase in the amount of light in the light scattering medium 11 .

[0048] As mentioned above, to ensure sufficient signal levels in the device, it is important to minimize light losses between the light collection site and the photodetector. The light may be guided to the slow light filter using a light guide. However, slow light filters need to be contained in a cryogenic device and stored at low temperatures in order to function as intended.

[0049] The present invention is advantageously used where there is medical benefit in measuring deep body tissue oxygenation. For example, the present invention may be used in the management of ischemic stroke. In emergency medical situations, it is important to rapidly diagnose the presence of ischemic regions in a patient's brain. The majority of patients who suffer an ischemic stroke undergo thrombectomy therapy to remove the blood clot. Problematically, in severe cases, the patient may not be able to wake up during thrombectomy therapy and it is not immediately clear to the surgeon whether the treatment procedure was successful or not. In such cases, monitoring cerebral oxygenation status is advantageous as a means of feedback for therapeutic treatment.

[0050] Another medical indication in which the present invention would be advantageous is monitoring myocardial oxygenation status for emergency treatment in the event of suspected myocardial infarction.

[0051] Another medical indication in which the present invention would be advantageous is in sports medicine, to assess oxygen uptake and metabolism in muscles.

[0052] Another medical indication in which the present invention may be of advantage is in the field of oncology, where oxygen saturation may be altered in tumors compared to healthy tissue. For example, within a tumor there may be necrotic areas that exhibit low oxygen saturation.

[0053] Slow light filters rely on the use of optical transitions of dopant ions within a crystalline material. In order to create a frequency filter narrow enough to separate the frequency-shifted light from the original frequency light, the linewidth of this transition needs to be narrow so that the ends of the Lorentzian absorption curve do not overlap between the original laser frequency and the acoustically shifted frequency. The width of these transition lines is limited by the inverse of the coherence time of the transition. And if this coherence time is lifetime limited, the shorter the lifetime of the higher state, the broader the line will be. For ions in a crystal lattice, the shortened lifetime is due to the interaction of the ion with phonons, which are vibrational quanta that propagate within the crystal structure. If the higher state of the ion in the optical transition is close to another state, the ion can transition to this state by absorbing or emitting one or more phonons, making events involving more phonons increasingly unlikely. And since this maximum phonon energy is a property of the host crystal, the lifetime of the high state can be extended by using a "soft" crystal with a lower maximum phonon energy. In this case, phonon transitions in these crystals require interactions with a larger number of phonons, resulting in a slower transition rate and a longer lifetime for the higher states. By using soft crystals, more optical transitions of a particular ion are available for use in filtration. This allows more wavelengths to be used for the assessment of tissue oxygenation using the same crystal.

[0054] A particular dopant ion, for example lanthanum, can only be used to create a filter at a wavelength specific to that ion. Thus, by using a number of different dopant ions, for example lanthanum and europium, it becomes possible to customize where the filter can be made. Multiple ions can be made accessible to the same signal light by doping multiple ions into the same host crystal, ie, "co-doping," or by combining different crystals. This combination can be achieved either by stacking or layering the crystals, ie, "sandwiching," or by fusing different crystals together, ie, "splicing."

[0055] The structures created in the absorption profile can be either one or several notch filters that select each individual frequency-shifted sideband of the original light. It may also be a bandstop filter where the only absorption that is not excluded is at the original frequency. This is beneficial because the frequency shifted light of the first sideband (±1 ultrasonic frequency) accounts for only half of the light removed from the original frequency for a given ultrasonic pulse. However, this bandstop filter does not have as high a slow light effect for light passing through the filter as a notch filter.

[0056] Depending on the application and the ultrasound pulse, both filters may be beneficial.

[0057] The maximum width of any spectral structure requires that the other ground state or states of the ion be optically pumped sufficiently far apart in frequency. This may be an inherent property of the ions and crystals themselves, where the ground state naturally splits into multiple hyperfine states. In some cases, this splitting can be achieved by applying electric and / or magnetic fields to cause the generally degenerate states to split. By applying an electric and / or magnetic field, the lifetime of the ground state can be extended, and therefore the filter lifetime. This increased filter life reduces the amount of time during an imaging sequence that must be allocated to filter preparation. In this way, increased filter life directly results in faster imaging speeds. The creation of a slow light filter has the additional effect that one or more so-called "anti-holes" are simultaneously created. Such antiholes are positions on the crystal absorption profile where absorption is rather increased. At least one antihole is located at a frequency distance equal to the ground state division from the center frequency of the slow light filter. This occurs due to optical pumping, when more ions occupy this state. In this way, by applying the correct electric and / or magnetic fields, in addition to extending the hole lifetime, it is possible to improve the efficiency of the slow light filter by providing antiholes at the original laser frequency.

[0058] Inside a cryogenic device, a magnetic field can be generated using permanent magnets. This has the advantage that no power supply or cooling is required as with external electromagnets, and no additional heat load is created through leads as with superconducting electromagnet coils or capacitors. Furthermore, when the magnets are placed inside the cryogenic device, a stronger magnetic field can be generated with a smaller magnet than when they are placed outside the device. In this manner, the use of permanent magnets allows for both a reduction in the size and cooling power of the low temperature cryostat device being used.

[0059] The shape of the ultrasound pulse also has a large effect on the spatial resolution of the image and the intensity of the frequency-shifted light. This ultrasound pulse can then be chirped, which, together with a slow light filter, allows probing multiple spatial points with the same ultrasound pulse. This effectively reduces acquisition time and allows for faster image construction.

[0060] A chirped ultrasound pulse refers to a pulse that has a gradient of central frequency along its temporal and spatial profile. With short light pulses, the frequency-shifted light can also be chirped in this manner, with each frequency component associated with a different spatial region. Different frequencies of filtered light propagate at different velocities, so the frequency components can be separated in time using slow light filters. Different times in the readout then correspond to different probe positions in the tissue.

[0061] Multiple ultrasonic pulses may be generated simultaneously at different frequencies, with each pulse frequency having the same effect. And if each of these pulse frequencies is matched to a slow-light filter with a different time delay, multiple points can be imaged with the same light pulse.

[0062] To optimize the detected signal, it is important that the signal light is collected from the light scattering medium 11 and guided to the filter as efficiently as possible. To accomplish this, it is preferred to use a light guide device, such as a single light guide or a fiber optic bundle. And when this light guide arrangement is coupled into a cryostat holding a filter, it is important to minimize light loss and maintain insulation of the cold chamber. An embodiment that accomplishes this is shown in FIG. A hollow light guide 27 having reflective walls directs light from an external light guide into and out of the crystal 21 located within the cryostat. In this way, light rays 28 that do not strike the filter and subsequently diverge too much to be imaged on the photodetector are contained by the hollow light guide 27 . This arrangement allows a vacuum to be maintained within the hollow light guide 27, preserving thermal integrity. At the same time, the hollow light guide 27 brings the active light receiving surface of the filter closer to the outside of the cryostat. Since no image needs to be stored within the optical system, the hollow light guide 27 allows for minimal light losses to and from the filter.

[0063] When presenting the oxygen saturation data to a user, it is preferable to simultaneously display a conventional ultrasound scan (eg, a B-mode scan) and an image indicative of oxygen saturation, either on an absolute or relative scale. Oxygen saturation can be represented as a 2D intensity monochrome image or as 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 approach, the oxygen saturation image can be toggled on / off using software functions or gradually blended onto the ultrasound image.

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

[0065] As used in this specification and claims, the indefinite articles "a" and "an" mean "at least one," unless otherwise specified. The term "and / or" as used in this specification and claims refers to either or both of the elements conjointly present in some cases and non-conjointly present in other cases.

Claims

1. In a light measurement system in a subject, The ultrasonic transducer includes an ultrasonic transducer, a laser that emits light within a wavelength range of infrared light (IR), visible light, or ultraviolet light, an optical filter, a photodetector, and a light reflecting member; the ultrasound transducer is configured to direct an ultrasound field having an ultrasound frequency to at least one location within a tissue site of the subject; the laser is configured to direct light having an optical frequency into the tissue site of the subject; the optical filter is disposed upstream of the photodetector so as to suppress the optical frequency while transmitting light having an optical frequency frequency shifted by the frequency of the ultrasonic wave; the photodetector detects the light frequency-shifted by the frequency of the ultrasonic wave; An optical measurement system configured such that the optical reflecting member is positioned on a surface of the tissue site of the subject.

2. The light measurement system of claim 1 , wherein the light reflecting member is disposed on the surface of the tissue site to increase the amount of light detected by the light detector.

3. The optical measurement system of claim 1 , wherein the optical reflecting member is configured to be disposed around an injection point of the laser.

4. The optical measurement system of claim 1 , wherein the light reflecting member is positioned about a point where the frequency-shifted light exits the tissue site toward the optical filter.

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

6. at least one location within the tissue site is occupied by at least one ultrasound field; The optical measurement system of claim 1 , wherein the laser directs light into the tissue site such that at least a portion of the emitted light passes through the location.

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

8. 10. The optical measurement system of claim 1, wherein the laser and ultrasound source are configured to scan at least one location to obtain a map or image of optical contrast within the tissue site, e.g., information regarding oxygen saturation at the tissue site.

9. 2. The light measurement system according to claim 1, wherein the light reflecting member is made of at least one of glass, a metal mirror, a metal surface, a light reflecting synthetic resin, or a surface coated 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 of claim 10 , wherein the slow light filter structure comprises two different absorption lines to form a filter at two wavelengths.

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

13. 13. The optical measurement system of claim 12, wherein the host crystal is either doped with a single ion, co-doped with multiple ions, or composed of multiple different crystals spliced ​​or sandwiched together.

14. 11. The light measurement system of claim 10, wherein the slow light filter structure comprises a notch filter configuration that blocks all light except the frequency-shifted light.

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

16. The light measurement system of claim 10 , wherein an element that creates an electric and / or magnetic field is configured to affect the lifetime of the slow light filter structure.

17. 17. The optical measurement system of claim 16, wherein the electric and / or magnetic fields position the antiholes of the slow light filter structure at the original laser frequency of the laser.

18. 17. The optical measurement system of claim 16, wherein the electric and / or magnetic field creating element comprises a permanent magnet.

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