Image acquisition device and image acquisition method

Abstract

An image acquisition device 1A comprises a measurement unit 10 and a processing unit 20. For each coincidence event in which a pair of gamma-ray photons generated through an electron-positron pair annihilation event in a positron emitting radionuclide 81 are detected in coincidence by a first detector 11 and a second detector 12, the processing unit 20 determines a position where a gamma-ray photon has been Compton-scattered in a subject 90 selectively if the gamma-ray photon has reached the second detector 12 after being Compton-scattered in the subject 90 and the energy of the gamma-ray photon represented by a signal outputted from the second detector 12 is lower than a first threshold value lower than 511 keV, on the basis of a position and a time of the gamma-ray photon detection by each of the first detector 11 and the second detector 12, and the position of the positron emitting radionuclide 81. Thus, an image acquisition device and an image acquisition method are achieved that make it possible to acquire a tomographic image representing anatomical information pertaining to a subject without performing image reconstruction processing.

Description

Image acquisition device and image acquisition method

[0001] The present disclosure relates to an image acquisition device and an image acquisition method.

[0002] Nuclear medicine diagnostic devices such as PET (Positron Emission Tomography) devices and SPECT (Single Photon Emission Computed Tomography) devices can acquire tomographic images of a subject administered with a drug labeled with a positron-emitting nuclide or a single-photon-emitting nuclide. The tomographic images acquired by these nuclear medicine diagnostic devices represent the distribution of the positron-emitting nuclide or single-photon-emitting nuclide (drug distribution) in the subject, and can be used to diagnose the health condition of the subject.

[0003] An X-ray CT scanner can also acquire three-dimensional tomographic images of a subject. The tomographic images acquired by an X-ray CT scanner represent anatomical information of the subject. Hereinafter, the tomographic images will be referred to as three-dimensional tomographic images.

[0004] The PET scanner and the X-ray CT scanner are used in combination, and the image quality of the PET image is improved by correcting the PET image using the anatomical information acquired by the X-ray CT scanner. However, since the X-ray CT scanner is expensive, research and development is being conducted on inexpensive devices that can acquire tomographic images that represent the anatomical information of the subject.

[0005] Non-Patent Document 1 and Patent Document 1 each describe an apparatus capable of acquiring both a tomographic image showing the distribution of positron-emitting nuclides in a subject and a tomographic image showing anatomical information.

[0006] The device described in Non-Patent Document 1 has the configuration of a PET device in which a number of detectors are arranged around a measurement space in which a subject is placed. ２ SiO ５Using detectors each having a (Ce) scintillator, gamma rays with energy 307 keV or 202 keV emitted from the Lu contained in the LSO scintillator of each detector are detected by another detector, and gamma rays that have passed through the subject are detected. This device then performs image reconstruction processing based on the detection results of the gamma rays with energy 307 keV or 202 keV, thereby obtaining tomographic images that represent anatomical information of the subject.

[0007] The device described in Patent Document 1 utilizes an Electron Tracking Compton Camera (ETCC). A typical Compton camera estimates that gamma rays have come from a position on a conical surface called a Compton cone, based on energy information from a scatterer and an absorber. In contrast, ETCC is said to be able to uniquely identify the direction of gamma ray arrival by tracking the tracks of recoil electrons using a gas detector.

[0008] This device uses ETCC to detect the direction of arrival of gamma rays that have been Compton scattered in the subject and have reduced energy, among gamma rays generated in the subject administered with a drug labeled with a positron-emitting nuclide.The device then uses an analytical or statistical method, i.e., performs image reconstruction processing, based on the detection results of the gamma ray arrival direction by ETCC to obtain a tomographic image that represents anatomical information of the subject.

[0009] Patent No. 6990412

[0010] Mohammadreza Teimoorisichani et al., "A CT-less approach to quantitative PET imaging using the LSO intrinsic radiation for long-axial FOV PET scanners", Med. Phys., Vol.49, pp.309-323 (2022)Gerard Arino-Estrada et al., "First Cerenkov charge-induction (CCI) TlBr detector for TOF-PET and proton range verification", Phys. Med. Biol., 64 175001 (2019)

[0011] In both the techniques described in Non-Patent Document 1 and Patent Document 1, it is necessary to perform image reconstruction processing based on gamma ray detection results in order to acquire tomographic images that represent anatomical information of the subject. The tomographic images acquired by the image reconstruction processing have a reduced image quality due to the image reconstruction processing, resulting in degraded anatomical information.

[0012] An object of the present invention is to provide an image acquisition device and an image acquisition method that can acquire tomographic images that represent anatomical information of a subject without performing image reconstruction processing.

[0013] An embodiment of the present invention is an image acquisition device, the image acquisition device comprising: (1) a measurement unit including a first detector and a second detector each for detecting gamma-ray photons, and outputting signals representing a detection position, a detection time, and a gamma-ray photon energy when the first detector and the second detector each detect a gamma-ray photon; and (2) a processing unit processing signals output from the first detector and the second detector, (3) in a first measurement mode, the measurement unit outputs signals representing a detection position, a detection time, and a gamma-ray photon energy of the gamma-ray photon by the first detector and the second detector, with a subject placed between the first detector and the second detector and a positron-emitting nuclide placed between the first detector or the second detector and the subject; and (4) in the first measurement mode, the processing unit outputs signals representing a pair of electrons generated by annihilation of a positron in the positron-emitting nuclide. For each coincidence event in which the first detector and the second detector simultaneously count gamma ray photons, when the gamma ray photon that arrives at one of the first detector and the second detector arrives without undergoing Compton scattering within the subject, and the gamma ray photon that arrives at the other detector arrives after undergoing Compton scattering within the subject, and the gamma ray photon energy represented by the signal output from the other detector is lower than a first threshold value that is lower than 511 keV, the positions where the gamma ray photons have been Compton scattered within the subject are selectively determined based on the detection positions and detection times of the gamma ray photons by the first detector and the second detector, respectively, and the positions of the positron-emitting nuclides, and a first tomographic image is created that represents the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence events.

[0014] An embodiment of the present invention is an image acquisition device, which includes: (1) a measurement unit including a first detector and a second detector each for detecting a gamma ray photon, and outputting a signal representing a detection position, a detection time, and a gamma ray photon energy when the first detector and the second detector each detect a gamma ray photon; (2) a processing unit for processing the signals output from the first detector and the second detector; and (3) the measurement unit is configured to measure a distance between the first detector and the second detector and a subject who has been administered a drug labeled with a positron-emitting nuclide, the distance being measured between the first detector and the second detector and the subject. With a positron-emitting nuclide placed between the first detector and the second detector, signals are output that represent the detection position, detection time, and gamma-ray photon energy of the gamma-ray photon by each of the first detector and the second detector, and (4) the processing unit, for each coincidence event in which the first detector and the second detector simultaneously count a pair of gamma-ray photons generated by an electron-positron annihilation event in the positron-emitting nuclide, determines whether (a) the gamma-ray photon that arrived at one of the first detector and the second detector arrived without Compton scattering within the subject, and whether the gamma-ray photon that arrived at the other detector (b) determining the position where the gamma-ray photon has been Compton scattered in the subject based on the detection position and detection time of the gamma-ray photon by each of the first detector and the second detector, and the position of a positron-emitting nuclide placed between the first detector or the second detector and the subject, when the gamma-ray photon arrives after Compton scattering in the subject and the gamma-ray photon energy represented by the signal output from the other detector is lower than a first threshold value lower than 511 keV; When a gamma ray photon arrives without undergoing Compton scattering within the subject, the position where the annihilation event occurred is determined based on the detection position and detection time of the gamma ray photon by the first detector and the second detector, respectively; (c) a first tomographic image is created that represents the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence counting events; a second tomographic image is created that represents the distribution of annihilation event positions in the subject determined for each of the multiple coincidence counting events; and the second tomographic image is corrected based on the first tomographic image.

[0015] An embodiment of the present invention is an image acquisition method, the image acquisition method comprising: (1) a measurement step of using a first detector and a second detector, each of which detects a gamma-ray photon, and outputting a signal representing a detection position, a detection time, and a gamma-ray photon energy when the first detector and the second detector each detect a gamma-ray photon; and (2) a processing step of processing the signals output from the first detector and the second detector, wherein (3) in a first measurement mode, the measurement step outputs signals representing a detection position, a detection time, and a gamma-ray photon energy by the first detector and the second detector, with a subject placed between the first detector and the second detector and a positron-emitting nuclide placed between the first detector or the second detector and the subject; and (4) in the processing step, in the first measurement mode, For each coincidence event in which the first detector and the second detector simultaneously count a pair of gamma ray photons generated, when the gamma ray photon arriving at one of the first and second detectors arrives without undergoing Compton scattering within the subject, and the gamma ray photon arriving at the other detector arrives after Compton scattering within the subject, and the gamma ray photon energy represented by the signal output from the other detector is lower than a first threshold value lower than 511 keV, the positions at which the gamma ray photons have been Compton scattered within the subject are selectively determined based on the detection positions and detection times of the gamma ray photons by the first and second detectors, respectively, and the positions of the positron-emitting nuclides, and a first tomographic image is created that represents the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence events.

[0016] An embodiment of the present invention is an image acquisition method, which includes: (1) a measurement step of using a first detector and a second detector, each of which detects a gamma-ray photon, and outputting a signal representing a detection position, a detection time, and a gamma-ray photon energy when the first detector and the second detector detect a gamma-ray photon; and (2) a processing step of processing the signals output from the first detector and the second detector, wherein (3) the measurement step includes placing a subject who has been administered a drug labeled with a positron-emitting nuclide between the first detector and the second detector, and processing the signals output from the first detector or the second detector. and the subject, and outputs signals representing the detection position, detection time, and gamma-ray photon energy of the gamma-ray photon by each of the first detector and the second detector, with a positron-emitting nuclide placed between the first detector and the subject; and (4) a processing step, for each coincidence event in which the first detector and the second detector simultaneously count a pair of gamma-ray photons generated by an annihilation event of an electron and a positron in the positron-emitting nuclide, (a) a gamma-ray photon that arrived at one of the first detector and the second detector arrived without Compton scattering within the subject, and When a gamma-ray photon arriving at the detector arrives after Compton scattering within the subject and the gamma-ray photon energy represented by the signal output from the other detector is lower than a first threshold value lower than 511 keV, the position where the gamma-ray photon has been Compton scattered within the subject is selectively determined based on the detection position and detection time of the gamma-ray photon by each of the first detector and the second detector, and the position of a positron-emitting nuclide placed between the first detector or the second detector and the subject; (b) determining whether the gamma-ray photon arrives at both the first detector and the second detector. When the incident gamma ray photons arrive without undergoing Compton scattering within the subject, the position where the pair annihilation event occurred is determined based on the detection positions and detection times of the gamma ray photons by the first detector and the second detector, respectively; (c) creating a first tomographic image representing the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence counting events, creating a second tomographic image representing the distribution of the positions where pair annihilation events occurred in the subject determined for each of the multiple coincidence counting events, and correcting the second tomographic image based on the first tomographic image.

[0017] According to the embodiment of the present invention, it is possible to acquire a tomographic image that represents anatomical information of a subject without performing image reconstruction processing.

[0018] FIG. 1 is a diagram showing the configuration of an image acquisition apparatus 1A of the first embodiment (particularly a diagram explaining the acquisition of a first tomographic image in the first measurement mode). FIG. 2 is a diagram explaining a method for determining the position of Compton scattering of gamma-ray photons in a subject. FIG. 3 is a diagram showing the configuration of the image acquisition apparatus 1A of the first embodiment (particularly a diagram explaining the acquisition of a second tomographic image in the second measurement mode). FIG. 4 is a diagram showing the configuration of an image acquisition apparatus 1B of the second embodiment. FIG. 5 is a diagram showing the configuration of an image acquisition apparatus 1C of the third embodiment. FIG. 6 is a diagram showing the configuration of an image acquisition apparatus 1D of the fourth embodiment. FIG. 7 is a diagram showing the configuration of an image acquisition apparatus 1E of the fifth embodiment. FIG. 8 is a diagram showing the configuration of an image acquisition apparatus 1F of the sixth embodiment. FIGS. 9(a) and 9(b) are diagrams explaining a change in the field of view of the apparatus when a positron-emitting nuclide 81 is moved in a direction parallel to the detection surface of the first detector 11 in the measurement unit 10F of the image acquisition apparatus 1F of the sixth embodiment. 10A and 10B are diagrams illustrating the image quality of a first tomographic image when a positron-emitting nuclide 81 is moved in a direction perpendicular to the detection surface of the first detector 11 in the measurement unit 10F of the image acquisition device 1F of the sixth embodiment. FIG. 11 is a diagram illustrating the configuration and arrangement of the measurement unit assumed in the simulation. FIG. 12 is a diagram illustrating the configuration of a phantom assumed as the subject 90 in the simulation. FIG. 13A and 13B are diagrams illustrating first tomographic images obtained in the simulation. FIG. 14A and 14B are diagrams illustrating first tomographic images obtained in the simulation. FIG. 15A and 15B are diagrams illustrating first tomographic images obtained in the simulation. FIG. 16A and 16B are diagrams illustrating first tomographic images obtained in the simulation.

[0019] Hereinafter, embodiments of an image acquisition device and an image acquisition method will be described in detail with reference to the accompanying drawings. In the description of the drawings, the same elements are designated by the same reference numerals, and duplicate explanations will be omitted. The present invention is not limited to these examples, but is defined by the claims, and all modifications within the meaning and scope equivalent to the claims are intended to be included.

[0020] (First embodiment)

[0021] 1 is a diagram showing the configuration of an image acquisition device 1A according to a first embodiment. The image acquisition device 1A includes a measurement unit 10, a processing unit 20, and a display unit 30. The measurement unit 10 includes a first detector 11 and a second detector 12 that are arranged opposite each other with a subject 90 in between. Each of the first detector 11 and the second detector 12 detects gamma-ray photons, and when a gamma-ray photon is detected, it outputs a signal representing the detection position, the detection time, and the gamma-ray photon energy.

[0022] The processing unit 20 processes the signals output from the first detector 11 and the second detector 12 to create a tomographic image of the subject 90. The display unit 30 displays the tomographic image etc. created by the processing unit 20. The processing unit 20 and the display unit 30 may be configured by, for example, a computer.

[0023] For example, a Cherenkov detector is used as each of the first detector 11 and the second detector 12. The Cherenkov detector is a detector that emits a Cherenkov radiator (for example, lead glass, lead fluoride PbF ２ , hafnium oxide HfO ２ The Cherenkov detector may be, for example, a two-dimensional array of small detectors that do not have position detection capability individually, or may be a combination of a Cherenkov radiator and a multi-anode MCP-PMT.

[0024] The Cherenkov detector is also a BGO (Bi ４ Ge ３ O １２The detector may be one that uses a slow scintillator such as a fluorine-containing fluorine atom (FWHM) as a Cherenkov radiator. A slow scintillator interacts with gamma rays to first emit Cherenkov light and then emits scintillation light, so it can be used as a Cherenkov radiator and achieve high time resolution. This allows for the construction of a detector that is less expensive than an LSO scintillator or the like.

[0025] Furthermore, each of the first detector 11 and the second detector 12 may be, for example, a semiconductor detector with high time resolution. The semiconductor detector with high time resolution may use, for example, thallium bromide (TlBr) and be equipped with a charge collection electrode and a high time resolution photodetector. Such a detector may be, for example, the one described in Non-Patent Document 2. The use of a semiconductor detector is expected to improve energy resolution, which in turn improves the ability to remove scattered components and improve image quality.

[0026] Considering that the spatial resolution of a tomographic image acquired by a nuclear medicine diagnostic device such as a PET device is about 3 to 5 mm, it is desirable that the spatial resolution required for each of the first detector 11 and the second detector 12 be equal to or better than that. Similarly, it is desirable that the temporal resolution required for each of the first detector 11 and the second detector 12 be 20 to 35 ps or less in terms of coincidence temporal resolution.

[0027] When the first detector 11 and the second detector 12 each include a Cherenkov radiator or a scintillator, it is preferable that the first detector 11 and the second detector 12 output a signal representing the position (detection position) and time (detection time) at which a gamma ray interacts with the Cherenkov radiator or the scintillator, rather than the position and time at which the Cherenkov light or the scintillation light is detected. In this case, the detection position is expressed as a three-dimensional coordinate value that specifies not only the position in two directions parallel to the detection surface of the detector, but also the position in a direction perpendicular to the detection surface.

[0028] The detection surface of each of the first detector 11 and the second detector 12 preferably has a size larger than the subject 90 (or a region of interest in the subject 90). For example, in the case of an image acquisition device that acquires tomographic images of a human brain, the detection surface of each of the first detector 11 and the second detector 12 preferably has a size that is approximately the same as or larger than the size of a human brain.

[0029] The image acquisition device 1A and the image acquisition method using the same acquire, in a first measurement mode, a tomographic image (first tomographic image) of the subject 90. The image acquisition device 1A and the image acquisition method can also acquire, in a second measurement mode, a tomographic image (second tomographic image) of the subject 90.

[0030] The first tomographic image is an image that represents the distribution of Compton scattering positions in the subject 90, and represents anatomical information of the subject 90. The second tomographic image represents the distribution of positron-emitting nuclides (distribution of a drug) in the subject 90, and can be used to diagnose the health condition of the subject 90.

[0031] The acquisition of the first tomographic image in the first measurement mode is performed by a first measurement step and a first processing step as follows: Fig. 1 is a diagram showing the configuration of an image acquisition device 1A of the first embodiment, and is a diagram particularly illustrating the acquisition of the first tomographic image in the first measurement mode.

[0032] In the first measurement step, the subject 90 is placed between the first detector 11 and the second detector 12. At this time, the subject 90 does not need to be administered with a positron-emitting nuclide. A positron-emitting nuclide 81 is placed between the subject 90 and the first detector 11 or the second detector 12. It is preferable to use a positron-emitting nuclide 81 that is as small as possible. In this figure, the positron-emitting nuclide 81 is placed between the first detector 11 and the subject 90.

[0033] A positron emitted from the positron-emitting nuclide 81 immediately annihilates with a nearby electron, and this electron-positron annihilation event generates a pair of gamma-ray photons traveling in opposite directions. When the first detector 11 and the second detector 12 detect a gamma ray, they each output a signal representing the detection position, detection time, and gamma-ray photon energy of the gamma ray photon.

[0034] In the first processing step, for each coincidence event in which the first detector 11 and the second detector 12 simultaneously count a pair of gamma ray photons generated by an electron-positron annihilation event in the positron-emitting nuclide 81, the processing unit 20 determines the position at which the gamma ray photon was Compton scattered based on the detection position and detection time of the gamma ray photon by each of the first detector 11 and the second detector 12 and the position of the positron-emitting nuclide 81, assuming that the gamma ray photon arriving at one of the first detector 11 and the second detector 12 arrived without Compton scattering within the subject, and the gamma ray photon arriving at the other detector arrived after Compton scattering within the subject.

[0035] The processing unit 20 then creates a first tomographic image that represents the distribution of Compton scattering positions in the subject 90 obtained for each of the multiple coincidence events. This first tomographic image represents anatomical information of the subject 90.

[0036] The processing unit 20 can determine whether or not a gamma ray photon arriving at the first detector 11 or the second detector 12 has undergone Compton scattering based on one or more of the position of the positron-emitting nuclide 81, the magnitude of the energy of the gamma ray photon, and the time of detection of the gamma ray photon by each of the first detector 11 and the second detector 12.

[0037] 1 , if a positron-emitting nuclide 81 is placed between the first detector 11 and the subject 90, it can be determined that the gamma rays arriving at the first detector 11 have not undergone Compton scattering within the subject. The energy of a pair of gamma ray photons generated by electron-positron annihilation is 511 keV, but the energy of the gamma rays is reduced by Compton scattering. Therefore, it can be determined whether or not the gamma rays have undergone Compton scattering based on the magnitude of the energy of the gamma rays. It can be determined whether or not the gamma ray has undergone Compton scattering within the subject based on the relationship between the time of detection of the gamma ray photons by the first detector 11 and the second detector 12, respectively.

[0038] 2 is a diagram for explaining a method for determining the position where a gamma ray photon has been Compton scattered within a subject. The position of the positron-emitting nuclide 81 is defined as P, and the position where the gamma ray is detected by the first detector 11 is defined as R. １ and the detection position of the gamma ray by the second detector 12 is R ２ The position where the gamma ray is Compton scattered is defined as C. The time when the gamma ray is detected by the first detector 11 is defined as t １ and the detection time of the gamma ray by the second detector 12 is t ２ Let's say.

[0039] Of the pair of gamma ray photons generated by the annihilation of an electron and a positron in the positron-emitting nuclide 81, the flight distance of one gamma ray is from position P to position R. １ Distance d to １ The flight distance of the other gamma ray is the distance d from position P to position C. ２１ and from position C to position R ２ Distance d to ２２ The sum of (d ２１ +d ２２ )

[0040] The difference in flight distance of a pair of gamma ray photons is L (= d ２１ +d ２２ -d １ ) is the difference in the detection times of these gamma-ray photons (t ２ -t １ ) multiplied by the speed of light c. Also, the distance between position P and position R １The line segment connecting the positions P and C is parallel to the line segment connecting the positions P and C. From the above, the detection position R of the gamma ray photon by the first detector 11 is １ and detection time t １ , the detection position R of the gamma ray photon by the second detector 12 ２ and detection time t ２ , and the position P of the positron-emitting nuclide 81, the position C at which the gamma ray photon is Compton scattered can be determined.

[0041] When neither of the pair of gamma-ray photons undergoes Compton scattering, the Compton scattering position determined by the method described in Fig. 2 coincides with the position P of the positron-emitting nuclide 81. This position is outside the subject 90, and can therefore be easily excluded.

[0042] The spatial resolution of the detection of the Compton scattering position C obtained as described above depends on the spatial resolution of each of the first detector 11 and the second detector 12, and the time resolution of each of the first detector 11 and the second detector 12. This will be explained below.

[0043] 1 and 2, an xyz orthogonal coordinate system is set having x- and y-axes parallel to the detection surfaces of the first detector 11 and the second detector 12. Position R １ The coordinate values ​​of (R １ｘ , R １ｙ , R １ｚ ) and position R ２ The coordinate values ​​of (R ２ｘ , R ２ｙ , R ２ｚ ), and the coordinate value of position C is (C ｘ , C ｙ , C ｚ ), and the difference in flight distance of a pair of gamma ray photons (d ２１ +d ２２ -d １ ) is defined as L.

[0044] The spatial resolution of the position detection of the first detector 11 is (dR １ｘ , dR １ｙ , dR １ｚ ), and the spatial resolution of the position detection of the second detector 12 is (dR ２ｘ , dR２ｙ , dR ２ｚ ), and the resolution of detecting the difference in flight distance L of a pair of gamma ray photons is dL. The spatial resolution of detecting position C is (ΔC ｘ , ΔC ｙ , ΔC ｚ ) then ΔC ｘ is expressed by the following formula (1): ｙ , ΔC ｚ is expressed by a similar formula. ｘ The formula is shown only for

[0045] In many cases, the resolution dL of the flight distance difference L based on the time resolution of the detector is larger than the spatial resolution of the detector. Therefore, as can be seen from the above equation (1), in order to reduce the spatial resolution of the detection of the position C, ｘ It is desirable that the partial derivative value of C is small. ｘ is expressed by the following formula (2): ｘ The partial derivative of is expressed by the following equation (3): Here, the position P of the positron-emitting nuclide 81 is set as the origin of the xyz orthogonal coordinate system.

[0046] Equation (3) can be rearranged to obtain equation (4) below, where θ is the Compton scattering angle (see FIG. 2). The Compton scattering angle θ is the angle of the direction in which the gamma rays travel after scattering position C relative to the direction in which the gamma rays travel from position P of the positron-emitting nuclide 81 to scattering position C.

[0047] As can be seen from equation (4), the smaller the Compton scattering angle θ (the closer to 0), the greater the C ｘ As mentioned above, the partial derivative of C by L is large. ｘ The larger the partial differential of , the greater the spatial resolution of detecting the position C. Also, the smaller the Compton scattering angle θ, the higher the energy of the gamma ray photon after Compton scattering. Therefore, in order to reduce the spatial resolution of detecting the position C, ｘ It is preferable that the partial derivative of θ is small, the Compton scattering angle θ is large, and the energy of the gamma ray photons detected by the second detector 12 is low.

[0048] Therefore, in the first processing step, when the gamma ray photon energy represented by the signal output from the second detector 12 is even lower than a first threshold E1 which is lower than 511 keV, the processing unit 20 selectively determines the position C where the gamma ray photon has Compton scattered within the subject based on the detection position and detection time of the gamma ray photon by each of the first detector 11 and the second detector 12 and the position of the positron-emitting nuclide 81.

[0049] In this way, the spatial resolution for detecting the position C can be reduced, and a first tomographic image with excellent spatial resolution can be acquired. The first threshold E1 is preferably sufficiently lower than 511 keV, and can be set to, for example, 450 keV, 400 keV, 370 keV, 350 keV, 330 keV, 300 keV, etc. It is also preferable that the first threshold E1 can be set arbitrarily.

[0050] The acquisition of the second tomographic image in the second measurement mode is performed by a second measurement step and a second processing step as follows: Fig. 3 is a diagram showing the configuration of the image acquisition device 1A of the first embodiment, and is a diagram particularly illustrating the acquisition of the second tomographic image in the second measurement mode.

[0051] In the second measurement step, a subject 90 administered with a drug labeled with a positron-emitting nuclide 83 is placed between the first detector 11 and the second detector 12. At this time, the positron-emitting nuclide does not need to be placed outside the subject 90. A pair of gamma-ray photons traveling in opposite directions is generated by an annihilation event of an electron and a positron in the positron-emitting nuclide 83 that labels the drug administered to the subject 90. When the first detector 11 and the second detector 12 detect a gamma ray, they each output a signal representing the detection position, detection time, and gamma-ray photon energy of the gamma-ray photon.

[0052] The positron-emitting nuclide 83 that labels the drug administered to the subject 90 in the second measurement step is preferably the same type as the positron-emitting nuclide 81 placed between the first detector 11 and the subject 90 in the first measurement step. By making the positron-emitting nuclide 83 and the positron-emitting nuclide 81 the same type, only one type of positron-emitting nuclide needs to be prepared, which makes measurement preparation easier. Note that the positron-emitting nuclide 81 may also be a calibration positron-emitting nuclide such as 68Ge / 68Ga.

[0053] In the second processing step, the processing unit 20 determines the position where the annihilation event occurred, for each coincidence event in which the first detector 11 and the second detector 12 simultaneously count a pair of gamma-ray photons generated by an electron-positron annihilation event in the positron-emitting nuclide 83. The position where the annihilation event occurred for each coincidence event can be determined based on the difference in the detection times of the gamma-ray photons by the first detector 11 and the second detector 12 on a line segment connecting the detection positions of the gamma-ray photons by the first detector 11 and the second detector 12.

[0054] The processing unit 20 then creates a second tomographic image that represents the distribution of positions of occurrence of pair annihilation events determined for each of the multiple coincidence counting events in the subject 90. Because the first detector 11 and the second detector 12 have good time resolution, the second tomographic image, which is a three-dimensional tomographic image, can be acquired without performing image reconstruction processing.

[0055] This second tomographic image represents the distribution of positron-emitting nuclides 83 (distribution of the drug) in the subject 90, and can be used to diagnose the health condition of the subject 90. The processing unit 20 further corrects the second tomographic image based on the first tomographic image, thereby obtaining the second tomographic image after correction for the gamma-ray absorption distribution in the subject 90.

[0056] Either the acquisition of the first tomographic image in the first measurement mode or the acquisition of the second tomographic image in the second measurement mode may be performed first, however, if the acquisition of the second tomographic image in the second measurement mode is performed first, an annihilation event in the positron-emitting nuclide 83 administered to the subject 90 at that time may affect the acquisition of the first tomographic image in the first measurement mode later, so it is preferable to perform the acquisition of the first tomographic image in the first measurement mode first.

[0057] In this embodiment, a tomographic image (first tomographic image) representing anatomical information of the subject 90 can be acquired without performing image reconstruction processing, thereby avoiding degradation of image quality due to image reconstruction processing and preventing deterioration of anatomical information. Furthermore, a tomographic image (second tomographic image) representing the distribution of positron-emitting nuclides (distribution of drug) in the subject can be acquired without performing image reconstruction processing.

[0058] In this embodiment, since a large-scale rotation mechanism such as that required in an X-ray CT scanner is not required, the device can be made compact and inexpensive. Furthermore, compared to the case where an X-ray CT scanner is used, this embodiment can reduce the amount of radiation exposure to the subject.

[0059] The device described in Non-Patent Document 1 has a PET device configuration in which many detectors are arranged around a measurement space in which a subject is placed, making it difficult to reduce its size, and also makes it difficult to reduce its price because it uses an LSO scintillator containing lutetium (Lu), a rare material. In contrast, the present embodiment does not have these problems and can be reduced in size and price.

[0060] The device described in Patent Document 1 uses a gas detector, making it difficult to improve detection efficiency. In contrast, this embodiment does not have this problem and makes it possible to improve detection efficiency.

[0061] Furthermore, since the Compton scattering position C is selectively determined for events in which the energy of Compton scattered gamma ray photons is lower than the first threshold, it is possible to acquire a first tomographic image with excellent spatial resolution.

[0062] Second Embodiment

[0063] 4 is a diagram showing the configuration of an image acquisition device 1B according to the second embodiment. The image acquisition device 1B includes a measurement unit 10, a processing unit 20, and a display unit 30. Compared to the first embodiment, the second embodiment differs in that the Compton scattering position in the subject 90 and the position of annihilation event occurrence in the subject 90 are determined during a common period.

[0064] In the measurement step, a subject 90 who has been administered a drug labeled with a positron-emitting nuclide 83 is placed between the first detector 11 and the second detector 12. Also, a positron-emitting nuclide 81 is placed between the first detector 11 or the second detector 12 and the subject 90. In this figure, the positron-emitting nuclide 81 is placed between the first detector 11 and the subject 90.

[0065] By using the same type of positron-emitting nuclide 83 that labels the drug administered to the subject 90 and the positron-emitting nuclide 81 placed between the first detector 11 and the subject 90, only one type of positron-emitting nuclide needs to be prepared, which facilitates preparation for measurement. The positron-emitting nuclide 81 may be a calibration positron-emitting nuclide such as 68Ge / 68Ga.

[0066] A pair of gamma ray photons traveling in opposite directions is generated by an annihilation event of an electron and a positron in each of the positron-emitting nuclides 81 and 83. When the first detector 11 and the second detector 12 detect a gamma ray, they each output a signal representing the detection position, detection time, and gamma ray photon energy of the gamma ray photon.

[0067] In the processing step, the processing unit 20 performs the following processing for each coincidence event in which the first detector 11 and the second detector 12 simultaneously count a pair of gamma ray photons generated by an electron-positron annihilation event in the positron-emitting nuclide 81 or the positron-emitting nuclide 83.

[0068] The processing unit 20 determines whether or not a gamma ray photon arriving at the first detector 11 or the second detector 12 has undergone Compton scattering based on one or more of the position of the positron-emitting nuclide 81, the magnitude of the energy of the gamma ray photon, and the detection time of the gamma ray photon by each of the first detector 11 and the second detector 12.

[0069] As a result of the above determination, when the gamma ray photon that arrived at one of the first detector 11 and the second detector 12 arrived without undergoing Compton scattering within the subject, and the gamma ray photon that arrived at the other detector arrived after undergoing Compton scattering within the subject, and the gamma ray photon energy represented by the signal output from the other detector is even lower than a first threshold E1 that is lower than 511 keV, the processing unit 20 selectively determines that the pair of gamma ray photons arrived from a positron-emitting nuclide 81 outside the subject 90, and determines the position where the gamma ray photon has Compton scattered in the subject 90 based on the detection position and detection time of the gamma ray photon by each of the first detector 11 and the second detector 12 and the position of the positron-emitting nuclide 81, by the calculation described using Figure 2.

[0070] On the other hand, if the result of the above judgment is that the gamma ray photons that arrived at both the first detector 11 and the second detector 12 arrived without undergoing Compton scattering within the subject, the processing unit 20 determines that the pair of gamma ray photons came from a positron-emitting nuclide 83 within the subject 90, and determines the position where the annihilation event occurred in the subject 90 based on the detection position and detection time of the gamma ray photons by the first detector 11 and the second detector 12, respectively.

[0071] After performing the above processing on multiple coincidence events, the processing unit 20 creates a first tomographic image representing the distribution of Compton scattering positions in the subject 90, and also creates a second tomographic image representing the distribution of positions where annihilation events occur in the subject 90.

[0072] The first tomographic image represents anatomical information of the subject 90. The second tomographic image represents the distribution of positron-emitting nuclides 83 (distribution of a drug) in the subject 90, and can be used to diagnose the health condition of the subject 90. The processing unit 20 further corrects the second tomographic image based on the first tomographic image, thereby obtaining the second tomographic image after correction for the gamma-ray absorption distribution in the subject 90.

[0073] In the second embodiment, in addition to achieving the same effects as the first embodiment, the Compton scattering position in the subject 90 and the position where the annihilation event occurs in the subject 90 can be determined in a common period, thereby shortening the time required to restrain the subject 90.

[0074] (Third embodiment)

[0075] 5 is a diagram showing the configuration of an image acquisition apparatus 1C according to the third embodiment. The image acquisition apparatus 1C includes a measurement unit 10, a processing unit 20, and a display unit 30. Compared to the previous embodiments, the third embodiment differs in that a positron-emitting nuclide 81 is placed between the first detector 11 and the subject 90, and a positron-emitting nuclide 82 is also placed between the second detector 12 and the subject 90. It is preferable that the positron-emitting nuclide 81 and the positron-emitting nuclide 82 are of the same type.

[0076] The processing unit 20 determines whether a gamma ray photon arriving at the first detector 11 or the second detector 12 has undergone Compton scattering and whether the gamma ray photon was generated in a positron-emitting nuclide 81 or a positron-emitting nuclide 82, based on one or more of the positions of the positron-emitting nuclide 81, 82, the magnitude of the energy of the gamma ray photon, and the detection times of the gamma ray photon by each of the first detector 11 and the second detector 12. Based on the determination result, the processing unit 20 determines the position where the gamma ray photon has been Compton scattered in the subject 90.

[0077] In this embodiment, the positron-emitting nuclides 81 and 82 can be arranged symmetrically with respect to the subject 90, so that a first tomographic image of higher quality can be acquired. Furthermore, the number of Compton scattering events per unit time in the subject 90 increases, so that the measurement time can be shortened.

[0078] (Fourth embodiment)

[0079] 6 is a diagram showing the configuration of an image acquisition device 1D according to the fourth embodiment. The image acquisition device 1D includes a measurement unit 10D, a processing unit 20, and a display unit 30. Compared to the previous embodiments, the image acquisition device 1D according to the fourth embodiment differs in that it includes a measurement unit 10D instead of the measurement unit 10.

[0080] The measurement unit 10D includes a first detector 11D and a second detector 12. The detection surface of the first detector 11D is narrower than that of the second detector 12. A positron-emitting nuclide 81 is placed between the first detector 11D and the subject 90. When a gamma-ray photon is detected, the first detector 11D and the second detector 12 each output a signal representing the detection position, the detection time, and the gamma-ray photon energy.

[0081] In order to acquire the first tomographic image, it is sufficient that one of a pair of gamma ray photons generated by an electron-positron annihilation event in the positron-emitting nuclide 81 is incident on the first detector 11D and the other is incident on the subject 90 (or a region of interest within the subject 90). Therefore, as long as this condition is satisfied, the detection surface of the first detector 11D can be made narrower. The closer the position at which the positron-emitting nuclide 81 is placed to the first detector 11D, the narrower the detection surface of the first detector 11D can be made. Since the first detector 11D can be made small in this way, the image acquisition device 1D can be configured at low cost.

[0082] Fifth Embodiment

[0083] 7 is a diagram showing the configuration of an image acquisition device 1E according to a fifth embodiment. The image acquisition device 1E includes a measurement unit 10E, a processing unit 20, and a display unit 30. Compared to the previous embodiments, the image acquisition device 1E according to the fifth embodiment differs in that it includes a measurement unit 10E instead of the measurement unit 10.

[0084] The measurement unit 10E includes a shield 13 in addition to the first detector 11 and the second detector 12. The shield 13 prevents gamma ray photons backscattered from either the first detector 11 or the second detector 12 from entering the other detector.

[0085] The shield 13 is placed between the first detector 11 and the second detector 12 at a position that does not interfere with measurement of the Compton scattering position in the subject 90 and measurement of the position where annihilation events occur in the subject 90. The shield 13 is preferably a plate-shaped shield made of a high-density material (e.g., lead) that can block gamma rays.

[0086] Sixth Embodiment

[0087] 8 is a diagram showing the configuration of an image acquisition device 1F according to the sixth embodiment. The image acquisition device 1F includes a measurement unit 10F, a processing unit 20, and a display unit 30. Compared to the previous embodiments, the image acquisition device 1F according to the sixth embodiment differs in that it includes a measurement unit 10F instead of the measurement unit 10.

[0088] The measurement unit 10F includes a moving unit 14 in addition to the first detector 11 and the second detector 12. The moving unit 14 moves the positron-emitting nuclide 81 between the first detector 11 or the second detector 12 and the subject 90. The moving unit 14 may move the positron-emitting nuclide 81 continuously over time, or may move the positron-emitting nuclide 81 so as to sequentially place the positron-emitting nuclide 81 at spaced positions. The processing unit 20 constantly keeps track of the position of the positron-emitting nuclide 81 in order to determine the position at which gamma-ray photons have Compton scattered in the subject 90.

[0089] The movement direction of the positron-emitting nuclide 81 may be one or two directions parallel to the detection surface of the first detector 11 or the second detector 12, may be a direction perpendicular to the detection surface, or may be three directions including a direction parallel to the detection surface and a direction perpendicular to the detection surface. By moving the positron-emitting nuclide 81 in a direction parallel to the detection surface, the field of view of the device can be expanded or uniformed. By moving the positron-emitting nuclide 81 in a direction perpendicular to the detection surface, the image quality of the acquired first tomographic image can be improved.

[0090] 9A and 9B are diagrams illustrating changes in the field of view of the measurement unit 10F of the image acquisition device 1F of the sixth embodiment when the positron-emitting nuclide 81 is moved in a direction parallel to the detection surface of the first detector 11. In this figure, the field of view of the device is indicated by a hatched area. As shown in Fig. 9A, when the positron-emitting nuclide 81 is located near the center of the detection surface of the first detector 11, both ends of the subject 90 may fall outside the field of view.

[0091] 9(b), when the positron-emitting nuclide 81 is located on the first end side of the detection surface of the first detector 11 (to the left of the center in the figure), the first end side of the subject 90 is within the field of view, but the second end side of the subject 90 (to the right of the center in the figure) may be significantly outside the field of view. Conversely, when the positron-emitting nuclide 81 is located on the second end side of the detection surface of the first detector 11, the second end side of the subject 90 is within the field of view, but the first end side of the subject 90 may be significantly outside the field of view. In this way, by moving the positron-emitting nuclide 81 in a direction parallel to the detection surface of the first detector 11, the field of view of the device can be expanded or made uniform.

[0092] 10 is a diagram illustrating the image quality of the first tomographic image when the positron-emitting nuclide 81 is moved in a direction perpendicular to the detection surface of the first detector 11 in the measurement unit 10F of the image acquisition device 1F of the sixth embodiment. In this figure, the position of the positron-emitting nuclide 81 is denoted by P, and the detection position of the gamma ray by the first detector 11 is denoted by R. １ and the detection position of the gamma ray by the second detector 12 is R ２ The position where the gamma ray is Compton scattered is C. Position P and position R １ The error range of the line segment connecting positions A and B and the error range of the line segment connecting positions P and C are indicated by hatched areas.

[0093] The actual gamma ray detection position R by the first detector 11 １ Since there is an error due to the spatial resolution in １ An error occurs in the estimation of the line segment connecting positions P and C, and further an error occurs in the estimation of the line segment connecting positions P and C, and finally an error occurs in the estimation of position C.

[0094] 10(a), the longer the distance between position P and position C, the larger the estimation error of position C, and as shown in Fig. 10(b), the shorter the distance between position P and position C, the smaller the estimation error of position C. Therefore, in order to improve the image quality of the acquired first tomographic image, it is preferable to place the positron-emitting nuclide 81 at a position close to the subject 90.

[0095] Since the size and shape of the subject 90 vary, it is preferable to move the positron-emitting nuclide 81 in three directions so as to expand or uniform the field of view of the device according to the size and shape of the subject 90 and improve the image quality of the first tomographic image obtained.

[0096] (Simulation example)

[0097] Next, we will explain the conditions and results of a simulation performed to acquire the first tomographic image of the subject (an image showing the distribution of Compton scattering positions in the subject) described using Figures 1 and 2. Here, we used Geant4, which can simulate particle tracks in a material using the Monte Carlo method.

[0098] 11 is a diagram showing the configuration and arrangement of the measurement unit assumed in the simulation. The first detector 11 and the second detector 12 each have dimensions of 500×500×5 mm. ３ The first detector 11 and the second detector 12 were each a Cherenkov detector including a Cherenkov radiator having a size of 1 mm. The spatial resolution of gamma ray photon detection by each of the first detector 11 and the second detector 12 was 3 mm in full width at half maximum in the directions of the x, y, and z axes, and the resolution dL of the flight distance difference L was 7.5 mm in full width at half maximum. The first detector 11 and the second detector 12 were arranged parallel to each other and separated by a distance of 300 mm.

[0099] A cylindrical phantom with a diameter of 200 mm and a height of 100 mm was assumed, and this was used as the subject 90. The subject 90 was placed at a central position between the first detector 11 and the second detector 12 so that the central axis of the cylinder of the subject 90 was perpendicular to the detection surfaces of the first detector 11 and the second detector 12. A positron-emitting nuclide 81 was placed at a central position between the first detector 11 and the subject 90. The size of the positron-emitting nuclide 81 was ignored.

[0100] 12 is a diagram showing the configuration of a phantom assumed as the subject 90 in the simulation. This diagram shows the phantom as the subject 90 having a cylindrical shape, viewed in the direction of the central axis (z direction). The subject 90 is assumed to have six types of cylindrical regions 91 to 96 extending in the z direction, and these regions are covered by a region 97.

[0101] Region 91 was a region made of bone with a diameter of 22 mm, and three regions were provided. Region 92 was a region made of bone with a diameter of 18 mm, and three regions were provided. Region 93 was a region made of air with a diameter of 14 mm, and six regions were provided. Region 94 was a region made of air with a diameter of 12 mm, and six regions were provided. Region 95 was a region made of air with a diameter of 10 mm, and six regions were provided. Region 96 was a region made of air with a diameter of 8 mm, and ten regions were provided. Region 97 was a region made of water with a diameter of 200 mm.

[0102] 13 to 16 are diagrams showing first tomographic images obtained by simulation. These diagrams are images of a cross section parallel to the xy plane of a phantom serving as a cylindrical subject 90. In each of these diagrams, the first tomographic image shown in (a) represents the frequency of Compton scattering occurrence with shades of gray, with the higher the frequency of Compton scattering occurrence being the lighter the color. Furthermore, the graph shown in (b) shows the distribution of the frequency of Compton scattering occurrence on the straight line shown in (a).

[0103] Fig. 13 is a first tomographic image obtained based on coincidence events in which the energy of gamma-ray photons after Compton scattering is in the range of 0 to 260 keV. Fig. 14 is a first tomographic image obtained based on coincidence events in which the energy of gamma-ray photons after Compton scattering is in the range of 260 to 370 keV. Fig. 15 is a first tomographic image obtained based on coincidence events in which the energy of gamma-ray photons after Compton scattering is in the range of 370 to 511 keV. Fig. 16 is a first tomographic image obtained based on coincidence events in which the energy of gamma-ray photons after Compton scattering is in the range of 0 to 511 keV.

[0104] 13 to 16, the resolution of the first tomographic image (FIG. 15) obtained based on coincidence events in which the energy of gamma-ray photons after Compton scattering is in the range of 370 to 511 keV is poor. In comparison, the resolution of the first tomographic image (FIG. 13) obtained based on coincidence events in which the energy of gamma-ray photons after Compton scattering is in the range of 0 to 260 keV is excellent, and the resolution of the first tomographic image (FIG. 14) obtained based on coincidence events in which the energy of gamma-ray photons after Compton scattering is in the range of 260 to 370 keV is even better.

[0105] In this way, by selectively determining the Compton scattering position C for events in which the energy of Compton scattered gamma ray photons is lower than the first threshold E1, a first tomographic image with excellent spatial resolution can be obtained.

[0106] 13 and 14, in the first tomographic image (FIG. 13) obtained based on coincidence events in which the energy of gamma-ray photons after Compton scattering is in the range of 0 to 260 keV, the frequency of Compton scattering occurrence is low in the region near the center of the subject 90, and is high in the region around the subject 90. This is thought to be due to the following phenomenon.

[0107] When one gamma ray photon of a gamma ray photon pair generated by the positron-emitting nuclide 81 is Compton scattered at a certain position in the central region of the subject 90, this means that the other gamma ray photon is incident on the first detector 11 at a substantially normal angle. In order for the energy of the gamma ray photon Compton scattered at a certain position in the central region of the subject 90 to be 260 keV or less, the Compton scattering angle θ is required to be 90° or more, and the scattered gamma ray photon will not be incident on the second detector 12. For this reason, in Fig. 13, the frequency of Compton scattering occurrence is low in the central region of the subject 90.

[0108] 13, some Compton scattering events are observed in the central region of the subject 90, which is thought to be due to the detection of events in which gamma ray photons that had energy of 260 keV or more after forward scattering in the central region of the subject 90 are further Compton scattered in the second detector 12 and then have energy of 260 keV or less. The same factor is also thought to be responsible for the increased frequency of Compton scattering in the peripheral region of the subject 90 in FIG.

[0109] Since it is not possible to determine whether a gamma-ray photon has been photoelectrically absorbed or Compton scattered by the second detector 12, it is preferable to exclude events in which the energy of the gamma-ray photon is equal to or less than a certain threshold value. That is, it is preferable that the processing unit 20 selectively determines the position C where the gamma-ray photon has Compton scattered within the subject when the gamma-ray photon energy represented by the signal output from the second detector 12 is lower than a first threshold value E1 and higher than a second threshold value E2.

[0110] By doing so, the spatial resolution for detecting position C can be further reduced, and a first tomographic image with even better spatial resolution can be acquired. The second threshold E2 can be set to, for example, 150 keV, 200 keV, 230 keV, 260 keV, 300 keV, 330 keV, etc., provided that the second threshold E2 is lower than the first threshold E1. It is also preferable that the second threshold E2 can be set arbitrarily.

[0111] In addition, by moving the positron-emitting nuclide 81 between the first detector 11 or the second detector 12 and the subject 90 as in the sixth embodiment (FIG. 8), the above problem occurring in the first tomographic image shown in FIG. 13 can be alleviated.

[0112] The image acquisition device and the image acquisition method are not limited to the above-described embodiments and configuration examples, and various modifications are possible. For example, the configurations of any two or more of the above-described embodiments may be combined.

[0113] The image acquisition device of the first aspect according to the above embodiment includes: (1) a measurement unit including a first detector and a second detector each for detecting gamma ray photons, and outputting signals representing the detection position, the detection time, and the gamma ray photon energy when the first detector and the second detector each detect a gamma ray photon; and (2) a processing unit for processing signals output from the first detector and the second detector, respectively; (3) in a first measurement mode, the measurement unit outputs signals representing the detection position, the detection time, and the gamma ray photon energy of the gamma ray photon by the first detector and the second detector, respectively, with a subject placed between the first detector and the second detector and a positron-emitting nuclide placed between the first detector or the second detector and the subject; and (4) in the first measurement mode, the processing unit outputs signals representing the detection position, the detection time, and the gamma ray photon energy of the gamma ray photon by the first detector and the second detector, respectively, with a subject placed between the first detector and the second detector and a positron-emitting nuclide placed between the first detector or the second detector and the subject; For each coincidence event in which the first detector and the second detector simultaneously count a pair of generated gamma ray photons, when the gamma ray photon that arrives at one of the first detector and the second detector arrives without undergoing Compton scattering within the subject, and the gamma ray photon that arrives at the other detector arrives after undergoing Compton scattering within the subject, and the gamma ray photon energy represented by the signal output from the other detector is even lower than a first threshold value that is lower than 511 keV, the position where the gamma ray photon has Compton scattered within the subject is selectively determined based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector and the position of the positron-emitting nuclide, and a first tomographic image is created that represents the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence events.

[0114] In the image acquisition device of the second aspect, in the configuration of the first aspect, the measurement unit may be configured to, in the second measurement mode, output signals representing the detection position, detection time, and gamma-ray photon energy of gamma-ray photons by each of the first and second detectors, with a subject administered a drug labeled with a positron-emitting nuclide being placed between the first and second detectors, and the processing unit, in the second measurement mode, determine the position of the annihilation event based on the detection position and detection time of the gamma-ray photon by each of the first and second detectors for each coincidence event in which the first and second detectors simultaneously count a pair of gamma-ray photons generated by an electron-positron annihilation event in the positron-emitting nuclide, create a second tomographic image representing the distribution of the positions of annihilation event occurrence in the subject determined for each of the multiple coincidence events, and correct the second tomographic image based on the first tomographic image.

[0115] The image acquisition device of the third aspect according to the above embodiment comprises: (1) a measurement unit including a first detector and a second detector each for detecting gamma ray photons, and outputting a signal representing the detection position, the detection time, and the gamma ray photon energy when the first detector and the second detector each detect a gamma ray photon; and (2) a processing unit for processing the signals output from the first detector and the second detector, respectively; and (3) the measurement unit is configured such that a subject who has been administered a drug labeled with a positron-emitting nuclide is placed between the first detector and the second detector, and the first detector or the second detector detects a gamma ray photon. With a positron-emitting nuclide placed between the first and second detectors and the subject, the first and second detectors output signals representing the detection position, detection time, and gamma-ray photon energy of the gamma-ray photon, respectively; and (4) a processing unit, for each coincidence event in which the first and second detectors simultaneously count a pair of gamma-ray photons generated by an electron-positron annihilation event in the positron-emitting nuclide, determines whether (a) the gamma-ray photon arriving at one of the first and second detectors arrived without Compton scattering within the subject, and (b) when a gamma ray photon arriving at the first detector arrives after Compton scattering within the subject and the gamma ray photon energy represented by a signal output from the other detector is lower than a first threshold value lower than 511 keV, the position where the gamma ray photon has been Compton scattered within the subject is determined selectively based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector, and the position of a positron-emitting nuclide placed between the first detector or the second detector and the subject; When the arriving gamma ray photons arrive without undergoing Compton scattering within the subject, the position where the pair annihilation event occurred is determined based on the detection positions and detection times of the gamma ray photons by the first detector and the second detector, respectively; (c) a first tomographic image is created that represents the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence counting events; a second tomographic image is created that represents the distribution of the positions where pair annihilation events occurred in the subject determined for each of the multiple coincidence counting events; and the second tomographic image is corrected based on the first tomographic image.

[0116] In the image acquisition device of the fourth aspect, in the configuration of any of the first to third aspects, the processing unit may be configured to selectively determine the position where the gamma ray photon has Compton scattered within the subject when the gamma ray photon energy represented by the signal output from the other detector is lower than a first threshold value and higher than a second threshold value.

[0117] In the image acquisition device of the fifth aspect, in the configuration of the second or third aspect, the positron-emitting nuclide placed between the first detector or the second detector and the subject and the positron-emitting nuclide that labels the drug administered to the subject may be of the same type.

[0118] In the image acquisition device of the sixth aspect, in the configuration of any of the first to fifth aspects, the processing unit may be configured to determine whether or not a gamma ray photon arriving at the first detector or the second detector has undergone Compton scattering based on any one or more of the position of the positron-emitting nuclide, the magnitude of the energy of the gamma ray photon, and the detection time of the gamma ray photon by each of the first detector and the second detector.

[0119] In the image acquisition device of the seventh aspect, in the configuration of any of the first to sixth aspects, the measurement unit may be configured to output signals representing the detection position, detection time, and gamma ray photon energy of gamma ray photons by each of the first detector and the second detector, with a positron-emitting nuclide placed between the first detector and the subject, and a positron-emitting nuclide also placed between the second detector and the subject.

[0120] In the image acquisition device of the eighth aspect, in the configuration of any of the first to seventh aspects, the measurement unit may be configured so that the detection surface of the first detector is narrower than the detection surface of the second detector, and a positron-emitting nuclide is placed between the first detector and the subject, and output signals representing the detection position, detection time, and gamma-ray photon energy of the gamma-ray photon by each of the first detector and the second detector.

[0121] In the image acquisition device of the ninth aspect, in the configuration of any of the first to eighth aspects, the measurement unit may further include a shield that prevents gamma ray photons backscattered from either the first detector or the second detector from entering the other detector.

[0122] In the image acquisition device of the tenth aspect, in the configuration of any of the first to ninth aspects, the measurement unit may further include a moving unit that moves the positron-emitting nuclide between the first detector or the second detector and the subject.

[0123] The image acquisition method of the first aspect according to the above embodiment includes: (1) a measurement step using a first detector and a second detector, each of which detects gamma-ray photons, and outputting signals representing the detection position, detection time, and gamma-ray photon energy when the first detector and the second detector each detect a gamma-ray photon; and (2) a processing step processing the signals output from the first detector and the second detector, respectively; (3) the measurement step, in a first measurement mode, outputs signals representing the detection position, detection time, and gamma-ray photon energy of the gamma-ray photon by the first detector and the second detector, respectively, with a subject placed between the first detector and the second detector and a positron-emitting nuclide placed between the first detector or the second detector and the subject; and (4) the processing step, in the first measurement mode, outputs signals representing the detection position, detection time, and gamma-ray photon energy of the gamma-ray photon by the first detector and the second detector, respectively, with a subject placed between the first detector or the second detector and the subject; For each coincidence event in which the first detector and the second detector simultaneously count a pair of gamma ray photons generated by an annihilation event, when the gamma ray photon arriving at one of the first and second detectors arrives without undergoing Compton scattering within the subject, and the gamma ray photon arriving at the other detector arrives after Compton scattering within the subject, and the gamma ray photon energy represented by the signal output from the other detector is lower than a first threshold value lower than 511 keV, the positions where the gamma ray photons have been Compton scattered within the subject are selectively determined based on the detection positions and detection times of the gamma ray photons by the first and second detectors, respectively, and the positions of the positron-emitting nuclides, and a first tomographic image is created that represents the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence events.

[0124] In the image acquisition method of the second aspect, in the configuration of the first aspect, the measurement step may be configured to, in a second measurement mode, output signals representing the detection position, detection time, and gamma-ray photon energy of a gamma-ray photon by each of the first and second detectors, with a subject administered a drug labeled with a positron-emitting nuclide being placed between the first and second detectors; and the processing step, in the second measurement mode, determine, for each coincidence event in which the first and second detectors simultaneously count a pair of gamma-ray photons generated by an electron-positron annihilation event in the positron-emitting nuclide, the position of the annihilation event based on the detection position and detection time of the gamma-ray photon by each of the first and second detectors; create a second tomographic image representing the distribution of the positions of annihilation event occurrence in the subject determined for each of the multiple coincidence events; and correct the second tomographic image based on the first tomographic image.

[0125] The image acquisition method of the third aspect according to the above embodiment includes: (1) a measuring step using a first detector and a second detector, each of which detects a gamma ray photon, and outputting a signal representing the detection position, the detection time, and the gamma ray photon energy when the first detector and the second detector each detect a gamma ray photon; and (2) a processing step of processing the signals output from the first detector and the second detector, and (3) the measuring step includes placing a subject who has been administered a drug labeled with a positron-emitting nuclide between the first detector and the second detector, and and (4) a processing step outputs signals representing the detection position, detection time, and gamma-ray photon energy of the gamma-ray photon by each of the first and second detectors, with a positron-emitting nuclide placed between the first or second detector and the subject, and for each coincidence event in which the first and second detectors simultaneously count a pair of gamma-ray photons generated by an annihilation event of an electron and a positron in the positron-emitting nuclide, (a) a gamma-ray photon that arrived at one of the first and second detectors without Compton scattering within the subject is counted. and when the gamma ray photon arriving at the other detector arrives after Compton scattering within the subject and the gamma ray photon energy represented by the signal output from the other detector is lower than a first threshold value lower than 511 keV, selectively determine the position where the gamma ray photon has been Compton scattered within the subject based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector, and the position of a positron-emitting nuclide placed between the first detector or the second detector and the subject; and (b) determine the position where the gamma ray photon has been Compton scattered within the subject based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector, If the gamma ray photons arriving at both detectors arrive without undergoing Compton scattering within the subject, the position where the annihilation event occurred is determined based on the detection positions and detection times of the gamma ray photons by the first detector and the second detector, respectively; (c) a first tomographic image is created that represents the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence counting events; a second tomographic image is created that represents the distribution of the positions where the annihilation events occurred in the subject determined for each of the multiple coincidence counting events; and the second tomographic image is corrected based on the first tomographic image.

[0126] In the image acquisition method of the fourth aspect, in the configuration of any of the first to third aspects, the processing step may be configured to selectively determine the position where the gamma ray photon has Compton scattered within the subject when the gamma ray photon energy represented by the signal output from the other detector is lower than a first threshold value and higher than a second threshold value.

[0127] In the image acquisition method of the fifth aspect, in the configuration of the second or third aspect, the positron-emitting nuclide placed between the first detector or the second detector and the subject and the positron-emitting nuclide that labels the drug administered to the subject may be of the same type.

[0128] In the image acquisition method of the sixth aspect, in the configuration of any of the first to fifth aspects, the processing step may be configured to determine whether or not a gamma ray photon arriving at the first detector or the second detector has undergone Compton scattering based on any one or more of the position of the positron-emitting nuclide, the magnitude of the energy of the gamma ray photon, and the detection time of the gamma ray photon by each of the first detector and the second detector.

[0129] In the image acquisition method of the seventh aspect, in the configuration of any of the first to sixth aspects, the measurement step may be configured to output signals representing the detection position, detection time and gamma-ray photon energy by each of the first detector and the second detector, with a positron-emitting nuclide placed between the first detector and the subject, and a positron-emitting nuclide also placed between the second detector and the subject.

[0130] In the image acquisition method of the eighth aspect, in the configuration of any of the first to seventh aspects, the measurement step may be configured to use a first detector having a narrower detection surface than the second detector, and to output signals representing the detection position, detection time, and gamma-ray photon energy of gamma-ray photons by each of the first detector and the second detector, with a positron-emitting nuclide placed between the first detector and the subject.

[0131] In the image acquisition method of the ninth aspect, in the configuration of any of the first to eighth aspects, the measurement step may be configured to use a shield to prevent gamma ray photons backscattered from either the first detector or the second detector from entering the other detector.

[0132] In the image acquisition method of the tenth aspect, in the configuration of any one of the first to ninth aspects, the measurement step may be configured to move a positron-emitting nuclide between the first detector or the second detector and the subject.

[0133] The present invention can be used as an image acquisition device and an image acquisition method that can acquire tomographic images that represent anatomical information of a subject without performing image reconstruction processing.

[0134] 1A to 1F...image acquisition device, 10, 10D, 10E, 10F...measurement unit, 11, 11D...first detector, 12...second detector, 13...shield, 14...movement unit, 20...processing unit, 30...display unit, 81, 82, 83...positron-emitting nuclide, 90...subject.

Claims

1. A measurement unit including a first detector and a second detector each for detecting a gamma ray photon, and outputting signals representing a detection position, a detection time and a gamma ray photon energy when the first detector and the second detector each detect a gamma ray photon; and a processing unit processing signals output from the first detector and the second detector, respectively; wherein in a first measurement mode, the measurement unit outputs signals representing a detection position, a detection time and a gamma ray photon energy of the gamma ray photon by the first detector and the second detector, respectively, with a subject placed between the first detector and the second detector and a positron emitting nuclide placed between the first detector or the second detector and the subject; an electron-positron annihilation event in a positron-emitting nuclide, the processing unit selectively determines a position where the gamma ray photon has been Compton scattered in the subject based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector and the position of the positron-emitting nuclide, when the gamma ray photon that has arrived at one of the first detector and the second detector has arrived without being Compton scattered in the subject and the gamma ray photon that has arrived at the other detector has arrived after Compton scattering in the subject, and the gamma ray photon energy represented by the signal output from the other detector is lower than a first threshold value that is lower than 511 keV, and creates a first tomographic image representing a distribution of Compton scattering positions in the subject determined for each of a plurality of coincidence events.

2. The image acquisition device of claim 1, wherein in the second measurement mode, the measurement unit outputs signals representing the detection position, detection time and gamma ray photon energy of the gamma ray photon by each of the first detector and the second detector while the subject, who has been administered a drug labeled with a positron-emitting nuclide, is placed between the first detector and the second detector; and in the second measurement mode, the processing unit determines, for each coincidence event in which the first detector and the second detector simultaneously count a pair of gamma ray photons generated by an electron-positron annihilation event in a positron-emitting nuclide, the position at which the annihilation event occurred based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector, creates a second tomographic image representing a distribution of the positions at which the annihilation events occurred in the subject determined for each of a plurality of coincidence events, and corrects the second tomographic image based on the first tomographic image.

3. A measuring unit including a first detector and a second detector each for detecting gamma ray photons, and outputting signals representing the detection position, detection time and gamma ray photon energy when the first detector and the second detector each detect a gamma ray photon; and a processing unit processing the signals output from the first detector and the second detector, respectively, wherein the measuring unit outputs signals representing the detection position, detection time and gamma ray photon energy by the first detector and the second detector, respectively, when a subject administered with a drug labeled with a positron emitting nuclide is placed between the first detector and the second detector and a positron emitting nuclide is placed between the first detector or the second detector and the subject, and the processing unit performs the following for each coincidence event in which the first detector and the second detector simultaneously count a pair of gamma ray photons generated by an electron-positron annihilation event in a positron emitting nuclide: selectively, when a gamma ray photon arriving at one of the first detector and the second detector has arrived without Compton scattering within the subject, and a gamma ray photon arriving at the other detector has arrived after Compton scattering within the subject, and the gamma ray photon energy represented by a signal output from the other detector is lower than a first threshold value lower than 511 keV, determining a position where the gamma ray photon has been Compton scattered within the subject based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector, and the position of a positron-emitting nuclide placed between the first detector or the second detector and the subject; when a gamma ray photon arriving at both the first detector and the second detector has arrived without Compton scattering within the subject, determining a position where the annihilation event has occurred based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector; creating a first tomographic image representing a distribution of Compton scattering positions in the subject obtained for each of a plurality of coincidence counting events, creating a second tomographic image representing a distribution of annihilation event occurrence positions in the subject obtained for each of a plurality of coincidence counting events, and correcting the second tomographic image based on the first tomographic image.

4. An image acquisition device described in any one of claims 1 to 3, wherein the processing unit selectively determines the position at which the gamma ray photon has Compton scattered within the subject when the gamma ray photon energy represented by the signal output from the other detector is lower than the first threshold value and higher than the second threshold value.

5. An image acquisition device as described in claim 2 or 3, wherein the positron-emitting nuclide placed between the first detector or the second detector and the subject and the positron-emitting nuclide that labels a drug administered to the subject are of the same type.

6. An image acquisition device as described in any one of claims 1 to 5, wherein the processing unit determines whether or not a gamma ray photon arriving at the first detector or the second detector has undergone Compton scattering based on any one or more of the position of a positron-emitting nuclide, the magnitude of energy of the gamma ray photon, and the detection time of the gamma ray photon by each of the first detector and the second detector.

7. An image acquisition device as claimed in any one of claims 1 to 6, wherein the measurement unit outputs signals representing the detection position, detection time and gamma ray photon energy by each of the first detector and the second detector, with a positron emitting nuclide placed between the first detector and the subject and a positron emitting nuclide also placed between the second detector and the subject.

8. An image acquisition device according to any one of claims 1 to 7, wherein the measurement unit outputs signals representing the detection position, detection time and gamma ray photon energy by each of the first detector and the second detector, when the detection surface of the first detector is narrower than that of the second detector and a positron-emitting nuclide is placed between the first detector and the subject.

9. An image acquisition device according to any one of claims 1 to 8, wherein the measurement unit further includes a shield that prevents gamma ray photons backscattered from either the first detector or the second detector from entering the other detector.

10. An image acquisition device according to any one of claims 1 to 9, wherein the measurement unit further includes a moving unit that moves a positron-emitting nuclide between the first detector or the second detector and the subject.

11. A method for detecting gamma ray photons comprising: a measurement step of using a first detector and a second detector, each of which detects gamma ray photons, and outputting signals representing the detection position, detection time and gamma ray photon energy when the first detector and the second detector each detect a gamma ray photon; and a processing step of processing the signals output from the first detector and the second detector, wherein in a first measurement mode, the measurement step outputs signals representing the detection position, detection time and gamma ray photon energy of the gamma ray photon by the first detector and the second detector, respectively, with a test subject placed between the first detector and the second detector and a positron-emitting nuclide placed between the first detector or the second detector and the test subject; the processing step, in the first measurement mode, for each coincidence event in which the first detector and the second detector simultaneously count a pair of gamma ray photons generated by an electron-positron annihilation event in a positron-emitting nuclide, when a gamma ray photon arriving at one of the first detector and the second detector arrives without Compton scattering within the subject and a gamma ray photon arriving at the other detector arrives after Compton scattering within the subject, and a gamma ray photon energy represented by a signal output from the other detector is lower than a first threshold value lower than 511 keV, selectively determines a position where the gamma ray photon has been Compton scattered within the subject based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector and the position of the positron-emitting nuclide, and creates a first tomographic image representing a distribution of Compton scattering positions in the subject determined for each of a plurality of coincidence events.

12. The image acquisition method according to claim 11, wherein the measuring step, in a second measurement mode, outputs signals representing the detection position, detection time and gamma ray photon energy of a gamma ray photon by each of the first detector and the second detector while the subject, who has been administered a drug labeled with a positron-emitting nuclide, is placed between the first detector and the second detector; and the processing step, in the second measurement mode, determines, for each coincidence event in which the first detector and the second detector simultaneously count a pair of gamma ray photons generated by an electron-positron annihilation event in a positron-emitting nuclide, the position at which the annihilation event occurred based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector, creates a second tomographic image representing a distribution of the positions at which the annihilation events occurred in the subject determined for each of a plurality of coincidence events, and corrects the second tomographic image based on the first tomographic image. 13.The present invention comprises a measuring step of using a first detector and a second detector, each of which detects gamma ray photons, to output signals representing a detection position, a detection time and a gamma ray photon energy when the first detector and the second detector each detect a gamma ray photon, and a processing step of processing the signals output from the first detector and the second detector, respectively, wherein the measuring step outputs signals representing a detection position, a detection time and a gamma ray photon energy by the first detector and the second detector, respectively, in a state in which a subject administered with a drug labeled with a positron emitting nuclide is placed between the first detector and the second detector and a positron emitting nuclide is placed between the first detector or the second detector and the subject, and the processing step comprises: for each coincidence event in which the first detector and the second detector simultaneously count a pair of gamma ray photons generated by an electron-positron annihilation event in a positron emitting nuclide, selectively, when a gamma ray photon arriving at one of the first detector and the second detector has arrived without Compton scattering within the subject, and a gamma ray photon arriving at the other detector has arrived after Compton scattering within the subject, and the gamma ray photon energy represented by a signal output from the other detector is lower than a first threshold value lower than 511 keV, determining a position where the gamma ray photon has been Compton scattered within the subject based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector, and the position of a positron-emitting nuclide placed between the first detector or the second detector and the subject; when a gamma ray photon arriving at both the first detector and the second detector has arrived without Compton scattering within the subject, determining a position where the annihilation event has occurred based on the detection position and detection time of the gamma ray photon by each of the first detector and the second detector; An image acquisition method comprising: creating a first tomographic image representing a distribution of Compton scattering positions in the subject obtained for each of a plurality of coincidence counting events; creating a second tomographic image representing a distribution of annihilation event occurrence positions in the subject obtained for each of a plurality of coincidence counting events; and correcting the second tomographic image based on the first tomographic image.

14. An image acquisition method according to any one of claims 11 to 13, wherein the processing step selectively determines the position at which the gamma ray photon has been Compton scattered within the subject when the gamma ray photon energy represented by the signal output from the other detector is lower than the first threshold value and higher than the second threshold value.

15. An image acquisition method as described in claim 12 or 13, wherein the positron-emitting nuclide placed between the first detector or the second detector and the subject and the positron-emitting nuclide that labels a drug administered to the subject are of the same type as each other.

16. An image acquisition method according to any one of claims 11 to 15, wherein the processing step determines whether or not a gamma ray photon arriving at the first detector or the second detector has undergone Compton scattering based on any one or more of the position of a positron-emitting nuclide, the magnitude of energy of the gamma ray photon, and the detection time of the gamma ray photon by each of the first detector and the second detector.

17. An image acquisition method according to any one of claims 11 to 16, wherein the measurement step includes outputting signals representing the detection position, detection time and gamma ray photon energy by each of the first detector and the second detector, with a positron-emitting nuclide placed between the first detector and the subject and a positron-emitting nuclide also placed between the second detector and the subject.

18. An image acquisition method according to any one of claims 11 to 17, wherein the measurement step uses the first detector having a narrower detection surface than the second detector, and outputs signals representing the detection position, detection time and gamma ray photon energy by each of the first detector and the second detector with a positron-emitting nuclide placed between the first detector and the subject.

19. An image acquisition method according to any one of claims 11 to 18, wherein the measurement step uses a shield to prevent gamma ray photons backscattered from either the first detector or the second detector from entering the other detector.

20. An image acquisition method according to any one of claims 11 to 19, wherein the measuring step comprises transferring a positron-emitting nuclide between the first detector or the second detector and the subject.

Images