Image acquisition device, and image acquisition method

Abstract

To provide an image acquisition device and image acquisition method that can acquire a tomographic image representing anatomical information on a subject without processing image reconstruction.SOLUTION: An image acquisition device 1A comprises: a measurement unit 10; and a processing unit 20. The processing unit 20 is configured to, when a gamma-ray photon arriving at a second detector 12 is the gamma-ray photon arriving after Compton scattering within an analyte 90 for each coincidence counting event in which a first detector 11 and second detector 12 perform a coincidence counting of a pair of gamma-ray photons to be generated by a pair annihilation event of an electron and positron in a positron-emitting radioisotope 81, and gamma-ray photon energy a signal output from the second detector 12 represents is further lower than a first threshold lower than 511keV, selectively obtain a position in which the gamma-ray photon performs the Compton scattering within the analyte 90 on the basis of a detection position and detection time of the gamma-ray photon by the first detector 11 and second detector 12, and a position of the positron-emitting radioisotope 81.SELECTED DRAWING: Figure 1

Description

[Technical field]

[0001] The present invention relates to an image acquisition device and an image acquisition method. [Background technology]

[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 (distribution of the drug) in the subject, and can be used to diagnose the health condition of the subject.

[0003] X-ray CT scanners can also obtain three-dimensional tomographic images of a subject. The tomographic images obtained by X-ray CT scanners represent anatomical information of the subject. Hereinafter, the tomographic images will be referred to as three-dimensional tomographic images.

[0004] In order to improve the quality of PET images, a PET scanner and an X-ray CT scanner are used in combination to correct the PET images using anatomical information acquired by the X-ray CT scanner. However, since X-ray CT scanners are 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 many detectors are arranged around a measurement space in which a subject is placed. This device uses a detector having an LSO (Lu2SiO5:Ce) scintillator to detect gamma rays with energy of 307 keV or 202 keV emitted from 176Lu contained in the LSO scintillator of each detector, and detects gamma rays that have passed through the subject using another detector. Then, this device performs image reconstruction processing based on the detection results of gamma rays with energy of 307 keV or 202 keV to obtain a tomographic image showing 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 a gamma ray has come from a position on a cone surface called a Compton cone, based on the energy information of the scatterer and absorber. In contrast, an ETCC is said to be able to uniquely identify the direction of a gamma ray by tracking the recoil electron's track using a gas detector. This device detects the direction of a gamma ray that has been Compton scattered by the subject and has a reduced energy, among the gamma rays generated by a subject administered with a drug labeled with a positron-emitting nuclide, by the ETCC. Then, this device obtains a tomographic image showing the anatomical information of the subject by using an analytical method or a statistical method, i.e., by performing image reconstruction processing, based on the detection result of the direction of the gamma ray by the ETCC. [Prior art documents] [Patent documents]

[0008] [Patent Document 1] Patent No. 6990412 [Non-patent literature]

[0009] [Non-Patent Document 1] Mohammadreza Teimoorisichani etal., Med. Phys., 2022, vol 49, pages 309-323 [Non-Patent Document 2] Gerard Arino-Estrada, et al.,Phys. Med. Biol., 64 (2019) 175001 Summary of the Invention [Problem to be solved by the invention]

[0010] In both of 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 obtain a tomographic image showing anatomical information of a subject. The tomographic image obtained by the image reconstruction processing has a deteriorated image quality due to the image reconstruction processing, and the anatomical information is deteriorated.

[0011] The present invention has been made to solve the above-mentioned problems, and aims to provide an image acquisition device and an image acquisition method that can acquire tomographic images representing anatomical information of a subject without performing image reconstruction processing. [Means for solving the problem]

[0012] The image acquisition device of the present invention includes (1) a measurement unit including a first detector and a second detector each for detecting a gamma ray photon, and outputs 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 unit for processing the signals output from the first detector and the second detector. The image acquisition device of the present invention can be in the following forms.

[0013] In a first aspect of the image acquisition device, the measurement unit and the processing unit are as follows: In a first measurement mode, the measurement unit 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, 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. 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 the gamma ray photon arriving at one of the first detector and the second detector arrives without Compton scattering and the gamma ray photon arriving at the other detector arrives after Compton scattering in 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 lower than 511 keV, the processing unit selectively determines the position where 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 and the second detector and the position of the positron-emitting nuclide, and creates a first tomographic image showing the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence events.

[0014] In a second aspect of the image acquisition device, in addition to the first aspect, the measurement unit and the processing unit are as follows. In a 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 and second detectors, with a subject administered with a drug labeled with a positron-emitting nuclide being placed between the first and second detectors. In the second measurement mode, the processing unit determines 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 annihilation event of an electron and a positron in a positron-emitting nuclide, and creates a second tomographic image representing the distribution of the annihilation event positions in the subject determined for each of the multiple coincidence events. Then, the processing unit corrects the second tomographic image based on the first tomographic image.

[0015] In a third aspect of the image acquisition device, the measurement unit and the processing unit are as follows: 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, with a subject administered with a drug labeled with a positron-emitting nuclide being placed between the first detector and the second detector, and with a positron-emitting nuclide being placed between the first detector or the second detector and the subject. 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 a positron-emitting nuclide, (a) 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 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, 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 placed between the first detector or the second detector and the subject; and (b). When gamma ray photons arriving at both the first detector and the second detector arrive without Compton scattering, a position where a pair annihilation event occurs is determined based on the detection positions and detection times of the gamma ray photons by the first detector and the second detector, respectively. Then, the processing unit creates a first tomographic image representing the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence counting events, creates a second tomographic image representing the distribution of the pair annihilation event occurrence positions in the subject determined for each of the multiple coincidence counting events, and corrects the second tomographic image based on the first tomographic image.

[0016] In a fourth aspect of the image acquisition device, in addition to any of the first to third aspects, the processing unit 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 a first threshold value and higher than a second threshold value.

[0017] In a fifth aspect of the image acquisition device, in addition to the second or third aspect, the positron-emitting nuclide placed between the first or second detector and the subject and the positron-emitting nuclide that labels the drug administered to the subject are of the same type as each other.

[0018] In a sixth aspect of the image acquisition device, in addition to any of the first to fifth aspects, 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 the 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.

[0019] In a seventh aspect of the image acquisition device, in addition to any of the first to sixth aspects, 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 also a positron-emitting nuclide placed between the second detector and the subject.

[0020] In an eighth aspect of the image acquisition device, in addition to any of the first to seventh aspects, 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 the detection surface of the first detector being narrower than the detection surface of the second detector and a positron-emitting nuclide being placed between the first detector and the subject.

[0021] In a ninth aspect of the image acquisition device, in addition to any of the first to eighth aspects, 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.

[0022] In a tenth aspect of the image acquisition device, in addition to any one of the first to ninth aspects, the measurement section further includes a moving section that moves the positron-emitting nuclide between the first detector or the second detector and the subject.

[0023] The image acquisition method of the present invention includes: (1) a measurement step of using a first detector and a second detector, each of which detects a gamma-ray photon, 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 step of processing the signals output from the first detector and the second detector. The image acquisition method of the present invention can be implemented in the following manner.

[0024] In a first embodiment of the image acquisition method, the measuring step and the processing step are as follows: In a first measurement mode, the measuring step 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, 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. 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 the gamma ray photon arriving at one of the first and second detectors arrives without 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 processing step selectively determines 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 the positron-emitting nuclide, and creates a first tomographic image showing the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence events.

[0025] In a second aspect of the image acquisition method, in addition to the first aspect, the measurement step and the processing step are as follows. In the measurement step, in a second measurement mode, a subject administered with a drug labeled with a positron-emitting nuclide is placed between the first and second detectors, and signals representing the detection position, detection time, and gamma-ray photon energy of the gamma-ray photon are outputted by each of the first and second detectors. In the processing step, in the second measurement mode, for each coincidence event in which the first and second detectors simultaneously count a pair of gamma-ray photons generated in the subject by an annihilation event of an electron and a positron in the positron-emitting nuclide, the position at which the annihilation event occurred is determined based on the detection position and detection time of the gamma-ray photon by the first and second detectors, respectively, and a second tomographic image is created that represents the distribution of the annihilation event positions in the subject determined for each of the multiple coincidence events. Then, in the processing step, the second tomographic image is corrected based on the first tomographic image.

[0026] In a third embodiment of the image acquisition method, the measuring step and the processing step are as follows: The measuring 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 subject administered with a drug labeled with a positron-emitting nuclide being placed between the first detector and the second detector, and with a positron-emitting nuclide being placed between the first detector or the second detector and the subject. The processing step includes, 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, (a) 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 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, selectively 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 the positron-emitting nuclide placed between the first detector or the second detector and the subject; and (b). When gamma ray photons arriving at both the first detector and the second detector arrive without Compton scattering within the subject, a position where a pair annihilation event has 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. Then, the processing step creates a first tomographic image showing the distribution of Compton scattering positions in the subject determined for each of the multiple coincidence counting events, creates a second tomographic image showing the distribution of the annihilation event occurrence positions in the subject determined for each of the multiple coincidence counting events, and corrects the second tomographic image based on the first tomographic image.

[0027] In a fourth aspect of the image acquisition method, in addition to any of the first to third aspects, 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 a first threshold value and higher than a second threshold value.

[0028] In a fifth aspect of the image acquisition method, in addition to the second or third aspect, the positron-emitting nuclide placed between the first or second detector and the subject and the positron-emitting nuclide that labels the drug administered to the subject are of the same type as each other.

[0029] In a sixth aspect of the image acquisition method, in addition to any of the first to fifth aspects, the processing step determines whether or not a gamma ray photon arriving at the first detector or the second detector has undergone Compton scattering within the subject based on any one or more of the position of the 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.

[0030] In a seventh aspect of the image acquisition method, in addition to any of the first to sixth aspects, 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 being placed between the first detector and the subject, and a positron-emitting nuclide also being placed between the second detector and the subject.

[0031] In an eighth aspect of the image acquisition method, in addition to any of the first to seventh aspects, the measurement step uses a first detector having a narrower detection surface than the second detector, and with a positron-emitting nuclide placed between the first detector and the subject, outputs signals representing the detection position, detection time and gamma ray photon energy by each of the first detector and the second detector.

[0032] In a ninth aspect of the image acquisition method, in addition to any of the first to eighth aspects, the measurement step includes using a shield to prevent gamma ray photons backscattered from either the first detector or the second detector from entering the other.

[0033] In a tenth aspect of the image acquiring method, in addition to any one of the first to ninth aspects, the measuring step transfers a positron-emitting nuclide between the first detector or the second detector and the subject. Effect of the Invention

[0034] According to the present invention, it is possible to obtain a tomographic image representing anatomical information of a subject without performing image reconstruction processing. [Brief description of the drawings]

[0035] [Figure 1] FIG. 1 is a diagram showing the configuration of an image acquisition device 1A of the first embodiment (particularly, a diagram for explaining acquisition of a first tomographic image in a first measurement mode). [Diagram 2] FIG. 2 is a diagram for explaining a method for determining the position where a gamma ray photon has been Compton scattered within a subject. [Diagram 3] FIG. 3 is a diagram showing the configuration of the image acquisition device 1A of the first embodiment (particularly, a diagram for explaining acquisition of a second tomographic image in the second measurement mode). [Figure 4] FIG. 4 is a diagram showing the configuration of an image acquisition device 1B according to the second embodiment. [Diagram 5] FIG. 5 is a diagram showing the configuration of an image acquisition device 1C according to the third embodiment. [Figure 6] FIG. 6 is a diagram showing the configuration of an image acquisition device 1D according to the fourth embodiment. [Figure 7] FIG. 7 is a diagram showing the configuration of an image acquisition device 1E according to the fifth embodiment. [Figure 8] FIG. 8 is a diagram showing the configuration of an image acquisition device 1F according to the sixth embodiment. [Figure 9] Figures 9(a) and 9(b) are figures explaining the change in the field of view of the device when a positron-emitting nuclide 81 is moved in a direction parallel to the detection surface of the first detector 11 in the measurement section 10F of the image acquisition device 1F of the sixth embodiment. [Figure 10] Figures 10(a) and 10(b) are figures explaining 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 section 10F of the image acquisition device 1F of the sixth embodiment. [Figure 11] FIG. 11 is a diagram showing the configuration and arrangement of the measurement unit assumed in the simulation. [Figure 12] FIG. 12 is a diagram showing the configuration of a phantom assumed as the subject 90 in the simulation. [Figure 13] FIG. 13 is a diagram showing a first tomographic image obtained in the simulation. [Figure 14] FIG. 14 is a diagram showing a first tomographic image obtained in the simulation. [Figure 15] FIG. 15 is a diagram showing a first tomographic image obtained in the simulation. [Figure 16] FIG. 16 is a diagram showing a first tomographic image obtained in the simulation. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0036] Hereinafter, the embodiments of the present invention will be described in detail with reference to the accompanying drawings. In the description of the drawings, the same elements are given the same reference numerals, and duplicated descriptions are omitted. The present invention is not limited to these examples, but is indicated by the claims, and is intended to include all modifications within the meaning and scope equivalent to the claims.

[0037] (First embodiment) FIG. 1 is a diagram showing the configuration of an image acquisition device 1A of the 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 arranged opposite to each other with a subject 90 in between. Each of the first detector 11 and the second detector 12 detects a gamma ray photon, and outputs a signal representing the detection position, the detection time, and the gamma ray photon energy when the gamma ray photon is detected. The processing unit 20 processes the signals output from each of 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 created by the processing unit 20. The processing unit 20 and the display unit 30 may be configured, for example, by a computer.

[0038] For example, a Cherenkov detector is used as each of the first detector 11 and the second detector 12. The Cherenkov detector includes a Cherenkov radiator (e.g., lead glass, lead fluoride PbF2, hafnium oxide HfO2, etc.) and a microchannel plate photomultiplier tube (MCP-PMT). 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.

[0039] The Cherenkov detector is made of BGO (Bi4Ge3O 12 ) may be used as a Cherenkov radiator. A slow scintillator interacts with gamma rays and first emits Cherenkov light, then emits scintillation light, so it can be used as a Cherenkov radiator and achieves high time resolution. This allows the construction of a cheaper detector than an LSO scintillator, etc.

[0040] Moreover, 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 is, for example, one that uses thallium bromide (TlBr) and is equipped with an electrode for collecting charges and a high time resolution photodetector. For example, it may be one described in Non-Patent Document 2. By using a semiconductor detector, the energy resolution is improved, and as a result, it is expected that the ability to remove scattered components is improved and image quality is improved.

[0041] 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 is equal to or better than that. Similarly, it is desirable that the time resolution required for each of the first detector 11 and the second detector 12 is 20 to 35 ps or less in terms of coincidence time resolution.

[0042] When the first detector 11 and the second detector 12 each include a Cherenkov radiator or a scintillator, it is preferable to output a signal indicating the position (detection position) and time (detection time) at which the 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 by 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 the perpendicular direction.

[0043] 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 approximately equal to or larger than the size of a human brain.

[0044] The image acquiring device 1A and the image acquiring method using the same acquire a tomographic image (first tomographic image) of the subject 90 in a first measurement mode. The image acquiring device 1A and the image acquiring method can also acquire a tomographic image (second tomographic image) of the subject 90 in a second measurement mode. 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.

[0045] Acquisition of a 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 in particular is a diagram explaining acquisition of a first tomographic image in the first measurement mode.

[0046] 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. 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.

[0047] 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 where 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 of the first detector 11 and the second detector 12 arrived after Compton scattering within the subject.

[0048] Then, the processing unit 20 creates a first tomographic image representing the distribution of Compton scattering positions in the subject 90 obtained for each of the multiple coincidence counting events. This first tomographic image represents anatomical information of the subject 90.

[0049] The processing unit 20 can determine whether or not the gamma ray photon arriving at the first detector 11 or the second detector 12 has undergone Compton scattering based on any 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. As shown in FIG. 1, if the positron-emitting nuclide 81 is placed between the first detector 11 and the subject 90, it can be determined that the gamma ray arriving at the first detector 11 has not undergone Compton scattering in the subject. The energy of a pair of gamma ray photons generated by annihilation of an electron and a positron is 511 keV, but the energy of the gamma ray is reduced by Compton scattering, so it can be determined whether or not the gamma ray has undergone Compton scattering based on the magnitude of the energy of the gamma ray. It can be determined whether or not the gamma ray has undergone Compton scattering in the subject based on the time relationship between the detection times of the gamma ray photon by each of the first detector 11 and the second detector 12.

[0050] FIG. 2 is a diagram for explaining a method for determining the position where a gamma ray photon has undergone Compton scattering within a subject. The position of the positron-emitting nuclide 81 is denoted as P, the position where the gamma ray is detected by the first detector 11 as R1, the position where the gamma ray is detected by the second detector 12 as R2, and the position where the gamma ray has undergone Compton scattering as C. The time when the gamma ray is detected by the first detector 11 as t1, and the time when the gamma ray is detected by the second detector 12 as t2. Of a pair of gamma ray photons generated by an electron-positron annihilation event in the positron-emitting nuclide 81, the flight distance of one gamma ray is the distance d1 from position P to position R1. The flight distance of the other gamma ray is the distance d from position P to position C. 21 and the distance d from position C to position R2 22 Sum of (d 21 +d 22 ).

[0051] The difference in flight distance of a pair of gamma ray photons L(=d 21 +d 22-d1) is equal to the difference (t2-t1) between the detection times of these gamma ray photons multiplied by the speed of light c. In addition, the line segment connecting position P and position R1 and the line segment connecting position P and position C are parallel to each other. From the above, it is possible to determine the position C where the gamma ray photon has been Compton scattered based on the detection position R1 and detection time t1 of the gamma ray photon by the first detector 11, the detection position R2 and detection time t2 of the gamma ray photon by the second detector 12, and the position P of the positron-emitting nuclide 81.

[0052] When neither of the pair of gamma ray photons undergoes Compton scattering, the Compton scattering position obtained 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.

[0053] 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 described below.

[0054] 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. The coordinate value of the position R1 is defined as (R 1x ,R 1y ,R 1z ), and the coordinate value of position R2 is (R 2x ,R 2y ,R 2z ), and the coordinate value of position C is (C x ,C y ,C z ) and the difference in the flight distance of a pair of gamma ray photons (d 21 +d 22 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 ,dR 2y ,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). ΔC y ,ΔC z is expressed by a similar formula. In the following, ΔC x The formula is shown only for .

[0055]

number

[0056] 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 position C, x It is desirable to have a small partial differential value. x is expressed by the following formula (2). x The partial differential of is expressed by the following formula (3): Here, the position P of the positron-emitting nuclide 81 is set as the origin of the xyz Cartesian coordinate system.

[0057]

number

[0058]

number

[0059] Rearranging equation (3), we obtain equation (4) below. Here, θ is the Compton scattering angle (see FIG. 2). Compton scattering angle θ is the angle of the direction in which gamma rays travel after scattering position C, relative to the direction in which they come from position P of the positron-emitting nuclide 81 to scattering position C.

[0060]

number

[0061] As can be seen from equation (4), the smaller the Compton scattering angle θ (the closer to 0), the smaller the C due to L. x The partial derivative of is large. As mentioned above, the C x The larger the partial differential of , the greater the spatial resolution of detecting 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 position C, x It is preferable that the partial differential of 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.

[0062] Therefore, in the first processing step, when the gamma ray photon energy represented by the signal output from the second detector 12 is lower than a first threshold E1 lower than 511 keV, the processing unit 20 selectively determines a position C where the gamma ray photon is Compton scattered in 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. In this way, the spatial resolution of the detection of the position C can be reduced, and a first tomographic image with excellent spatial resolution can be obtained. 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.

[0063] 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 in particular, a diagram explaining the acquisition of the second tomographic image in the second measurement mode.

[0064] 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.

[0065] 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 it easy to prepare for measurement. The positron-emitting nuclide 81 may be a calibration positron-emitting nuclide such as 68Ge / 68Ga.

[0066] In the second processing step, the processing unit 20 determines the position where the pair 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, 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. The position where the pair annihilation event occurred for each coincidence event can be determined based on the difference in the detection time of the gamma ray photon by each of the first detector 11 and the second detector 12 on a line segment connecting the detection positions of the gamma ray photon by each of the first detector 11 and the second detector 12.

[0067] Then, the processing unit 20 creates a second tomographic image showing the distribution of the positions of occurrence of annihilation events in the subject 90 obtained for each of the multiple coincidence counting events. Since 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 image reconstruction processing. This second tomographic image shows the distribution of the positron-emitting nuclide 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, so that the second tomographic image can be acquired after correction for the gamma-ray absorption distribution in the subject 90.

[0068] 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, so it is preferable to perform the acquisition of the first tomographic image in the first measurement mode first.

[0069] In this embodiment, a tomographic image (first tomographic image) showing anatomical information of the subject 90 can be obtained without performing image reconstruction processing, and therefore, deterioration of image quality caused by image reconstruction processing can be avoided, and deterioration of anatomical information can be avoided. Also, a tomographic image (second tomographic image) showing the distribution of positron-emitting nuclides (distribution of drug) in the subject can be obtained without performing image reconstruction processing.

[0070] In this embodiment, since a large-scale rotation mechanism such as that in an X-ray CT scanner is not required, a compact and inexpensive device can be obtained. Also, compared with the case where an X-ray CT scanner is used, this embodiment can reduce the amount of radiation exposure of the subject.

[0071] The device described in Non-Patent Document 1 has a PET device configuration in which many detectors are arranged around the measurement space in which the subject is placed, making it difficult to reduce its size, and also makes it difficult to reduce its cost 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 made smaller and less expensive.

[0072] 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.

[0073] Furthermore, since the Compton scattering position C is selectively determined for an event in which the energy of the Compton scattered gamma ray photon is lower than the first threshold, it is possible to obtain a first tomographic image with excellent spatial resolution.

[0074] Second embodiment 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 with 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 obtained during a common period.

[0075] In the 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. 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. 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 is required, making it easy to prepare for measurement. The positron-emitting nuclide 81 may be a calibration positron-emitting nuclide such as 68Ge / 68Ga. Annihilation events of electrons and positrons in the positron-emitting nuclide 81 and the positron-emitting nuclide 83 generate a pair of gamma ray photons flying 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.

[0076] 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.

[0077] 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 any one or more of the position of the positron-emitting nuclide 81, the magnitude of 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.

[0078] When the result of the above determination is 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, and the gamma ray photon energy represented by the signal output from the other detector is even lower than a first threshold E1 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 was 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.

[0079] On the other hand, if the result of the above judgment is that the gamma ray photons arriving 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 positions and detection times of the gamma ray photons by the first detector 11 and the second detector 12, respectively.

[0080] After performing the above-mentioned processing on the multiple coincidence events, the processing unit 20 creates a first tomographic image showing the distribution of Compton scattering positions in the subject 90, and creates a second tomographic image showing the distribution of annihilation event occurrence positions in the subject 90. The first tomographic image shows anatomical information of the subject 90. The second tomographic image shows the distribution of positron-emitting nuclides 83 (distribution of drugs) in the subject 90, and can be used to diagnose the health condition of the subject 90. The processing unit 20 can further correct the second tomographic image based on the first tomographic image, thereby acquiring the second tomographic image after correction for the gamma-ray absorption distribution in the subject 90.

[0081] 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 an annihilation event occurs in the subject 90 can be determined during a common period, thereby shortening the time required to restrain the subject 90.

[0082] Third embodiment 5 is a diagram showing the configuration of an image acquisition device 1C of the third embodiment. The image acquisition device 1C includes a measurement unit 10, a processing unit 20, and a display unit 30. Compared with 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 the same type as each other.

[0083] The processing unit 20 judges whether the 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 the positron-emitting nuclide 81 or the 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 time of the gamma ray photon by each of the first detector 11 and the second detector 12. The processing unit 20 determines the position where the gamma ray photon has been Compton scattered in the subject 90 based on the judgment result.

[0084] In this embodiment, the positron-emitting nuclides 81, 82 can be arranged symmetrically with respect to the subject 90, so that a first tomographic image of higher quality can be acquired. In addition, the number of Compton scattering events per unit time in the subject 90 increases, so that the measurement time can be shortened.

[0085] (Fourth embodiment) 6 is a diagram showing the configuration of an image acquisition device 1D of the fourth embodiment. The image acquisition device 1D includes a measurement unit 10D, a processing unit 20, and a display unit 30. Compared with the previous embodiments, the image acquisition device 1D of the fourth embodiment differs in that it includes a measurement unit 10D instead of the measurement unit 10.

[0086] 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. Each of the first detector 11D and the second detector 12 outputs a signal representing the detection position, the detection time, and the gamma-ray photon energy when a gamma-ray photon is detected.

[0087] 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 in the subject 90), so 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 constructed at low cost.

[0088] Fifth embodiment 7 is a diagram showing the configuration of an image acquisition device 1E of the fifth embodiment. The image acquisition device 1E includes a measurement unit 10E, a processing unit 20, and a display unit 30. Compared with the previous embodiments, the image acquisition device 1E of the fifth embodiment differs in that it includes a measurement unit 10E instead of the measurement unit 10.

[0089] 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. The shield 13 is disposed between the first detector 11 and the second detector 12 at a position that does not interfere with the measurement of the Compton scattering position in the subject 90 and the measurement of the position of the annihilation event occurrence in the subject 90. The shield 13 is preferably a plate-shaped member made of a high-density material (e.g., lead) capable of blocking gamma rays.

[0090] Sixth embodiment 8 is a diagram showing the configuration of an image acquisition device 1F of the sixth embodiment. The image acquisition device 1F includes a measurement unit 10F, a processing unit 20, and a display unit 30. Compared with the previous embodiments, the image acquisition device 1F of the sixth embodiment differs in that it includes a measurement unit 10F instead of the measurement unit 10.

[0091] 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 it 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 where the gamma-ray photons have been Compton scattered in the subject 90.

[0092] The direction of movement 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.

[0093] FIG. 9 is a diagram for explaining the change in the field of view of the apparatus when the 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. In this figure, the field of view of the apparatus is indicated by a hatched area. As shown in FIG. 9(a), 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 be out of the field of view. In contrast, as shown in FIG. 9(b), when the positron-emitting nuclide 81 is located on the first end side of the detection surface of the first detector 11 (the side to the left of the center in the figure), the first end side of the subject 90 may be included in the field of view, but the second end side of the subject 90 (the side to the right of the center in the figure) may be significantly out of 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 may be included in the field of view, but the first end side of the subject 90 may be significantly out of 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.

[0094] 10 is a diagram for explaining 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 diagram, the position of the positron-emitting nuclide 81 is P, the position of the gamma ray detected by the first detector 11 is R1, the position of the gamma ray detected by the second detector 12 is R2, and the position where the gamma ray is Compton scattered is C. The error range of the line segment connecting the position P and the position R1 and the error range of the line segment connecting the position P and the position C are indicated by hatched areas.

[0095] Since an error due to spatial resolution exists in the actual detection position R1 of gamma rays by the first detector 11, this causes an error in the estimation of the line segment connecting the position P and the position R1, and further causes an error in the estimation of the line segment connecting the position P and the position C, and finally causes an error in the estimation of the position C. As shown in Fig. 10(a), the longer the distance between the position P and the position C, the larger the estimation error of the position C, and as shown in Fig. 10(b), the shorter the distance between the position P and the position C, the smaller the estimation error of the 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.

[0096] 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 apparatus according to the size and shape of the subject 90 and to improve the image quality of the first tomographic image obtained.

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

[0098] FIG. 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. 3The first detector 11 and the second detector 12 were assumed to be 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 assumed to be 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 assumed to be 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 facing each other at a distance of 300 mm. A cylindrical phantom with a diameter of 200 mm and a height of 100 mm was assumed to be the subject 90. The subject 90 was placed at the center 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 the center between the first detector 11 and the subject 90. The size of the positron-emitting nuclide 81 was ignored.

[0099] FIG. 12 is a diagram showing the configuration of a phantom assumed as the subject 90 in the simulation. This figure shows a phantom as the subject 90 having a cylindrical shape, viewed in the central axis direction (z direction). The subject 90 has six types of cylindrical regions 91 to 96 extending in the z direction, and these are covered by a region 97. The region 91 is a region made of bone and has a diameter of 22 mm, and three regions are provided. The region 92 is a region made of bone and has a diameter of 18 mm, and three regions are provided. The region 93 is a region made of air and has a diameter of 14 mm, and six regions are provided. The region 94 is a region made of air and has a diameter of 12 mm, and six regions are provided. The region 95 is a region made of air and has a diameter of 10 mm, and six regions are provided. The region 96 is a region made of air and has a diameter of 8 mm, and ten regions are provided. The region 97 is a region made of water and has a diameter of 200 mm.

[0100] 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 as a subject 90 having a cylindrical shape. In these diagrams, the first tomographic image shown in (a) shows the frequency of Compton scattering occurrence in shades of gray, with a lighter color indicating a higher frequency of Compton scattering occurrence. The graph shown in (b) shows the distribution of the frequency of Compton scattering occurrence on the straight line shown in (a).

[0101] FIG. 13 is a first tomographic image obtained based on a coincidence event in which the energy of the gamma-ray photon after Compton scattering is in the range of 0 to 260 keV. FIG. 14 is a first tomographic image obtained based on a coincidence event in which the energy of the gamma-ray photon after Compton scattering is in the range of 260 to 370 keV. FIG. 15 is a first tomographic image obtained based on a coincidence event in which the energy of the gamma-ray photon after Compton scattering is in the range of 370 to 511 keV. FIG. 16 is a first tomographic image obtained based on a coincidence event in which the energy of the gamma-ray photon after Compton scattering is in the range of 0 to 511 keV.

[0102] As can be seen by comparing Figures 13 to 16, the resolution of the first tomographic image (Figure 15) obtained based on the coincidence events in which the energy of the 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 (Figure 13) obtained based on the coincidence events in which the energy of the 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 (Figure 14) obtained based on the coincidence events in which the energy of the gamma-ray photons after Compton scattering is in the range of 260 to 370 keV is even better. In this way, by selectively determining the Compton scattering position C for events in which the energy of the Compton scattered gamma-ray photons is lower than the first threshold value E1, a first tomographic image with excellent spatial resolution can be obtained.

[0103] 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 phenomena.

[0104] 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, the other gamma ray photon is incident on the first detector 11 approximately perpendicularly. 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 gamma ray photon after scattering does not enter 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. Note that, although some Compton scattering occurrence events are observed in the central region of the subject 90 in FIG. 13, this is considered to be due to the detection of an event in which a gamma ray photon whose energy was 260 keV or more after forward scattering in the central region of the subject 90 is further Compton-scattered in the second detector 12 and whose energy became 260 keV or less. Moreover, the fact that the frequency of Compton scattering occurrence is high in the region around the subject 90 in FIG. 13 is also considered to be due to the same factor.

[0105] Since the second detector 12 cannot determine whether a gamma-ray photon has been photoelectrically absorbed or Compton-scattered, it is preferable to exclude events in which the energy of the gamma-ray photon is equal to or less than a certain threshold. That is, it is preferable that the processing unit 20 selectively determines the position C where the gamma-ray photon has been Compton-scattered in the subject when the gamma-ray photon energy represented by the signal output from the second detector 12 is lower than the first threshold E1 and higher than the second threshold E2. In this way, the spatial resolution of the detection of the position C can be further reduced, and a first tomographic image with even better spatial resolution can be obtained. The second threshold E2 can be set to, for example, 150 keV, 200 keV, 230 keV, 260 keV, 300 keV, 330 keV, etc. However, the second threshold E2 is lower than the first threshold E1. It is also preferable that the second threshold E2 can be set arbitrarily.

[0106] 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.

[0107] (Modification) The present invention is not limited to the above-described embodiment, and various modifications are possible. For example, the configurations of any two or more of the above-described embodiments may be combined. [Explanation of symbols]

[0108] 1A to 1F... image acquisition device, 10, 10D, 10E, 10F... measurement section, 11, 11D... first detector, 12... second detector, 13... shield, 14... moving section, 20... processing section, 30... display section, 81, 82, 83... positron-emitting nuclide, 90... specimen.

Claims

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; a processing unit that processes signals output from the first detector and the second detector; Equipped with In a first measurement mode, the measurement unit outputs signals representing detection positions, detection times, and gamma ray photon energies of gamma ray photons by the first detector and the second detector, with a test object being placed between the first detector and the second detector, and a positron-emitting nuclide being placed between the first detector or the second detector and the test object; 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 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 processing unit 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. Image acquisition device.

2. In the second measurement mode, the measurement unit outputs signals representing detection positions, detection times, and gamma ray photon energies of the gamma ray photons by the first detector and the second detector, with the subject administered with a drug labeled with a positron emitting nuclide being placed between the first detector and the second detector; The processing unit includes: In the second 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, a position at which 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, and a second tomographic image is created that represents a distribution of the positions at which the annihilation events occurred in the subject determined for each of the multiple coincidence events; correcting the second tomographic image based on the first tomographic image; The image acquisition device of claim 1 .

3. 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; a processing unit that processes signals output from the first detector and the second detector; Equipped with the measurement unit outputs signals representing detection positions, detection times, and gamma ray photon energies of gamma ray photons by the first detector and the second detector, 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; The processing unit includes: 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 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 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, determining a position where the annihilation event occurred based on the detection positions and detection times of the gamma ray photons by the first detector and the second detector, respectively, when the gamma ray photons that have arrived at both the first detector and the second detector have arrived without being Compton scattered within the subject; 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; Image acquisition device.

4. the processing unit selectively determines a position where 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. The image acquisition device according to any one of claims 1 to 3.

5. the positron-emitting nuclide placed between the first detector or the second detector and the subject and the positron-emitting nuclide labeling the drug administered to the subject are the same type as each other; 4. An image acquisition device according to claim 2 or 3.

6. The processing unit determines whether or not the 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. The image acquisition device according to any one of claims 1 to 3.

7. the measurement unit outputs signals representing detection positions, detection times, and gamma ray photon energies of the gamma ray photons by the first detector and the second detector, in a state in which a positron emitting nuclide is placed between the first detector and the subject, and a positron emitting nuclide is also placed between the second detector and the subject; The image acquisition device according to any one of claims 1 to 3.

8. the measurement unit outputs signals representing detection positions, detection times, and gamma ray photon energies of the gamma ray photons detected by the first detector and the second detector, in a state in which a 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; The image acquisition device according to any one of claims 1 to 3.

9. The measurement unit further includes a shield that prevents gamma ray photons backscattered from one of the first detector and the second detector from entering the other detector. The image acquisition device according to any one of claims 1 to 3.

10. The measurement unit further includes a transfer unit that transfers a positron-emitting nuclide between the first detector or the second detector and the subject. The image acquisition device according to any one of claims 1 to 3.

11. a measuring 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; a processing step of processing signals output from each of the first detector and the second detector; Equipped with In the measurement step, in a first measurement mode, a subject 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 signals representing detection positions, detection times, and gamma-ray photon energies of gamma-ray photons by the first detector and the second detector, respectively, are outputted; In the first measurement mode, the processing step selectively 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, 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 being Compton scattered 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, and creates a first tomographic image showing a distribution of Compton scattering positions in the subject determined for each of a plurality of coincidence events. Image acquisition method.

12. In the measurement step, in a second measurement mode, a signal representing a detection position, a detection time, and a gamma ray photon energy by each of the first detector and the second detector is outputted 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; The processing steps include: In the second 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, a position at which 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, and a second tomographic image is created that represents a distribution of the positions at which the annihilation events occurred in the subject determined for each of the multiple coincidence events; correcting the second tomographic image based on the first tomographic image; The image acquisition method according to claim 11.

13. a measuring 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; a processing step of processing signals output from each of the first detector and the second detector; Equipped with The measuring step includes: placing a subject administered with a drug labeled with a positron-emitting nuclide between the first detector and the second detector; and outputting signals representing detection positions, detection times, and gamma-ray photon energies of gamma-ray photons by the first detector and the second detector, with a positron-emitting nuclide being placed between the first detector or the second detector and the subject; The processing steps include: 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 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 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, determining a position where the annihilation event occurred based on the detection positions and detection times of the gamma ray photons by the first detector and the second detector, respectively, when the gamma ray photons that have arrived at both the first detector and the second detector have arrived without being Compton scattered within the subject; 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; Image acquisition method.

14. The processing step selectively determines a position where 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 a second threshold value. The image acquisition method according to any one of claims 11 to 13.

15. the positron-emitting nuclide placed between the first detector or the second detector and the subject and the positron-emitting nuclide labeling the drug administered to the subject are the same type as each other; 14. The image acquisition method according to claim 12 or 13.

16. The processing step determines whether or not the 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 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. The image acquisition method according to any one of claims 11 to 13.

17. the measuring step includes placing a positron-emitting nuclide between the first detector and the subject, and placing a positron-emitting nuclide between the second detector and the subject, and outputting signals representing detection positions, detection times, and gamma-ray photon energies of the gamma-ray photons by the first detector and the second detector, respectively. The image acquisition method according to any one of claims 11 to 13.

18. The measuring step uses the first detector having a narrower detection surface than the second detector, and outputs signals representing detection positions, detection times, and gamma ray photon energies of the gamma ray photons detected by the first detector and the second detector, respectively, in a state in which a positron-emitting nuclide is placed between the first detector and the subject. The image acquisition method according to any one of claims 11 to 13.

19. The measuring step includes using a shield to prevent gamma ray photons backscattered from one of the first detector and the second detector from entering the other detector. The image acquisition method according to any one of claims 11 to 13.

20. The measuring step comprises transferring a positron-emitting nuclide between the first detector or the second detector and the subject. The image acquisition method according to any one of claims 11 to 13.

Images