JP2013083580A - Positron emission tomographic device and method - Google Patents

Abstract

To improve image resolution.
A positron emission tomography apparatus according to an embodiment includes an FE circuit that is a counting information output unit, a coincidence counting information generation unit, and an image generation unit. The FE circuit 15 outputs a gamma ray detection position, energy, and detection time as count information of a detector having a plurality of detector modules 14 that count light derived from gamma rays. The coincidence information generation unit 25 calculates a predetermined number of pieces of count information obtained by almost simultaneously counting three or more predetermined numbers of gamma rays emitted by the pair annihilation phenomenon, and the sum of energy values generated by the pair annihilation phenomenon. And a combination of a predetermined number of searched count information is generated as the first coincidence information. The image generation unit 26 generates the first image based on the occurrence location of the pair annihilation phenomenon identified from the detection position of each piece of count information constituting the first coincidence information.
[Selection] Figure 1

Description

  Embodiments described herein relate generally to a positron emission tomography apparatus and method.

  2. Description of the Related Art A positron emission computed tomography (PET) apparatus is known as a medical image diagnostic apparatus capable of performing functional diagnosis on a living tissue of a subject. In the PET examination, a drug labeled with a positron emitting nuclide is administered to a subject. Conventionally, a PET device takes in a drug by utilizing the fact that two photons (two gamma rays) are emitted in approximately opposite directions when a positron emitted from the drug combines with an electron and annihilates. A PET image showing the tissue distribution of the subject is reconstructed.

  Specifically, the PET apparatus reconstructs a PET image on the assumption that there is a location where a pair annihilation occurs on a line connecting two detection positions at which two photons emitted in substantially opposite directions are detected substantially simultaneously. ing. The line connecting the two detection positions is called LOR (Line Of Response). Conventionally, the location where the pair disappearance cannot be determined uniquely on the LOR has been a cause of resolution degradation of the PET image.

  For this reason, development of a technique for improving the resolution of a PET image is being promoted. For example, in recent years, development of TOF (Time Of Flight) -PET apparatus that identifies a place where a pair annihilation is highly likely to occur on the LOR using a difference in detection time between two gamma rays emitted by pair annihilation. Is underway. However, in order to perform TOF-PET, it is required to measure the detection time with an accuracy of picosecond order, for example.

JP 2009-42029 A

  The problem to be solved by the present invention is to provide a positron emission tomography apparatus and method capable of improving the image resolution.

  The positron emission tomography apparatus according to the embodiment includes a counting information output unit, a coincidence counting information generation unit, and an image generation unit. The count information output unit outputs a detection position, energy, and detection time of gamma rays as count information of a detector having a plurality of detector modules that count light derived from gamma rays. The coincidence information generation unit calculates the predetermined number of pieces of count information obtained by simultaneously counting three or more predetermined numbers of gamma rays emitted by the pair annihilation phenomenon, and sums the energy values generated by the pair annihilation phenomenon. And a combination of the searched predetermined number of pieces of count information is generated as first coincidence count information. An image generation part produces | generates a 1st image based on the generation | occurrence | production location of the pair annihilation phenomenon specified from the detection position of each count information which comprises said 1st coincidence count information.

FIG. 1 is a diagram for explaining a configuration example of a PET apparatus according to the present embodiment. FIG. 2 is a diagram for explaining a configuration example of the detector module according to the present embodiment. FIG. 3 is a diagram (1) for explaining a processing example of the FE circuit according to the present embodiment. FIG. 4 is a diagram (2) for explaining a processing example of the FE circuit according to the present embodiment. FIG. 5 is a diagram (3) for explaining a processing example of the FE circuit according to the present embodiment. FIG. 6 is a diagram for explaining an example of the count information collection / storage unit according to the present embodiment. FIG. 7 is a diagram for explaining the two-photon coincidence counting information. FIG. 8 is a diagram for explaining the three-photon coincidence counting information. FIG. 9 is a diagram for explaining the identification process of the occurrence portion of the pair annihilation phenomenon performed by the image generation unit according to the present embodiment. FIG. 10 is a diagram for describing a specific example of display control of the control unit according to the present embodiment. FIG. 11 is a flowchart for explaining processing of the PET apparatus according to the present embodiment.

  Hereinafter, embodiments of a positron emission computed tomography (PET) apparatus will be described in detail with reference to the accompanying drawings. In the following, the positron emission tomography apparatus is abbreviated as PET apparatus.

(Embodiment)
FIG. 1 is a diagram for explaining a configuration example of a PET apparatus according to the present embodiment. As shown in FIG. 1, the PET apparatus according to this embodiment includes a gantry device 10 and a console device 20.

  The gantry device 10 is a device that counts gamma rays emitted by a positron emitting nuclide that is administered to a subject P and selectively taken into a living tissue of the subject P in a predetermined monitoring period. As shown in FIG. 1, the gantry device 10 includes a top plate 11, a bed 12, a bed driving unit 13, a detector module 14, an FE (Front End) circuit 15, and a counting information collection unit 16. . As shown in FIG. 1, the gantry device 10 has a cavity serving as a photographing port.

  The top plate 11 is a bed on which the subject P lies, and is placed on the bed 12. The bed driving unit 13 moves the subject P into the imaging port of the gantry device 10 by moving the bed 12 under the control of a bed control unit 23 described later.

  The detector module 14 is a photon counting detector that detects gamma rays emitted from the subject P. For example, as shown in FIG. 1, the gantry device 10 according to the present embodiment includes a detector having a plurality of detector modules 14 arranged so as to surround the subject P in a ring shape. FIG. 2 is a diagram for explaining a configuration example of the detector module according to the present embodiment.

  For example, as shown in FIG. 2, the detector module 14 is an Anger-type detector module having a scintillator 141, a photomultiplier tube (PMT) 142, and a light guide 143.

  The scintillator 141 is a crystal such as NaI, LYSO, or BGO that converts gamma rays emitted from the subject P and incident to visible light. In the detector module 14, as shown in FIG. 2, a plurality of scintillators 141 are two-dimensionally arranged. The photomultiplier tube 142 is a device that multiplies the visible light output from the scintillator 141 and converts it into an electric signal. As shown in FIG. 2, a plurality of photomultiplier tubes 142 are arranged densely via a light guide 143. ing. The light guide 143 is used to transmit visible light output from the scintillator 141 to the photomultiplier tube 142, and is made of, for example, a plastic material having excellent light transmittance such as methyl methacrylate (MMA).

  The photomultiplier tube 142 includes a photocathode that receives scintillation light and generates photoelectrons, a multistage dynode that provides an electric field that accelerates the generated photoelectrons, and an anode that is an outlet for electrons. Electrons emitted from the photocathode due to the photoelectric effect are accelerated toward the dynode, collide with the surface of the dynode, and knock out a plurality of electrons. By repeating this phenomenon over multiple dynodes, the number of electrons is avalancheally increased, and the number of electrons at the anode reaches about 1 million. In such an example, the gain factor of the photomultiplier tube 142 is 1 million times. In addition, a voltage of 600 volts or more is normally applied between the dynode and the anode for amplification using the avalanche phenomenon.

  Thus, the detector module 14 counts the gamma rays emitted from the subject P by converting the gamma rays into visible light by the scintillator 141 and converting the converted visible light into an electric signal by the photomultiplier tube 142. To do.

  The FE circuit 15 shown in FIG. 1 is connected to the subsequent stage of each of the plurality of photomultiplier tubes 142 included in each of the plurality of detector modules 14 for counting light derived from gamma rays, and is connected to the front stage (Front End) of the count information collecting unit 16. ). The FE circuit 15 outputs a gamma ray detection position, energy, and detection time as count information of a detector having a plurality of detector modules 14. That is, the FE circuit 15 functions as a count information output unit. For example, the FE circuit 15 performs the following measurement processing from the electrical signal output from each photomultiplier tube 142 to generate measurement data of the detection position, energy, and detection time of the gamma rays, and the measurement data is counted. Is output to the counting information collecting unit 16. 3 to 5 are diagrams for explaining processing examples of the FE circuit according to the present embodiment.

  The FE circuit 15 measures the energy of the detected (counted) gamma rays by performing waveform shaping processing on the analog waveform data of the electrical signal output from each photomultiplier tube 142. For example, as shown in FIG. 3, the FE circuit 15 performs arithmetic processing (integration processing and differentiation processing) on the analog waveform of the electrical signal output from each photomultiplier tube 142, so that the wave height becomes energy. Generate data. Using such data, the FE circuit 15 measures the energy (E) of gamma rays converted into visible light.

Further, the FE circuit 15 measures the time (detection time) when the gamma rays are detected from the analog waveform data of the electrical signal output from each photomultiplier tube 142. For example, as shown in FIG. 4, the FE circuit 15 measures the time when the threshold value (TH) of a preset voltage value is reached in the analog waveform data as the gamma ray detection time (T). For example, the FE circuit 15 measures the detection time (T) with an accuracy of 10 ⁻¹² seconds (psec). Here, the detection time (T) may be the absolute time (time), or may be the relative time from the PET image capturing start time.

  Further, the FE circuit 15 discriminates the incident position of the gamma rays by, for example, an anger type position calculation process. Specifically, the FE circuit 15 converts the plurality of visible lights output from the scintillator 141 into electric signals at substantially the same timing, and the gamma rays corresponding to the intensity of each electric signal. The position of the center of gravity is calculated from the energy of. Then, the FE circuit 15 determines a scintillator number (P) indicating the position of the scintillator where the gamma rays are incident from the position of the center of gravity obtained as the calculation result. When the photomultiplier tube 142 is a position detection type photomultiplier tube, the measurement data of the detection position is output from the photomultiplier tube 142.

  Then, the FE circuit 15 outputs the measurement data generated by the above measurement process to the count information collecting unit 16 as the count information of the detector. For example, as shown in FIG. 5, the FE circuit 15 associates “P: scintillator number”, “E: energy” and “T: associated with“ module ID ”for uniquely identifying the detector module 14. The “detection time” ”is output to the count information collecting unit 16 as count information.

  The count information collection unit 16 illustrated in FIG. 1 collects the measurement information output from the FE circuit 15 and transmits the collected count information to the console device 20.

  The console device 20 shown in FIG. 1 is a device that receives an operation of the PET device by an operator and generates a PET image from the count information collected by the gantry device 10.

  As shown in FIG. 1, the console device 20 includes an input unit 21, a display unit 22, a bed control unit 23, a count information storage unit 24, a coincidence count information generation unit 25, an image generation unit 26, data Each unit of the console device 20 includes a storage unit 27 and a control unit 28, and is connected via an internal bus.

  The input unit 21 includes a mouse, a keyboard, and the like used by the operator of the PET apparatus for inputting various instructions and various settings, and transfers the instructions and setting information received from the operator to the control unit 28. For example, the input unit 21 receives conditions for generating a PET image from the operator.

  The display unit 22 is a monitor that is referred to by the operator, for displaying a PET image under the control of the control unit 28, and for receiving various instructions and various settings from the operator via the input unit 21. GUI (Graphical User Interface) is displayed.

  The bed control unit 23 moves the subject P into the imaging port of the gantry device 10 by controlling the bed driving unit 13.

  The count information storage unit 24 stores the count information collected by the count information collection unit 16. FIG. 6 is a diagram for explaining an example of the count information storage unit according to the present embodiment.

  For example, as illustrated in FIG. 6, the count information storage unit 24 includes “P: P11, E: E11, T: T11” as count information collected from the count result by the detector module 14 of “module ID: D1”. And “P: P12, E: E12, T: T12” and the like are stored. In the figure, “P”, “E”, and “T” indicate “scintillator number”, “energy”, and “detection time”, respectively.

  Further, as shown in FIG. 6, the count information storage unit 24 similarly stores count information collected from the count results of the “module ID: D2” and “module ID: D3” detector modules 14.

  The simultaneous count information generation unit 25 illustrated in FIG. 1 searches the count information stored in the count information storage unit 24 for a plurality of pieces of count information obtained by counting a plurality of gamma rays emitted by the pair annihilation phenomenon substantially simultaneously. Then, the coincidence count information generation unit 25 generates a combination of the searched plural pieces of count information as coincidence count information.

  Here, PET inspection using the PET apparatus will be described again. In the PET examination, a drug labeled with a positron emitting nuclide is administered to the subject P. For example, in a PET examination for cancer screening, 18F-labeled deoxyglucose labeled with “18F (fluorine)” which is a positron emitting nuclide is administered to the subject P. Such drugs accumulate at a specific site such as a tumor site of the subject P and emit positrons. Positrons combine with surrounding electrons to annihilate pairs, and multiphotons (multiple gamma rays) are emitted by pair annihilation. Since the total energy generated by the pair annihilation is “1022 keV”, the total energy of the multiphotons emitted by the pair annihilation is “1022 keV”. Also, from the law of conservation of momentum, the sum of the motion vectors of each multiphoton is a zero vector.

  Conventionally, by utilizing the fact that two photons (two gamma rays) having energy of 511 keV are emitted in substantially opposite directions due to the pair annihilation phenomenon, a PET image showing the tissue distribution of the subject that has taken in the drug is reconstructed. doing. In the two-photon emission, the energy value of each photon (gamma ray) is “511 keV”, and the angle between the motion vectors of the two photons is 180 degrees from the law of conservation of momentum. FIG. 7 is a diagram for explaining the two-photon coincidence counting information. Conventionally, as shown in FIG. 7A, it is assumed that there is a location where pair annihilation occurs on a line connecting two detection positions (scintillator numbers) at which two photons are detected almost simultaneously, and the PET image is reproduced. It was composed. The line connecting the two detection positions is called LOR (Line Of Response).

  For example, in the case of generating two-photon coincidence counting information, the coincidence counting information generating unit 25 abbreviates two photons according to the condition set by the operator or initially set (two-photon coincidence counting information generation condition). Two pieces of count information counted simultaneously are searched. Then, the coincidence count information generation unit 25 generates a combination of the searched two pieces of count information as two-photon coincidence count information. As an example, a time window width “t” is set as the two-photon coincidence counting information generation condition. In addition, as a two-photon coincidence count information generation condition, an energy window width “(511−e1) ≦ E ≦ (511 + e2)” may be set. Note that “t”, “e1”, and “e2” are set according to, for example, the measurement accuracy for time measurement and energy measurement of the FE circuit 15.

  For example, as shown in FIG. 7B, the coincidence information generation unit 25 generates ““ | Tn1−Tm2 | ≦ t ”,“ (511−e1) ≦ En1 ≦ (511 + e2) ”,“ (511− e1) ≦ Em2 ≦ (511 + e2) ”” is searched for two pieces of count information. Thereby, as shown in FIG. 7B, the coincidence counting information generation unit 25 counts information “P: Pn1, E: En1, T: Tn1” derived from the detector module 14 of “module ID: n”. And a combination of the count information “P: Pm2, E: Em2, T: Tm2” derived from the detector module 14 of “module ID: m” is generated as two-photon coincidence count information in which two photons are detected substantially simultaneously. .

  By generating the two-photon coincidence information, the location where the pair annihilation has occurred is specified to exist on the LOR connecting “Pn1” and “Pm2”, for example. However, conventionally, the location where the pair annihilation has occurred cannot be uniquely determined on the LOR, which has been a cause of degradation in resolution of the PET image.

  Here, it is known that the number of gamma rays emitted by the pair annihilation phenomenon is not limited to two but may be three or more. For example, it is known that three-photon (three gamma rays) emission is generated by the pair annihilation phenomenon with a probability of about “1/370” of the probability that two photons (two gamma rays) are emitted.

  Therefore, the coincidence information generation unit 25 according to the present embodiment uses a predetermined number of pieces of count information obtained by almost simultaneously counting three or more predetermined numbers of gamma rays emitted by the pair annihilation phenomenon, and the sum of energy values is pair annihilated. The search is made on the condition that the total energy generated by the phenomenon is substantially equal. Then, the coincidence counting information generating unit 25 according to the present embodiment generates a combination of the searched predetermined number of counting information as the first coincidence counting information. Specifically, the coincidence counting information generating unit 25 according to the present embodiment sets the predetermined number to “3” and generates three-photon coincidence counting information as the first coincidence counting information. FIG. 8 is a diagram for explaining the three-photon coincidence counting information.

  As described above, the total energy of the multiphotons emitted by pair annihilation is “1022 keV”. That is, as shown in FIG. 8, the sum of the energy values of the three pieces of count information obtained by detecting the three photons emitted simultaneously due to the pair annihilation phenomenon is substantially equal to “1022 keV”.

  When generating the three-photon coincidence information, the coincidence information generation unit 25 searches for three pieces of count information obtained by counting three photons almost simultaneously according to the operator or the initial three-photon coincidence information generation condition. To do. Then, the coincidence counting information generation unit 25 generates a combination of the searched three pieces of counting information as three-photon coincidence counting information.

  As an example, the three-photon coincidence information generation condition is set with a time window width “t” as in the two-photon coincidence information generation condition, and the energy window width is “the sum of the three energy values. , (1022-e3) or more and (1022 + e4) or less ”is set.

  Note that “t”, “e3”, and “e4” are set according to, for example, the measurement accuracy for time measurement and energy measurement of the FE circuit 15. Furthermore, “e3” and “e4” do not incorporate the counting information of one of the two photons emitted at the same time as one of the three counting information of the three photons emitted at the same time into the three-photon coincidence counting information. Thus, it is a set value.

  For example, as shown in FIG. 8B, the coincidence information generation unit 25 generates ““ | Tk3-Tj4 | ≦ t ”,“ | Tk3-Tw4 | ≦ t ”,“ | Tj4-Tw4 | ≦ t ”. ”” And three pieces of count information satisfying “(1022−e3) ≦ (Ek3 + Ej4 + Ew4) ≦ (1022 + e4)” are searched. Thereby, as shown in FIG. 8B, the coincidence counting information generation unit 25 counts information “P: Pk3, E: Ek3, T: Tk3” derived from the detector module 14 of “module ID: k”. And the counting information “P: Pj4, E: Ej4, T: Tj4” derived from the detector module 14 with “module ID: j” and the counting information “P: from the detector module 14 with“ module ID: w ”. "Pw4, E: Ew4, T: Tw4" are generated as the three-photon coincidence counting information in which three photons are simultaneously detected.

  From the count information stored in the count information storage unit 24, the coincidence count information generation unit 25 according to the present embodiment performs a search process using the three-photon coincidence count information generation condition described as an example in FIG. Three-photon coincidence count information is generated by combining three pieces of count information as first coincidence information.

  Furthermore, the coincidence counting information generation unit 25 according to the present embodiment stores the counting information stored in the counting information storage unit 24 by the search process using the two-photon coincidence counting information generation condition described as an example in FIG. Thus, the two-photon coincidence information is generated as the second coincidence information as in the conventional case.

  The coincidence information generation unit 25 stores the first coincidence information and the second coincidence information in the coincidence information data 27a of the data storage unit 27 shown in FIG. For example, the coincidence information generation unit 25 adds incidental information indicating that it is three-photon coincidence information to the first coincidence information, and incidental information that indicates that it is two-photon coincidence information as the second coincidence information. And stored in the coincidence counting information data 27a. In other words, the coincidence information generation unit 25 assigns incidental information indicating that it is projection data (sinogram) of gamma rays to the second coincidence information and stores it in the coincidence information data 27a.

  Note that a plurality of visible light outputs derived from a plurality of gamma rays incident on the same detector module 14 at the same time are usually excluded from measurement targets by the FE circuit 15. Therefore, the first coincidence information (three-photon coincidence information) generated by the coincidence information generation unit 25 according to the present embodiment includes three counts obtained by counting three gamma rays simultaneously incident on three different detector modules 14. The data is a combination of information. However, since the angle at which the three photons are emitted is random, the probability that two photons or three photons are incident on the same detector module 14 during the three-photon emission is low. For this reason, the first coincidence information generated in the present embodiment can be regarded as data that substantially covers the three-photon emission phenomenon.

  The image generation unit 26 shown in FIG. 1 generates a PET image using the coincidence counting information stored in the coincidence counting information data 27a. Specifically, the image generation unit 26 generates a three-photon PET image based on the occurrence point of the pair annihilation phenomenon specified from the detection position of each piece of count information constituting the first coincidence information. Hereinafter, a three-photon PET image based on the first coincidence information that is the three-photon coincidence information is referred to as a “first image” or a “three-photon image”. FIG. 9 is a diagram for explaining the identification process of the occurrence portion of the pair annihilation phenomenon performed by the image generation unit according to the present embodiment.

  As described above, from the law of conservation of momentum, the sum of the motion vectors of each multiphoton is a zero vector. Therefore, if the detection position of a plurality of photons of three or more photons is specified, the occurrence location of the pair annihilation phenomenon can be uniquely specified. For example, as shown in FIG. 9, the position of “Pk3” is point A, the position of “Pj4” is point B, the position of “Pw4” is point C, and the location where the pair annihilation phenomenon occurs is point G. . The image generation unit 26 calculates the position where the sum of the “vector vector unit vector GA ′”, “vector GB unit vector GB ′”, and “vector GC unit vector GC ′” is “zero vector” by an arithmetic process. As a result, the occurrence point “G” of the pair annihilation phenomenon is specified. In other words, “G” is the center of gravity of “triangle A′B′C ′”.

  The image generation unit 26 performs the above-described center-of-gravity calculation, thereby specifying the occurrence location with each first coincidence information. And the image generation part 26 produces | generates a 1st image by allocating the pixel value according to specific frequency to the pixel corresponding to each specified generation | occurrence | production location.

  Further, the image generation unit 26 generates a PET image by reconstructing the second coincidence information. Hereinafter, a two-photon PET image based on the second coincidence information that is the two-photon coincidence information is referred to as a “second image” or a “two-photon image”.

  Specifically, the image generation unit 26 reconstructs the second image by a successive approximation method using second coincidence information that is gamma ray projection data (sinogram). For example, the image generation unit 26 reconstructs the second image using an MLEM (Maximum Likelihood Expectation Maximization) method or an OSEM (Ordered Subset MLEM) method as a successive approximation method.

  Then, the image generation unit 26 stores the first image (three-photon image) and the second image (two-photon image) in the image data 27b of the data storage unit 27 illustrated in FIG.

  The control unit 28 illustrated in FIG. 1 performs overall control of the PET apparatus by controlling the operations of the gantry device 10 and the console device 20. Specifically, the control unit 28 controls the movement of the bed 12 and the counting information collecting process in the counting information collecting unit 16. Further, the control unit 28 controls the coincidence counting information generation process in the coincidence counting information generation unit 25 and the image generation process in the image generation unit 26. In addition, the control unit 28 performs control so that the image data stored in the image data 27 b is displayed on the display unit 22. FIG. 10 is a diagram for describing a specific example of display control of the control unit according to the present embodiment.

  Specifically, as shown in FIG. 10A, the control unit 28 causes the display unit 22 to display the three-photon image that is the first image and the two-photon image that is the second image in parallel.

  Alternatively, the control unit 28 causes the display unit 22 to display a superimposed image obtained by superimposing the first image and the second image after changing the color tone of each other. For example, the image generation unit 26 converts the three-photon image, which is the first image, from a gray scale to a color scale image based on red. Then, the image generation unit 26 generates a superimposed image in which the converted three-photon image and the grayscale two-photon image are superimposed. Thereby, the display part 22 displays the superimposed image shown to (B) of FIG. When displaying the superimposed image, the two-photon image may be converted from a gray scale to a color scale image, and a superimposed image obtained by superimposing the converted two-photon image and the gray-scale three-photon image may be displayed.

  As described above, the probability that three gamma rays are emitted by the pair annihilation phenomenon is “1/370” of the probability that two gamma rays are emitted by the pair annihilation phenomenon. That is, the first image is a PET image in which a region where the occurrence frequency of the pair annihilation phenomenon is high due to accumulation of a large amount of drugs is depicted as a set of points with high resolution. On the other hand, the second image uses the LOR formed by detecting the two-photon emission that occurs with high probability due to the pair annihilation phenomenon, so that the region where the annihilation phenomenon may have occurred is low resolution. It becomes a comprehensive PET image.

  For example, the operator can efficiently perform image diagnosis by comparing and referring to the two-photon image and the three-photon image shown in FIG. 10A or referring to the superimposed image shown in FIG. Can be done automatically. That is, the operator can roughly grasp the tissue in which the medicine is accumulated and the distribution in the tissue from the two-photon image. The operator can further specify the position in the tissue where the medicine is accumulated with fine granularity by using the three-photon image. For example, by referring to the superimposed image shown in FIG. 10B, the operator can confirm that the central portion (highly integrated portion) of the integrated region depicted by the two-photon image is actually several mm (millimeters). It can be grasped that it is shifted to the left side. Further, by referring to the superimposed image, for example, the operator can grasp that the integrated region depicted by the two-photon image is actually an area formed by a plurality of highly integrated locations. .

  Note that this embodiment is applicable even when only the first image, that is, the three-photon image is generated and displayed.

  Next, processing of the PET apparatus according to the present embodiment will be described using FIG. FIG. 11 is a flowchart for explaining processing of the PET apparatus according to the present embodiment. In the following, a case where the superimposed display of the first image and the second image is set will be described.

  As shown in FIG. 11, the PET apparatus according to the present embodiment receives a PET image capturing request from the operator via the input unit 21 after moving the subject P into the imaging port of the gantry device 10. It is determined whether or not (step S101). Here, when an imaging request is not accepted (No in step S101), the PET apparatus according to the present embodiment stands by until an imaging request is accepted.

  On the other hand, when the imaging request is received (step S101 Yes), the count information collection unit 16 collects the count information output from the FE circuit 15 (step S102). That is, the FE circuit 15 generates count information obtained by measuring the detection position, energy, and detection time of the gamma rays based on the count result of each detector module 14 in a predetermined monitoring period. The count information collecting unit 16 collects the count information output from the FE circuit 15 for each of the plurality of detector modules 14, for example.

  Then, the count information collection unit 16 stores the collected count information in the count information storage unit 24 of the console device 20 (step S103), and the coincidence count information generation unit 25 performs the two-photon coincidence count information generation condition and the three-photon coincidence. Based on the count information generation condition, two-photon coincidence count information (second coincidence count information) and three-photon coincidence count information (first coincidence count information) are generated (step S104).

  Thereafter, the image generation unit 26 generates a two-photon image (second image) and a three-photon image (first image) (step S105). That is, the image generation unit 26 generates a two-photon image by reconstructing the two-photon coincidence information. In addition, the image generation unit 26 generates a three-photon image by specifying the occurrence location of the pair annihilation phenomenon from the detection positions of the three pieces of count information configuring the three-photon coincidence information.

  Subsequently, under the control of the control unit 28, the image generation unit 26 generates a superimposed image of the two-photon image and the three-photon image (step S106). Then, under the control of the control unit 28, the display unit 22 displays a superimposed image (step S107) and ends the process.

  As described above, in the present embodiment, the position where the multiphotons of three or more photons are detected almost simultaneously is specified using the sum of energy values as a search condition, and the pair annihilation phenomenon is detected from the specified three or more detected positions. Uniquely identify the location. In the present embodiment, a first image that is a PET image is generated using the identified occurrence point of the pair annihilation phenomenon. That is, the first image is a PET image in which a region where the occurrence frequency of the pair annihilation phenomenon is high is depicted as a set of points with high resolution. That is, in the present embodiment, the occurrence location of the pair annihilation phenomenon is specified within a range of one pixel or several pixels. Therefore, in this embodiment, the image resolution can be improved. Further, for example, if the detection efficiency of the three-photon emission generated by the pair annihilation phenomenon can be increased, the time for image output can be shortened.

  Further, in the present embodiment, a second image, which is a conventional PET image that covers a region that has a low resolution but may have disappeared, is displayed in parallel with or superimposed on the first image. That is, the operator who is a doctor refers to, for example, the two-photon image and the three-photon image shown in (A) of FIG. 10 by comparison or the superimposed image shown in (B) of FIG. In addition, the distribution in the tissue in which the medicine is accumulated can be specified with high accuracy. Therefore, in this embodiment, image diagnosis in PET inspection can be efficiently supported.

  The second image (two-photon image) may be a PET image generated by TOF (Time Of Flight). In the case of TOF as well, the location of two-photon emission can be uniquely specified on the LOR, but the accuracy of the specified location greatly depends on the measurement accuracy of the detection time. Further, the location on the LOR specified by the TOF is a location with a high probability of being a location where the pair annihilation phenomenon occurs. On the other hand, the process of specifying the detection positions of three or more multiphoton emissions performed in the present embodiment can be performed with an accuracy lower than the time measurement accuracy required by TOF. Further, in the present embodiment, since the occurrence location of the pair annihilation phenomenon can be uniquely specified, the resolution of the occurrence location of the pair annihilation phenomenon specified in this embodiment is higher than that of TOF.

  In the present embodiment, the case where the predetermined number is “3” and the first coincidence information is generated for three gamma rays emitted by the pair annihilation phenomenon has been described. However, the present embodiment is applicable even when four or more gamma rays emitted by the pair annihilation phenomenon are targeted. However, the location where the pair annihilation phenomenon occurs can be specified by specifying three or more gamma ray detection positions. In addition, the occurrence probability of multiphotons decreases by, for example, two digits each time the number of generated photons increases by one. Therefore, in the present embodiment, it is desirable to set the predetermined number to 3.

  In the above description, the case has been described in which the count information is accumulated in the console device 20 and the coincidence count information (first and second coincidence count information) is generated in the console device 20 by, for example, software processing. However, the present embodiment is a case where the gantry device 10 generates coincidence counting information (first and second coincidence counting information) by hardware processing and transmits the generated coincidence counting information to the console device 20. Also good.

  However, in the PET device according to the present embodiment, as described above, the count information is accumulated in the console device 20, so that even after the generation of the PET image (first image and second image), The counting information collected during shooting can be held. With such a configuration, for example, the first coincidence information generation condition and the second coincidence information generation condition can be changed, and a PET image (first image or second image) can be generated again.

  In addition, the search process for the first coincidence information requires a search with a combination of three or more as compared with the search process for the conventional coincidence information (second coincidence information). Increase. That is, the generation of the first coincidence information may be assumed not to be completed at the end of the PET examination. For this reason, it is desirable that the configuration of the present embodiment holds the count information in the console device 20 so that the first coincidence count information can be generated even after the PET examination.

  In the above description, the case where the two-photon image and the three-photon image are displayed in parallel or superimposed in order to supplementarily use the three-photon image in the image diagnosis using the two-photon image has been described. However, the present embodiment may be a case where an image obtained by correcting the two-photon image is generated and displayed based on the information of the three-photon image. In such a case, for example, under the control of the control unit 28, the image generation unit 26 generates a corrected image in which a region that is positive (drug accumulation) in the two-photon image is replaced with a portion that is positive in the three-photon image. . Then, under the control of the control unit 28, the display unit 22 displays the corrected image.

  Or this embodiment may be a case where the information based on 3 photon coincidence information is reflected in the reconstruction process of 2 photon coincidence information which is projection data. In this case, for example, under the control of the control unit 28, the image generation unit 26 specifies the LOR passing through the location “G” specified from the three-photon coincidence information from the two-photon coincidence information. Then, under the control of the control unit 28, the image generation unit 26 performs a process of increasing the weight for projecting and backprojecting to the “G” for the specified LOR when performing the successive approximation method. The image is reconstructed, and the display unit 22 displays the second image.

  Note that the positron emission tomography method described in this embodiment may be executed by a personal computer, a workstation, or the like to which the count information for each PET examination is transferred from the count information storage unit 24. That is, the positron emission tomography method described in the present embodiment can be realized by executing a prepared program on a computer such as a personal computer or a workstation.

  As described above, according to the present embodiment, the image resolution can be improved.

  Although several embodiments of the present invention have been described, these embodiments are presented by way of example and are not intended to limit the scope of the invention. These embodiments can be implemented in various other forms, and various omissions, replacements, and changes can be made without departing from the spirit of the invention. These embodiments and their modifications are included in the scope and gist of the invention, and are also included in the invention described in the claims and the equivalents thereof.

DESCRIPTION OF SYMBOLS 10 Stand apparatus 11 Top plate 12 Bed 13 Bed drive part 14 Detector module 141 Scintillator 142 Photomultiplier tube 143 Light guide 15 FE (Front End) circuit 16 Count information collection part 20 Console apparatus 21 Input part 22 Display part 23 Bed control Unit 24 count information storage unit 25 coincidence count information generation unit 26 image generation unit 27 data storage unit 27a coincidence count information data 27b image data 28 control unit

Claims (5)

As counting information of a detector having a plurality of detector modules that count light derived from gamma rays, a counting information output unit that outputs a detection position, energy, and detection time of gamma rays,
The predetermined number of pieces of count information obtained by counting three or more predetermined numbers of gamma rays emitted due to the pair annihilation phenomenon at the same time is set on condition that the sum of energy values is substantially equal to the sum of energy generated by the pair annihilation phenomenon. A coincidence information generating unit that generates a combination of the predetermined number of pieces of the searched information as first coincidence information;
An image generation unit that generates a first image based on the occurrence location of the pair annihilation phenomenon identified from the detection position of each count information constituting the first coincidence information;
A positron emission tomography apparatus comprising: A control unit for displaying the first image on a display unit;
The positron emission tomography apparatus according to claim 1, further comprising: The coincidence information generation unit further searches for two pieces of count information obtained by counting two gamma rays emitted by the pair annihilation phenomenon substantially simultaneously, and generates a combination of the two pieces of searched count information as second coincidence information. ,
The image generation unit further generates a second image by reconstructing the second coincidence information,
The control unit displays the first image and the second image in parallel on the display unit, or a superimposed image obtained by superimposing the first image and the second image after changing each other's color tone. The positron emission tomography apparatus according to claim 1, wherein: is displayed on the display unit.   The positron emission tomography apparatus according to claim 1, wherein the predetermined number is three. As a counting information of a detector having a plurality of detector modules for counting light derived from gamma rays, a counting information output step for outputting a gamma ray detection position, energy, and detection time;
The predetermined number of pieces of count information obtained by counting three or more predetermined numbers of gamma rays emitted due to the pair annihilation phenomenon at the same time is set on condition that the sum of energy values is substantially equal to the sum of energy generated by the pair annihilation phenomenon. A coincidence counting information generating step for generating a combination of the searched predetermined number of counting information as first coincidence information;
An image generating step of generating a first image based on a location where a pair annihilation phenomenon is identified from a detection position of each count information constituting the first coincidence information;
A positron emission tomography method characterized by comprising:

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