JP4588042B2 - Nuclear medicine diagnostic equipment - Google Patents

Description

  The present invention relates to a nuclear medicine diagnostic apparatus, and more particularly to an apparatus for detecting a gamma ray to identify a place where the gamma ray is generated and imaging a spatial distribution of the gamma ray source.

  Imaging devices that detect a gamma ray generated from a radiolabeled drug administered to a subject and acquire a spatial distribution image of a radioisotope are used for diagnostic imaging in clinical fields such as hospitals.

As a typical apparatus, there is a PET apparatus that performs positron emission tomography (PET). The PET device detects a pair of 511 keV annihilation gamma rays radiated outside the body from the accumulation site of the positron emission type radiolabeled drug administered to the subject by a gamma ray detector arranged oppositely, and performs simultaneous counting. The configuration is such that the determination is performed. Specifically, based on the measurement information of a pair of annihilation gamma rays regarded as being simultaneously counted, the location where the annihilation gamma rays are generated is estimated on the line connecting the simultaneously counted detectors. The PET apparatus is an apparatus that performs image reconstruction based on the gamma-ray direction information obtained in this way, and acquires an in-vivo distribution image of the radioisotope.
JP 2006-189274 A

  Positron tomography using a device that estimates the direction of gamma rays based only on coincidence determination has the problem that image quality deterioration due to angular fluctuation cannot be prevented in principle. In addition to this problem, it is impossible to efficiently remove the effects of simultaneous incidents and subject internal scattering events, and false signals are generated, resulting in poor image quality and image quantification. have.

  A supplementary explanation will be given below regarding the image quality deterioration due to the above-mentioned angle fluctuation. A positron released from the radiolabeled drug reacts with an electron in the body of the subject, and a pair of annihilation gamma rays is generated. In general, positrons emitted from radioisotopes have momentum (ie, are not stationary). In a positron-electron centroid system (a system in which the sum of positron-electron momentum vectors is zero), a pair of annihilation gamma rays are generated 180 degrees opposite to each other. Then, a pair of annihilation gamma rays do not occur 180 degrees opposite each other. The deviation from 180 degrees is the angle swing. In the conventional PET apparatus, it is assumed that the positron loses its initial momentum due to ionization in the body (after resting) and then undergoes a pair annihilation reaction with the electron in the body, and the angle fluctuation is the range of the positron in the body. Convert to (movement distance from drug accumulation point). The converted range of the positron in the body is the fundamental limit of the resolution of the PET image in the conventional PET apparatus, and contributes to the deterioration of the image quality.

  In addition, in the conventional PET apparatus using only coincidence determination, an event in which annihilation gamma rays that do not make a pair are accidentally coincidentally counted (incident coincidence event), or annihilation gamma rays are scattered in the body of the subject. The effect of an event (scattering event in the subject) that changes the original direction of annihilation gamma rays cannot be removed efficiently. These events become false signals and cause deterioration in image quality and quantitativeness.

  The object of the present invention has been made in view of the above-mentioned circumstances, and overcomes the fundamental limitations of the measurement method of the conventional PET apparatus and can obtain a high-definition PET image having high quantitativeness. To provide an apparatus.

  The nuclear medicine diagnostic apparatus according to claim 1, which has been made in order to solve the above-mentioned problem, calculates a generation location of photons emitted from a subject using Compton scattering kinematics, and calculates the calculated information. It is characterized in that different image reconstruction is performed according to the above.

  According to the present invention, an apparatus (referred to as a Compton camera) that uniquely determines the direction of incident gamma rays by Compton scattering kinematics is basically included in the configuration. The Compton camera scatters any gamma rays (photons) emitted from the collection site of the radiolabeled drug administered to the subject to the Compton scatter by the front detector, and detects the Compton scattered gamma rays by the back detector. At the same time, the electrons rebounded by the gamma rays are detected by the former detector. Based on the kinematic information of the scattered gamma rays and recoil electrons, the incident direction of the gamma rays with respect to the detector is calculated, and the generation location of the gamma rays is estimated on the calculated line. In the present invention, a pair of annihilation gamma rays is simultaneously measured to estimate a point where the gamma ray is generated as a point intersected by the calculated two directional lines. The present invention performs different image reconstruction according to the estimated information.

  When considering a positron with a relatively large initial momentum, the rate of pair annihilation in the body of the positron is not negligibly small compared to the rate of ionization (that is, the pair annihilation occurs before the positron comes to rest). In such a case, however, in the present invention, if a pair of annihilation gamma rays are simultaneously measured by a Compton camera, it is not necessary to convert all the influences of the positron angular fluctuations into the range of the positron in the body. That is, all of the angle fluctuations do not cause deterioration in image quality. Further, in the present invention, the location where a gamma ray is generated can be estimated using only information on a single gamma ray, so that an accidental coincidence event does not cause deterioration in image quality or quantification.

  If a pair of annihilation gamma rays is simultaneously measured by the configuration of the Compton camera as in the present invention, the influence of angular fluctuation can be reduced, and as a result, image quality deterioration can be prevented. In addition, when the direction of a pair of annihilation gamma rays deviates from 180 degrees, the present invention can estimate the occurrence location of gamma rays at one point in the three-dimensional space, and as a result, the image quantitativeness can be improved.

  According to the present invention, even when the direction of a pair of annihilation gamma rays is not deviated from 180 degrees, a judgment based on the coincidence counting judgment and the Compton scattering kinematics is imposed (the best mode for carrying out the invention later). (Explained in detail in the column), the simultaneous coincidence event and the in-subject scattering event can be efficiently removed, and the false signal can be reduced. As a result, the image quality and the quantitativeness of the image can be improved. it can. According to the present invention, since a false signal can be reduced and a true signal can be generated, the shooting time can be shortened. Similarly, since the dose of the drug to the subject can be reduced, the exposure of the subject can also be reduced.

  According to the present invention, there is an effect that it is possible to provide a nuclear medicine diagnostic apparatus that can overcome the fundamental limitations of the measurement method of a conventional PET apparatus and obtain a PET image having high definition and high quantitativeness.

  Hereinafter, the present invention will be described with reference to the drawings. Here, when the present invention is applied to a nuclear medicine diagnostic apparatus, that is, a drug labeled with a radioisotope is administered to a subject, and the distribution of the drug in the body is based on information on gamma rays emitted from the subject. An example in the case of imaging will be described. In particular, an example in which a positron emission type radioisotope is used as a radiolabeled drug and thereby annihilation gamma rays are detected will be described.

<First Embodiment>
FIG. 1 is a block diagram showing the apparatus configuration of the nuclear medicine diagnosis apparatus of the present invention including a subject. In FIG. 1, reference numeral 150 indicates a subject who is administered a radiopharmaceutical and lies on an examination table (not shown). Reference numerals 100 and 100 'denote detector heads (detection means). The detector heads 100 and 100 'mainly include front-stage detectors 101 and 101' and rear-stage detectors 102 and 102 'as main components. The detector heads 100 and 100 ′ are arranged opposite to each other around the subject 150.

  In addition to the detector heads 100 and 100 ', the nuclear medicine diagnostic apparatus of the present invention collects signals from the detector heads 100 and 100' and processes them, and a processing unit 110 (signal processing means). An image reconstruction device (image reconstruction means) 120 for reconstructing an image from the reconstructed signal, and an image display device 130 for displaying the reconstructed image. The signal processing device 110, the image reconstruction device 120, and the image display device 130 include a CPU, a memory, and the like, and have a microcomputer function. That is, it has control and calculation functions.

  There may be at least two detector heads 100, 100 'as shown. The arrangement of the detector heads 100, 100 ′ is desirable as the solid angle between the detector heads 100, 100 ′ and the subject 150 increases. An examination table (not shown) may have a drive mechanism. There may be at least one upstream detector 101, 101 ′ for each detector head 100, 100 ′, but there may be a plurality of detectors 101, 101 ′. The upstream detectors 101 and 101 ′ may be gas detectors, liquid detectors, or solid state detectors. Moreover, these combinations may be sufficient. There may be at least one post-stage detector 102, 102 'in each detector head 100, 100', but there may be a plurality of post-stage detectors 102, 102 '. The post-stage detectors 102 and 102 ′ may be gas detectors, liquid detectors, or solid-state detectors similarly to the pre-stage detectors 101 and 101 ′. Moreover, these combinations may be sufficient.

  Each detector head 100, 100 ′ is one device constituting a Compton camera (device based on the advanced Compton camera method), and the front detectors 101, 101 ′ are radiolabeled drugs administered to the subject 150. Compton scatters arbitrary gamma rays radiated outside the body from the accumulation point of the laser beam and detects electrons recoiled by the gamma rays. The rear detectors 102 and 102 'are Compton detectors at the front detectors 101 and 101'. It is a device that detects scattered gamma rays.

  The detector heads 100 and 100 ′ having such a configuration are provided, and the signal processing device 110 and the image reconstruction device 120 perform the following data processing. FIG. 2 is a block diagram relating to data processing performed by the signal processing device 110 and the image reconstruction device 120 of FIG. Note that FIG. 1 is also referred to if necessary.

  In FIG. 2, reference numeral 200 is gamma ray data obtained from each detector head 100, 100 ′, and reference numeral 300 is gamma ray incident on each detector head 100, 100 ′ oppositely arranged from gamma ray data 200. Based on the Compton scattering kinematics, reference numeral 400 guarantees that the energy of each gamma ray and the two gamma rays intersect, Kinematic determination means (method and apparatus) for calculating an angle (crossing angle) formed by gamma rays and determining whether the energy and the crossing angle are within a set range are shown.

  In addition, the reference numeral 500 is simultaneously counted, and a gamma ray generation point is estimated as a certain point in a three-dimensional space by one event from Compton scattering kinematics using a pair of gamma rays whose energy and crossing angle are within a set range. A two-photon Compton camera image reconstruction device that performs image reconstruction, reference numeral 600 is a gamma ray that is not simultaneously counted, and a gamma ray whose energy and crossing angle are outside the set range, and the kinematics of Compton scattering 1 shows a single-photon Compton camera image reconstruction device that estimates an occurrence point of a gamma ray on a certain line in a three-dimensional space in one event and performs image reconstruction. The above-mentioned “single photon” and “two photon” are names in Compton camera image reconstruction based on the meaning of “two photons” based on “one photon”.

  Further, reference numeral 700 indicates an image obtained by the two-photon Compton camera image reconstruction device 500 or the single-photon Compton camera image reconstruction device 600. The image 700 is displayed on the screen of the image display device 130.

  As can be seen from FIG. 2, event selection is performed, and image reconstruction is performed by the two-photon Compton camera image reconstruction device 500 or the single-photon Compton camera image reconstruction device 600. In other words, different image reconstruction is performed according to event selection.

  With the configuration described above, the effects of angular fluctuations can be reduced, accidental simultaneous events and internal scattering events can be efficiently removed, the effects of false signals can be reduced, and high-definition, good quantitative images can be obtained. The resulting device. This point will be specifically described below.

  First, the reason why the image quantification is improved by performing event selection as shown in FIG. 2 and performing image reconstruction using the two-photon Compton camera image reconstruction apparatus 500 will be described. FIG. 3 is an explanatory diagram illustrating the reason why the quantitativeness of the image is improved. In addition, FIG.1 and FIG.2 is also referred as needed.

  In FIG. 3, when a positron 2002 having a relatively large initial momentum is released from the radiolabeled drug administered to the subject 150 and immediately after that, a pair annihilation reaction occurs with the electron 2003 in the subject 150. A pair of annihilation gamma rays 2004 and 2004 'are produced. The annihilation gamma ray 2004 (2004 ′) is incident on the front detector 101 (101 ′) and Compton scattered. Scattered gamma rays 2005 (2005 ′) generated by Compton scattering are detected by the post-stage detector 102 (102 ′). Recoil electrons 2006 (2006 ′) recoiled by Compton scattering are detected by the front detector 101 (101 ′).

  In such a case, if only the simultaneous determination means 300 is implemented (executed), in other words, in the case of the conventional apparatus only for simultaneous counting determination, the generation location of the gamma rays is detected by the preceding detectors 101 of the annihilation gamma rays 2004 and 2004 ′. And 101 'are estimated uniformly on a line 3001 connecting the incident points to 101' (3001 is a generation point of a gamma ray estimated only by coincidence determination). As described above, since the gamma ray source is estimated on the line 3001 in the related art, counting is performed as if there is a gamma ray source even in a place where the gamma ray source does not exist, and the quantitativeness of the image deteriorates.

  However, the present invention imposes the annihilation gamma ray 2004 (2004 ′) on the front-stage detector 101 (101 ′) by imposing a determination based on the kinematics of Compton scattering on the events simultaneously counted by the kinematic determination unit 400. The direction is estimated on the estimated direction 4002 (4002 ′) calculated from the Compton scattering kinematics, and the intersection 4001 of the estimated directions 4002 and 4002 ′ is estimated at the location where the gamma ray is generated (4001 is a coincidence determination and motion) Gamma-ray generation location estimated by academic judgment). As described above, according to the present invention, the gamma ray source can be locally estimated at a three-dimensional point and the gamma rays can be counted, so that the quantitativeness of the image is improved.

  Processing of the simultaneous determination unit 300 and the kinematic determination unit 400 will be described with reference to FIGS. 4 and 5. 4 is a flowchart of “simultaneous determination” according to the simultaneous determination unit 300 of FIG. 2, and FIG. 5 is a flowchart of “kinematic determination” according to the kinematic determination unit 400 of FIG.

  In FIG. 4, the simultaneous determination means 300 determines whether or not the data collection times are synchronized within the set time based on the gamma ray data collected by the detector heads 100 and 100 ′. If they are synchronized, they are determined as “simultaneous counting events” as indicated by Yes in the figure, and if they are not synchronized, they are determined as “events that are not coincident counting” as indicated by No in the figure. The determination of “simultaneous counting event” corresponds to YES in FIG. 2, and the determination of “event not simultaneous counting” corresponds to NO in FIG. 2.

  In FIG. 5, in the kinematic determination means 400, since each detector head 100, 100 'has already collected gamma ray data, the gamma ray is determined from the collected gamma ray data, that is, from the measurement data based on the Compton scattering kinematics. The direction A and the direction A ′ are calculated to determine whether the direction A and the direction A ′ of the gamma rays are spatially overlapped. Then, if they are spatially overlapped, it is determined as “estimate the point where the gamma rays are generated in the direction A and the direction A ′” as indicated by Yes in the figure. Is determined as “estimate the location where the gamma ray is generated on each line in the direction A and the direction A ′” as indicated by No in the figure. The determination “estimate the point where the gamma ray is generated at the point where the direction A and the direction A ′ overlap” corresponds to YES in FIG. 2, and “determine the point where the gamma ray is generated on each line in the direction A and the direction A ′. Is equivalent to NO in FIG.

  Next, the reason why the image quality is improved by performing event selection as shown in FIG. 2 and performing image reconstruction using the two-photon Compton camera image reconstruction apparatus 500 will be described. FIG. 6 is an explanatory diagram showing the reason why the image quality is improved. In addition, FIG.1 and FIG.2 is also referred as needed.

  In FIG. 6, when a positron 2002 having a relatively large initial momentum is released from the radiolabeled drug administered to the subject 150 and causes a pair annihilation reaction with the electron 2003 in the subject 150, a pair of annihilation occurs. Gamma rays 2004 and 2004 'are generated.

  Here, the upper diagram (a) of FIG. 6 shows an example in which the annihilation gamma ray 2004 (2004 ′) is detected by a conventional gamma ray detector 1001 (1001 ′) having only the coincidence counting determination function, The lower diagram (b) of FIG. 6 shows an example in which the annihilation gamma ray 2004 (2004 ′) is detected by the detector head 100 (100 ′) as shown in the present invention, that is, an example in which a Compton camera is used. And

  (A-1) and (b-1) of FIG. 6 are diagrams showing a case where the positron 2002 has caused a pair annihilation reaction with the electron 2003 in the subject 150 at the same location as the drug accumulation location 2000. Further, (a-2) and (b-2) in FIG. 6 show that the positron 2002 moves from the drug accumulation point 2000 to a position away from the distance 2001 and causes a pair annihilation reaction with the electron 2003 with a non-zero momentum. FIG. Furthermore, (a-3) and (b-3) in FIG. 6 show that after the positron 2002 has moved to a location away from the drug accumulation location 2000 by a distance of 2001 ′ and lost all kinetic energy (the momentum becomes zero). It is a figure which shows the case where a pair annihilation reaction is caused with the electron 2003. FIG.

  First, FIG. 6 (a-1) and FIG. 6 (b-1) are compared. In FIG. 6 (a-1) related to the prior art, the occurrence location of the gamma ray is estimated on the line 3001 shown in the figure, and the drug accumulation that becomes the actual occurrence location by the distance 3002 from the estimated location to the location where the pair annihilation reaction has occurred. Deviation from point 2000. This shift directly affects the resolution of the image, and as a result, the image quality deteriorates. However, in FIG. 6B-1 according to the present invention, there is no such influence, and the image quality does not deteriorate. That is, by imposing a determination based on the Compton scattering kinematics, the incident direction of the annihilation gamma ray 2004 (2004 ′) is estimated on the estimated direction 4002 (4002 ′) calculated from the Compton scattering kinematics as shown in FIG. Then, since the intersection of the estimated directions 4002 and 4002 ′ is estimated at the medicine accumulation location 2000 where the gamma ray is generated, the image quality is not deteriorated.

  Next, FIG. 6A-2 is compared with FIG. 6B-2. In FIG. 6 (a-2) according to the prior art, the location where a gamma ray is generated is estimated on the line 3001, and the distance 3002 ′ from the estimated location to the location where the pair annihilation reaction occurs and the pair annihilation from the drug accumulation location 2000 are shown. Only the distance 2001 to the place where the reaction has occurred is deviated from the actual drug accumulation place 2000. On the other hand, in FIG. 6B-2 according to the present invention, the location where the gamma rays are generated is estimated at the point of the distance 2001 from the drug accumulation location 2000 to the location where the annihilation reaction has occurred, and only the distance 2001 is actually generated. Since it does not deviate from the medicine accumulation location 2000 as a location, image quality degradation is reduced compared to the conventional case.

  Subsequently, FIG. 6A-3 is compared with FIG. 6B-3. In this case, the gamma ray generation location estimated for both is shifted from the actual drug accumulation location 2000 by a distance 2001 ′ from the drug accumulation location 2000 to the location where the pair annihilation reaction has occurred.

  From the above considerations, it can be seen that image quality is improved by performing image reconstruction using the two-photon Compton camera image reconstruction device 500.

  Next, event selection is performed as shown in FIG. 2, and image reconstruction is performed using the single photon Compton camera image reconstruction device 600, so that the coincident coincidence event can be efficiently removed, and the image quality and quantification of the image can be eliminated. The reason why the property is improved will be described. FIG. 7 is an explanatory diagram showing the reason why the image quality and quantitativeness of an image are improved. In addition, FIG.1 and FIG.2 is also referred as needed.

  Here, FIG. 7 shows a case where two pairs of annihilation gamma rays (2004 and 2004 ′ are a pair and 2014 and 2014 ′ are a pair) are counted simultaneously. FIG. 7A shows a case where measurement is performed by a conventional gamma ray detector 1001 (1001 ′) having only a coincidence determination function, and FIG. 7B shows the former detectors 101 and 101 of the present invention. It is assumed that measurement is performed with the configuration of 'and the subsequent detectors 102 and 102'. In the figure, reference numeral 2015 indicates scattered gamma rays and reference numeral 2016 recoil electrons, and the other reference numerals are the same as described above.

  In the case of FIG. 7A according to the prior art, the occurrence location of the gamma ray is estimated on the line 3001, and a false signal is generated in this event. Therefore, the image quality and quantitativeness of the image are deteriorated. On the other hand, in the case of FIG. 7B according to the present invention, the locations where gamma rays are generated are estimated on the lines 4002 and 4002 ′. Therefore, clearly, unlike the case of FIG. 7A, a true signal is generated without generating a false signal, and as a result, the image quality and quantitativeness of the image are improved.

  Subsequently, as shown in FIG. 2, event selection is performed, and image reconstruction is performed using the single photon Compton camera image reconstruction device 600, so that the in-vivo scattering event can be efficiently removed, and the image The reason why the image quality and the quantitativeness are improved will be described. FIG. 8 is an explanatory diagram showing the reason why the image quality and quantitativeness of an image are improved in the case of a subject internal scattering event. In addition, FIG.1 and FIG.2 is also referred as needed.

  Here, FIG. 8 shows a case where one of the pair of annihilation gamma rays 2004 and 2004 ′ is scattered inside the subject 150 to become scattered rays 2025. 8A shows a case where measurement is performed by a conventional gamma ray detector 1001 (1001 ′) having only the coincidence counting determination function, and FIG. 8B shows the former detectors 101 and 101 of the present invention. It is assumed that measurement is performed with the configuration of 'and the subsequent detectors 102 and 102'. In the figure, reference numeral 2035 indicates scattered gamma rays, and other reference numerals are the same as described above.

  In the case of FIG. 8A according to the prior art, the occurrence location of the gamma ray is estimated on the line 3001, and a false signal is generated in this event. Therefore, the image quality and quantitativeness of the image are deteriorated. On the other hand, in the case of FIG. 8B according to the present invention, the occurrence location of the gamma ray is estimated on the line 4002. The scattered radiation 2025 is determined by this energy and is not an estimation target of the location where gamma rays are generated. Therefore, obviously, unlike the case of FIG. 8A, a true signal is generated without generating a false signal, and as a result, the image quality and quantitativeness of the image are improved.

<Second Embodiment>
In the first embodiment, an example in which the simultaneous determination unit 300 and the kinematic determination unit 400 based on Compton scattering kinematics are imposed to perform image reconstruction using the two-photon Compton camera image reconstruction device 500 is shown. However, in this embodiment, as shown in FIG. 9, a PET reconstruction apparatus 800 is used instead of the two-photon Compton camera image reconstruction apparatus 500. This PET reconstruction apparatus 800 is an apparatus for reconstructing an image that is used when a gamma ray generation location is estimated on a line connecting two detectors that have been counted simultaneously. The PET reconstruction apparatus 800 has an apparatus configuration that uses a known image reconstruction method of a PET apparatus.

  The kinematics determination means 400 in the second embodiment has a slightly different processing procedure from that of the first embodiment, and this point will be described below with reference to FIG. FIG. 10 is a flowchart of “kinematic determination” according to the kinematic determination means 400 of FIG.

  In FIG. 10, the kinematic determination means 400 in the second embodiment uses gamma rays based on Compton scattering kinematics from gamma ray data (measured data) collected by each detector head 100, 100 ′ (see FIG. 1). The direction A and the direction A ′ are calculated, and it is determined whether or not the angle formed by the direction A and the direction A ′ of the gamma ray is 180 degrees. If the angle between the direction A and the direction A ′ is 180 degrees, it is determined as “estimate the location where the gamma ray is generated on one line” as indicated by Yes in the figure, and the direction A and the direction A If the angle formed by ′ is not 180 degrees, it is determined as “estimate the location where the gamma rays are generated on the lines of the respective directions A and A ′” as indicated by No in the figure. The determination of “estimating the occurrence location of gamma rays on a line in one direction” corresponds to YES of the kinematics determination unit 400 in FIG. 9, and “estimating the occurrence location of gamma rays on the lines in the respective directions A and A ′. The determination of “Yes” corresponds to NO of the kinematics determination means 400 of FIG.

  Returning to FIG. 9, in the second embodiment, when the kinematics generation unit 400 estimates the location where the gamma ray is generated on a line in one direction, in other words, on the line connecting two detectors that are simultaneously counted. When the location where gamma rays are generated is estimated, image reconstruction using the PET reconstruction device 800 is performed. On the other hand, in the determination by the kinematics determination means 400, when the location where the gamma rays are generated is estimated on the lines of the respective directions A and A ′, the image reconstruction using the single photon Compton camera image reconstruction device 600 is performed. To implement.

  The second embodiment will be described in more detail. For example, in the kinematics determination unit 400, the gamma ray energy is set to 511 keV, the crossing angle is set to 180 degrees, and PET reconstruction is performed for an event (for example, see FIG. 11) that satisfies this condition. Image reconstruction using the device 800 is performed. For events that do not satisfy the conditions, image reconstruction is performed using the single-photon Compton camera image reconstruction device 600. In the second embodiment as described above, the coincidence coincidence event and the internal scattering event can be removed and the generation of the false signal can be suppressed. As a result, the image quality and quantitativeness of the image are improved.

<Third Embodiment>
As shown in FIG. 12, the third embodiment classifies events in more detail by a kinematic determination means 400 based on the kinematics of Compton scattering, and according to the classified events, a two-photon Compton camera image reconstruction device. Image reconstruction using 500, image reconstruction using the single photon Compton camera image reconstruction device 600, and image reconstruction using the PET reconstruction device 800 are performed.

  The kinematics determination means 400 in the third embodiment has a slightly different processing procedure from that in the first embodiment and the second embodiment, and this will be described below with reference to FIG. FIG. 13 is a flowchart of “kinematic determination” according to the kinematic determination means 400 of FIG.

  In FIG. 13, the kinematics determination means 400 in the third embodiment uses gamma rays based on Compton scattering kinematics from gamma ray data (from measurement data) collected by each detector head 100, 100 ′ (see FIG. 1). The direction A and the direction A ′ are calculated, and it is determined whether or not the angle formed by the direction A and the direction A ′ of the gamma ray is 180 degrees. When the angle between the direction A and the direction A ′ is 180 degrees, it is determined as “estimate the location where a gamma ray is generated on one line” as indicated by Yes in the drawing. On the other hand, if the angle formed by the direction A and the direction A ′ is not 180 degrees, the process proceeds to the process indicated by No in the figure, and the deviation from the angle formed by the direction A and the direction A ′ of the gamma ray from 180 degrees is set. It is determined whether it is within the range (angle range).

  When the deviation from 180 degrees of the angle between the direction A and the direction A ′ is within the set range, as indicated by Yes in the figure, “the point where the gamma ray generation points overlap (the overlap between the direction A and the direction A ′). It is determined as “estimated to the point)”. In the case where the deviation from the angle formed by the direction A and the direction A ′ from 180 degrees is not within the set range, as indicated by No in the figure, “Gamma ray generation location is estimated on the lines of the respective directions A and A ′. Is determined.

  The determination of “estimating the occurrence location of gamma rays on a line in one direction” corresponds to “180 degrees” of the kinematics determination unit 400 in FIG. 12, and the determination of “estimating the occurrence location of gamma rays to the overlapping point”. Corresponds to “within range” of the kinematics determination unit 400 of FIG. 12, and the determination of “estimating the location where the gamma rays are generated on the lines of the respective directions A and A ′” is performed by the kinematics determination unit 400 of FIG. Corresponds to “out of range”.

  Explaining the third embodiment in more detail, for example, the kinematics determination unit 400 sets a plurality of crossing angles. When the crossing angle is 180 degrees, image reconstruction using the PET reconstruction apparatus 800 is performed, and when the crossing angle is within the range of angular fluctuation estimated from the initial momentum of the positron, two-photon Compton camera image reconstruction If image reconstruction is performed using the apparatus 500, and if the image reconstruction is performed using the single photon Compton camera image reconstruction apparatus 600 when out of the range, it is possible to remove the coincidence coincidence event and the internal scattering event, and the false signal As a result, the image quality and quantitativeness of the image are improved.

<Fourth embodiment>
In the first embodiment, the case of a plurality of detector heads fixed around the subject has been described. On the other hand, in this embodiment, as shown in FIG. 14, an example of the case of the detector head 100 (100 ′) having a mechanism for rotating around the subject 150 is shown. In the present embodiment, the rotary track 5000 may be a circular track, an elliptical track, or a subject closest track. By arranging the object around the subject in this way, three-dimensional isotropic data can be collected quickly, which contributes to shortening of the photographing time.

<Fifth Embodiment>
In the first embodiment, the case of a plurality of detector heads fixed around the subject has been described. On the other hand, in this embodiment, as shown in FIGS. 15 and 16, an example in the case of the detector head 100 arranged in a ring around the subject 150 is shown. By arranging the object around the subject in this way, three-dimensional isotropic data can be collected quickly, which contributes to shortening of the photographing time.

  In addition, it goes without saying that the present invention can be variously modified without departing from the spirit of the present invention.

  Hereinafter, a supplementary explanation regarding the present invention will be given.

  In recent years, with the spread of positron tomography equipment (hereinafter referred to as PET equipment) and FDG (radiopharmaceuticals that accumulate specifically in cancer cells due to active sugar metabolism of cancer cells), expectations for nuclear medicine diagnosis have increased greatly. . In addition, with the rapid progress of molecular biology and cell biology, a new field, molecular imaging, has emerged in which radioisotope-labeled compounds are widely used to elucidate the pathology and pharmacology from a basic scientific perspective. . Development of a high-performance imaging device that can cope with such diversified diagnosis, clinical experiment, and animal experiment is demanded.

  As various types of drugs are developed, there is an increasing need for using these drugs simultaneously and observing the pharmacokinetics for diagnosis or research. Actually, in the nuclear medicine diagnosis of the myocardium using a gamma camera (including a SPECT device), there is a diagnostic method using two kinds of drugs simultaneously.

  In the PET apparatus, in principle, it is not possible to select and measure gamma rays simultaneously emitted from various radiopharmaceuticals. The energy of a pair of opposed gamma rays generated as a result of positron decay (a positron causes a pair annihilation reaction with an electron in the material) is constant regardless of the type of radioisotope. In the PET apparatus, image reconstruction is performed based on straight line information obtained from a pair of opposed gamma rays. Therefore, the PET apparatus cannot obtain an image in which different types of radioisotopes are distinguished by gamma ray energy.

  A gamma camera (including a SPECT device) performs image reconstruction based on information on gamma rays that have linearly passed through the holes of the mechanical collimator. Therefore, different radioisotopes can be distinguished by the energy of gamma rays. However, since a gamma camera uses a mechanical collimator, for example, when using a radiopharmaceutical for PET, a thick collimator is required. As a result, spatial resolution or detection efficiency is significantly reduced. Generally, in a gamma camera, the spatial resolution deteriorates as the energy of gamma rays increases.

  The present invention is useful because it is a device capable of using both existing PET preparations and SPECT preparations and capable of dealing with new radiopharmaceuticals in terms of detection principle.

It is a block diagram which shows the apparatus structure of the nuclear medicine diagnostic apparatus of this invention in the state including a subject. FIG. 2 is a block diagram relating to data processing performed by the signal processing device and the image reconstruction device in FIG. 1. It is explanatory drawing about the reason which the quantitative property of an image improves. It is a flowchart of "simultaneous determination" concerning the simultaneous determination means of FIG. 3 is a flowchart of “kinematic determination” according to the kinematic determination means of FIG. 2. It is explanatory drawing about the reason which the image quality of an image improves. It is explanatory drawing about the reason for which the image quality and quantitative property of an image improve. It is explanatory drawing about the reason why the image quality and quantitative property of an image improve in the case of a subject internal scattering event. It is a figure which concerns on 2nd Embodiment, and is a block diagram which concerns on description of the data processing performed with a signal processing apparatus and an image reconstruction apparatus. 10 is a flowchart of “kinematic determination” according to the kinematic determination means of FIG. 9. It is a figure which shows an example of the event which satisfy | filled the conditions in 2nd Embodiment. It is a figure which concerns on 3rd Embodiment, and is a block diagram which concerns on description of the data processing performed with a signal processing apparatus and an image reconstruction apparatus. 13 is a flowchart of “kinematic determination” according to the kinematic determination means of FIG. 12. It is a figure which concerns on 4th Embodiment, and is a figure which shows the example about the case of the detector head provided with the mechanism which rotates the circumference | surroundings of a subject. It is a figure which concerns on 5th Embodiment, and is a figure which shows the example about the case of the detector head arrange | positioned circularly around the subject. It is a figure which concerns on 5th Embodiment, and is a figure which shows the example about the case of the detector head arrange | positioned circularly around the subject.

Explanation of symbols

100, 100 'detector (detection means)
101, 101 ′ Pre-stage detector 102, 102 ′ Sub-stage detector 110 Signal processing device (signal processing means)
120 Image reconstruction device (image reconstruction means)
DESCRIPTION OF SYMBOLS 130 Image display apparatus 150 Subject 300 Simultaneous determination means 400 Kinematic determination means 500 Two-photon Compton camera image reconstruction apparatus 600 Single-photon Compton camera image reconstruction apparatus 700 Image 800 PET reconstruction apparatus 2000 Drug accumulation location 2002 Positron 2003 Electron 2004, 2004 'annihilation gamma ray 2005, 2005' scattered gamma ray 2006, 2006 'recoil electron 4001 intersection 4002, 4002' estimated direction calculated from Compton scattering kinematics

Claims (1)

A nuclear medicine diagnostic apparatus characterized by calculating a generation location of photons emitted from a subject using Compton scattering kinematics and performing different image reconstruction according to the calculated information.

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