JP2010101682A - Nuclear medicine diagnosis apparatus - Google Patents

Abstract

Provided is a low-cost nuclear medicine diagnostic apparatus. A detector ring includes a plurality of detectors arranged on a circumference around a subject. Each of the plurality of detectors 12 receives at least one scintillator 14 that receives annihilation gamma rays emitted from the radioisotope in the subject P and generates scintillation light, and the scintillator 14 in the direction of the central axis A of the detector ring 1. And at least two photomultiplier tubes 16 that receive the scintillation light and generate electrical signals. The scintillator 14 has a strip shape and is arranged so that its longitudinal direction coincides with the direction of the central axis A. The scintillator 14 can secure a wide imaging field of view FOV in the direction of the central axis A without having to arrange a plurality of detector 1 rings in the direction of the central axis A.
[Selection] Figure 2

Description

  The present invention relates to a nuclear medicine diagnosis apparatus of the PET (Positron Emission Tomography) type.

  In a PET-type nuclear medicine diagnostic apparatus that is currently in widespread use, a plurality of detector rings in which a plurality of detectors are arranged circumferentially are arranged in a direction substantially perpendicular to the body axis of the subject. . Each detector includes a plurality of scintillators that convert gamma rays into scintillation light, and a plurality of photomultiplier tubes that amplify the scintillation light and convert it into an electrical signal. The photomultiplier tube is attached to one end of the scintillator in the radial direction orthogonal to the central axis direction of the ring. Thus, a PET-type nuclear medicine diagnostic apparatus requires a large number of photomultiplier tubes. Therefore, the manufacturing cost of the PET apparatus becomes expensive.

  An object of the present invention is to provide a low-cost nuclear medicine diagnostic apparatus.

  A nuclear medicine diagnostic apparatus according to an aspect of the present invention includes a plurality of detectors arranged on a circumference around a subject, and each of the plurality of detectors includes a radioisotope in the subject. At least one scintillator that generates radiated γ-rays to generate scintillation light, and at least two scintillators that are provided at both ends of the scintillator in the direction of the central axis of the circumference, and that generates electrical signals by receiving the scintillation light A photomultiplier tube.

  According to the present invention, it is possible to provide a nuclear medicine diagnostic apparatus with a low manufacturing cost.

  Hereinafter, a PET-type nuclear medicine diagnostic apparatus according to an embodiment of the present invention will be described with reference to the drawings.

  First, the problems of the conventional PET-type nuclear medicine diagnostic apparatus pointed out by the present inventors will be described in detail.

  As is well known, radioisotopes include positron nuclides that emit positrons. The positron nuclide administered to the subject accumulates at a specific site in the subject according to its chemical properties. The positrons emitted from the positron nuclide combine with surrounding electrons (electrons) to generate a pair of annihilation gamma rays. Each pair of annihilation gamma rays has 511 keV and is emitted in directions opposite to each other by approximately 180 degrees. On the other hand, the nuclear medicine diagnosis apparatus has a plurality of detectors arranged in a ring shape around the subject. A pair of annihilation gamma rays generated from positron nuclides are incident on the pair of detectors almost simultaneously. The nuclear medicine diagnosis apparatus simultaneously measures a pair of annihilation gamma rays incident on the detector one after another within a predetermined time frame. The nuclear medicine diagnostic apparatus has a positron nuclide on a straight line (LOR: Line Of Response) connecting the incident positions of a pair of annihilation gamma rays in each of the pair of detectors only when a pair of annihilation gamma rays are simultaneously measured. I reckon. The nuclear medicine diagnosis apparatus collects output signals from a pair of detectors relating to LOR as projection data. In this way, the nuclear medicine diagnostic apparatus collects projection data for 360 degrees. The nuclear medicine diagnostic apparatus generates image data based on the collected 360-degree projection data set.

  FIG. 10 is a schematic cross-sectional view of a detector ring 100 in a conventional nuclear medicine diagnostic apparatus. FIG. 11 is a schematic longitudinal cross-sectional view (cross-section 11-11 in FIG. 10) of the conventional detector ring 100. FIG. As shown in FIGS. 10 and 11, the detector ring 100 includes a plurality of detectors 200 arranged on a circumference having the body axis of the subject P placed on the top plate 500 as the central axis A. Have. Each detector 200 includes a plurality of scintillators 300 and a plurality of photomultiplier tubes 400. The photomultiplier tube 400 is provided at one end of the scintillator 300 in the radial direction orthogonal to the central axis A of the detector ring 100. That is, in the detector ring 100, the plurality of scintillators 300 and the plurality of photomultiplier tubes 400 are arranged concentrically (concentric cylinder). As shown in FIG. 11, in this conventional nuclear medicine diagnostic apparatus, when a wide imaging field of view (FOV) is ensured with respect to the central axis A, it is necessary to arrange a plurality of detector rings 100 along the direction of the central axis A. there were. Therefore, since the conventional nuclear medicine diagnostic apparatus has many photomultiplier tubes 400, the manufacturing cost is high.

  The nuclear medicine diagnosis apparatus according to the present embodiment can ensure a wide imaging field of view (FOV) in the direction of the central axis A (subject's body axis) at a low manufacturing cost. Specifically, the detector ring of the nuclear medicine diagnosis apparatus according to this embodiment is composed of a plurality of detectors having a unique structure. Due to the unique structure, the nuclear medicine diagnosis apparatus according to the present embodiment can realize a wide imaging field of view in the central axis direction even when one detector ring is used. Accordingly, the number of photomultiplier tubes can be reduced as compared with the conventional structure. Therefore, the nuclear medicine diagnosis apparatus according to this embodiment can secure a wide imaging field of view in the central axis direction at a low manufacturing cost.

  Hereinafter, the structure of the detector ring and each detector in the nuclear medicine diagnostic apparatus according to the present embodiment will be described.

  FIG. 1 is a schematic cross-sectional view of a detector ring 1 included in a PET-type nuclear medicine diagnostic apparatus according to an embodiment of the present invention. FIG. 2 is a schematic longitudinal section (section 2-2 in FIG. 1) of the detector ring 1. As shown in FIGS. 1 and 2, the detector ring 1 includes a plurality of detectors 12 arranged on a circumference with the body axis of the subject P placed on the top 10 as the central axis A. Composed. The number of detectors 12 is better for increasing the spatial resolution, but typically several tens of detectors 12 are arranged. Further, any number of detector rings 10 may be provided for one nuclear medicine diagnostic apparatus. However, as will be described later, a wide imaging field of view in the central axis direction can be secured with one detector ring 10, so that the number of photomultiplier tubes 16 can be reduced with respect to one nuclear medicine diagnostic apparatus. One is preferably provided.

  Next, the structure of one detector 12 will be described. FIG. 3 is a schematic perspective view of the detector 12. FIG. 4 is a schematic XY sectional view of the detector 12. As shown in FIGS. 3 and 4, the detector 12 includes a plurality of scintillators 14 and a plurality of photomultiplier tubes 16.

  The scintillator 14 typically has a strip shape. The scintillator 14 is arranged so that its longitudinal direction is parallel to the direction of the central axis A of the detector ring 1. The plurality of scintillators 14 are arranged along the thickness direction. Here, the central axis A direction (longitudinal direction of the scintillator 14) is the X direction, the arrangement direction of the scintillators 14 (thickness direction of the scintillator 14) is the Y direction, and the direction orthogonal to the X direction and the Y direction (width of the scintillator 14). Direction) is defined in the Z direction. This XYZ orthogonal coordinate system is an intra-crystal coordinate system set in the scintillator crystal 14.

The length of the scintillator 14 in the X direction is determined according to the required length of the imaging field in the X direction. The length of the scintillator 14 in the X direction is typically an appropriate length in the direction of the central axis A of a group of detector rings 100 when several conventional detector rings 100 are arranged in the direction of the central axis A. It is. For example, the length of the scintillator 14 in the X direction is designed to be about 100 mm to 200 mm. The length of the scintillator 14 in the Y direction is determined according to the spatial resolution. The length of the scintillator 14 in the Z direction is determined based on the effective atomic number of the scintillator. The scintillator 14 is, for example, NaI (sodium iodide), BGO (bismuth acid germanate), LSO (a lutetium silicate added with a certain amount of cerium), LaBr ₃ : Ce, LYSO (mixture of LSO and yttrium silicate). Crystal) and other scintillator crystals. The scintillator 14 may be formed of plastic or the like instead of the scintillator crystal.

  The photomultiplier tubes 16 are provided at both ends of the scintillator 14 in the X direction. The scintillator 14 and the photomultiplier tube 16 are bonded with, for example, grease. In the following description, when the two photomultiplier tubes 16 bonded to both ends of the scintillator 14 are distinguished, the photomultiplier tube 161 bonded to the + X direction is referred to as the one bonded to the −X direction. This will be called a photomultiplier tube 162.

  A reflective material 18 is provided between the scintillators 14. Further, the reflector 18 is also attached to the surface of the scintillator 14 where the photomultiplier tube 16 is not bonded. The reflective material 18 is formed of an extremely thin thin film. The material of the reflector 18 is, for example, a Teflon (registered trademark) tape. The reflector 18 is provided so as not to leak scintillation light propagating through the scintillator 14 to the outside.

  The detector 12 shown in FIG. 3 and FIG. 4 has three scintillators 14 as an example. However, the present invention is not limited to this, and the number of scintillators 14 provided in one detector 12 may be three or more or three or less. For example, the detector 14 may be designed in a structure having one scintillator 14.

  As described above, the scintillator 14 generates scintillation light when a pair of annihilation gamma rays are incident. The generated scintillation light propagates through the scintillator 14 and reaches the photomultiplier tube 16. The amount of scintillation light reaching the photomultiplier tube 16 varies according to the type of scintillator and the distance from the light emission position of the scintillation light to the photomultiplier tube 16. That is, since the length of the scintillator 14 is constant, the amount of scintillation light reaching the photomultiplier tube 161 and the amount of scintillation light reaching the photomultiplier tube 162 according to the light emission position of the scintillation light in the scintillator 14 The ratio of varies. Specifically, the photomultiplier tube 16 having a short distance from the light emitting position receives a larger amount of scintillation light than the photomultiplier tube 16 having a long distance from the light emitting position. The photomultiplier tube 16 amplifies the reached scintillation light and generates an electric signal having an intensity proportional to the amount of light.

  Further, the time from when the scintillation light is generated until it reaches the photomultiplier tube 16 varies depending on the type of the scintillator 14 and the distance from the emission position of the scintillation light to the photomultiplier tube 16. . That is, the difference between the generation time of the electric signal by the photomultiplier tube 161 and the generation time of the electric signal by the photomultiplier tube 162 changes according to the light emission position of the scintillation light in the scintillator 14. Specifically, the photomultiplier tube 16 having a short distance from the light emitting position is longer than the photomultiplier tube 16 having a long distance from the light emitting position until the generation of an electrical signal after the scintillation light is generated. The time is short.

  As described above, to reconstruct an image, the LOR must be specified. For this purpose, it is necessary to specify the emission positions of scintillation light derived from a pair of annihilation gamma rays. Considering the XYZ in-crystal coordinate system, the light emission position is defined by the X coordinate and the Y coordinate.

  First, consider the Y coordinate of the light emission position. In the detector 12, a plurality of scintillators 14 that are thin with respect to the Y direction are closely arranged along the Y direction. Therefore, the Y coordinate of the light emission position can be specified with good spatial resolution by specifying the scintillator 14 that is the source of scintillation light or the photomultiplier tube 16 that is the source of electrical signals.

  Next, consider the X coordinate of the light emission position. The detector 12 is provided with only one scintillator 14 along the X direction. In addition, the scintillator 14 has a length of about 100 mm to 200 mm in the X direction and is very long. Therefore, the X coordinate cannot be specified with good spatial resolution only by identifying the scintillator 14 that is the source of the scintillation light or the photomultiplier tube 16 that is the source of the electrical signal.

  Therefore, the PET-type nuclear medicine diagnostic apparatus according to this embodiment uses at least one of two types of specifying methods unique to this embodiment in order to specify the X coordinate with a good spatial resolution. One is a method that uses the difference between the generation time of the electrical signal by the photomultiplier tube 161 and the generation time of the electrical signal by the photomultiplier tube 162. The other is a method that uses the ratio of the amount of scintillation light reaching the photomultiplier 161 and the amount of scintillation light reaching the photomultiplier 162.

  Hereinafter, the principle of the method for specifying the X coordinate of the emission position of the scintillation light will be described in detail with reference to FIG. FIG. 5 is a longitudinal sectional view of the detector ring 1. As shown in FIG. 5, it is assumed that the detector 12a is arranged on the + Z side and the detector 12b is arranged on the -Z side. A photomultiplier 161a is attached to the + X side of the scintillator 14a of the detector 12a, and a photomultiplier 162a is attached to the -X side. Further, a photomultiplier 161b is attached to the + X side of the scintillator 14b of the detector 12b, and a photomultiplier 162b is attached to the -X side. The length l in the X direction of the scintillator 14a and the scintillator 14b is the same.

  It is assumed that the annihilation gamma ray Ra and the annihilation gamma ray Rb are generated at the generation position P0. The annihilation gamma ray Ra and the annihilation gamma ray Rb travel in the air at the speed of light c. The annihilation gamma ray Ra and the annihilation gamma ray Rb are incident on the scintillator 14a and the scintillator 14b, respectively. The annihilation gamma ray Ra incident on the scintillator 14a interacts with the scintillator at the light emission position Pa at time ta, and scintillation light is generated. Similarly, scintillator light is generated in the scintillator 14b at the light emission position Pb at time tb. The scintillation light continues to be emitted during the fluorescence time determined according to the type of the scintillator 14 and the like.

  It is assumed that the time ta and the time tb are within a predetermined time frame (for example, about 6 ns to 18 ns) and are “simultaneous”. A straight line connecting the light emission position Pa and the light emission position Pb constitutes an LOR. That is, it can be considered that the generation position P0 of the annihilation gamma ray Ra and the annihilation gamma ray Rb exists somewhere on the LOR.

The scintillation light generated at the light emission position Pa propagates through the scintillator 14a at a speed v determined according to the refractive index of the scintillator, and enters the photomultiplier tube 161a and the photomultiplier tube 161b. When the photomultiplier tube 161a and the photomultiplier tube 161b receive the scintillation light, the photomultiplier tube 161a generates an electrical signal having an intensity corresponding to the amount of the received scintillation light. As described above, the difference between the generation time of the electrical signal in the photomultiplier tube 161a and the generation time of the electrical signal in the photomultiplier tube 162a depends on the time-of-flight difference (TOF information) of the scintillation light in the scintillator 14. Here, it is assumed that the light emission position Pa is located at a distance l1a from the photomultiplier tube 161a and a distance l2a from the photomultiplier tube 162a. Further, it is assumed that the photomultiplier tube 161a generates an electric signal at time ta1, and the photomultiplier tube 162a generates an electric signal at time ta2. Then, the following expressions (1) and (2) are established.
l = l1a + l2a (1)
l1a-l2a = v (ta1-ta2) (2)
Here, the length l of the scintillator and the velocity v of the scintillation light are known. Therefore, the distance l1a and the distance l2a can be calculated based on the difference between the electrical signal generation time ta1 in the photomultiplier tube 161a and the electrical signal generation time ta2 in the photomultiplier tube 162a. That is, the X coordinate of the light emission position Pa can be calculated. Under the same conditions, the X coordinate of the light emission position Pb in the detector 12b can also be calculated.

  The intensity of the electrical signal generated by the photomultiplier tube 16 has a certain mathematical correspondence with the amount of received scintillation light. This correspondence is determined according to the performance of the photomultiplier tube 16. The amount of scintillation light received by the photomultiplier tube 16 has a certain mathematical correspondence with the distance from the light emission position. This correspondence is determined according to the type of scintillator. That is, the ratio between the intensity of the electric signal in the photomultiplier tube 161a and the intensity of the electric signal in the photomultiplier tube 161b is determined according to the position of the light emission position Pa in the scintillator 14a. Therefore, the distance l1a and the distance l2a can be calculated based on the ratio between the electric signal intensity of the photomultiplier tube 161a and the electric signal intensity of the photomultiplier tube 162a. That is, the X coordinate of the light emission position Pa can be calculated.

  Next, the configuration of the nuclear medicine diagnostic apparatus according to the present embodiment that includes the detector ring 1 and performs image reconstruction using the above-described light emission position specifying method will be described. FIG. 6 is a diagram illustrating a configuration of the nuclear medicine diagnosis apparatus 20 according to the present embodiment. As shown in FIG. 6, the nuclear medicine diagnostic apparatus 20 includes a detector ring 1, a signal processing circuit 3, a simultaneous measurement unit 5, an image reconstruction unit 7, and a display unit 9.

  As described above, the detector ring 1 has a plurality (n) of detectors 12 arranged on a circumference centered on the central axis. The detector 12 electrically detects the incidence of the annihilation gamma ray by photoelectrically converting the scintillation light generated when the annihilation gamma ray enters the scintillator 14 by the photomultiplier tube 16. Specifically, each of the n detectors 12 includes a plurality of scintillators 14. A plurality of photomultiplier tubes 161 are bonded to one end of the plurality of scintillators 14 in the central axis direction, and a plurality of photomultiplier tubes 162 are bonded to the other end, respectively. When the photomultiplier tube 161 receives the scintillation light from the scintillator 14, the photomultiplier tube 161 outputs a first output signal, which is an electric signal having an intensity corresponding to the amount of the received scintillation light, to the signal processing unit 3. Similarly, when the photomultiplier 162 receives the scintillation light from the scintillator 14, it outputs a second output signal, which is an electric signal having an intensity corresponding to the amount of the received scintillation light, to the signal processing circuit 3. More specifically, the photomultiplier tube 16 integrates an electric signal generated during a predetermined minute time interval, and generates an output signal having an intensity corresponding to the integrated value at each minute interval to generate a signal processing circuit. 3 is output. The minute interval depends on the time resolution of the photomultiplier tube 161. The detector ring 1 is embedded in a housing (not shown).

  The signal processing circuit 3 includes the same number n of signal processing units 4-1 to 4-n as the number of detectors 12. Each signal processing unit 4-1 to 4-n is electrically connected to each detector 12-1 to 12-n in a one-to-one correspondence. Each of the n signal processing units 4-1 to 4-n generates gamma ray data by performing signal processing on the first output signal from the photomultiplier tube 161 and the second output signal from the photomultiplier tube 162. To do. The gamma ray data is data in which the energy data related to the energy value of the annihilation gamma ray incident on the scintillator 14 is associated with the emission position of the scintillation light generated by the incident annihilation gamma ray and the information related to the emission time of the scintillation light. The gamma ray data is output to the simultaneous measurement unit 5 by each signal processing unit 4.

  Hereinafter, details of each of the signal processing units 4-1 to 4-n will be described with reference to FIG. Since the signal processing units 4-1 to 4-n have the same configuration and function, the signal processing units 4-1 to 4-n will be collectively described as the signal processing unit 4 without being distinguished. As shown in FIG. 7, the signal processing unit 4 includes an energy calculation unit 41, a light emission position calculation unit 43, a light emission time calculation unit 45, and a gamma ray data generation unit 47.

  The energy calculation unit 41 adds the first output signal from the photomultiplier tube 161 and the second output signal from the photomultiplier tube 162. The scintillation light generated in the scintillator 14 is collected by two photomultiplier tubes 161 and 162. The added value of the first output signal output from the photomultiplier tube 161 and the second output signal output from the photomultiplier tube 162 depends on the light amount of the scintillation light. Represents energy. This added value is called energy data. The generated energy data is output to the gamma ray data generation unit 47 by the energy calculation unit 41.

  The light emission position calculation unit 43 specifies the emission position of the annihilation gamma ray based on the principle of the method for specifying the emission position of the annihilation gamma ray described above. As a method for specifying the X coordinate of the light emission position, as described above, the light emission position calculation unit 43 includes the generation time of the first output signal in the photomultiplier tube 161 and the generation time of the second output signal in the photomultiplier tube 162. Based on the difference, the X coordinate of the light emission position is calculated. For example, the light emission position calculation unit 43 has a clock generation circuit, and the generation time of each output signal can be specified by the clock signal from the clock generation circuit. When the generation time of each output signal is specified, the light emission position calculation unit 43 calculates the X coordinate of the light emission position based on each specified generation time. For example, the X coordinate of the light emission position is defined by the distance from the reference position in the X direction of the scintillator 14. The reference position is set, for example, at the center point in the X direction of the scintillator 14 or at one end of both ends.

  As another X coordinate specifying method, the light emission position calculation unit 43 determines the X of the light emission position based on the ratio between the output value of the output signal of the photomultiplier tube 161 and the output value of the output signal of the photomultiplier tube 162. Coordinates may be calculated. For example, the light emission position calculation unit 43 stores a table in which the output value of the first output signal and the output value of the second output signal are input and the X coordinate of the light emission position is output for each material of the scintillator 14. deep. When each output signal is received from each photomultiplier 161, 162, the light emission position calculation unit 43 inputs each output signal to the table and outputs the X coordinate of the light emission position.

  As another specifying method, the X coordinate of the light emission position may be specified by a method combining the above two X coordinate specifying methods.

  As a method for specifying the Y coordinate, the light emission position calculation unit 43 specifies the Y coordinate of the light emission position based on the position of the scintillator 14 that is the source of the scintillation light. The position of the scintillator 14 is defined by an identification number for uniquely specifying the photomultiplier tube 16 that is the output source of the output signal. Since the photomultiplier tube 161 or the photomultiplier tube 162 that is the output source of the output signal has a one-to-one correspondence with the scintillator 14, the photomultiplier tube 16 that is the output source of the output signal is specified. This is synonymous with the specification of the scintillator 14 that has generated the scintillation light.

  Data relating to the specified light emission position is output by the light emission position calculation unit 43 to the light emission time calculation unit 45 and the gamma ray data generation unit 47. Further, data relating to the generation time of the output signal is also output by the light emission position calculation unit 43 to the light emission time calculation unit 45.

  The light emission time calculation unit 45 calculates the light emission time of the scintillation light based on the light emission position calculated by the light emission position calculation unit 43 and the generation time of each output signal. Data regarding the emission time is output to the gamma ray data generation unit 47 by the emission time calculation unit 45.

  The gamma ray data generation unit 47 generates gamma ray data based on the energy data from the energy signal calculation unit 41, the light emission position data from the light emission position calculation unit 43, and the light emission time data from the light emission time calculation unit 45. The gamma ray data is data in which light emission position data and light emission time data are associated with energy data. The gamma ray data is output to the coincidence counting unit 5 by the gamma ray data generating unit 37.

  Above, the detailed description about the signal processing part 4 is complete | finished.

  Returning to FIG. 6, the simultaneous measurement unit 5, the image reconstruction unit 6, and the display unit 9 will be described.

  As shown in FIG. 6, the simultaneous measurement unit 5 collects gamma ray data sequentially supplied from the signal processing units 4-1 to 4 -n in real time during scanning, and selects a predetermined amount from the collected gamma ray data. Two gamma ray data that fall within the time frame are extracted repeatedly. Specifically, the simultaneous measurement unit 5 refers to the occurrence time information associated with the collected gamma ray data, and the two gamma rays whose difference in occurrence time is within a time frame (for example, about 6 ns to 18 ns). Extract data. The two extracted gamma ray data are derived from annihilation gamma rays detected “simultaneously” by the pair of detectors 12. When the pair of gamma ray data is extracted, the simultaneous measurement unit 5 calculates the incident direction of the pair of annihilation gamma rays based on the two light emission position information associated with each. The simultaneous measurement unit 5 associates light emission position data, light emission time data, and incident direction data with energy data of a pair of gamma ray data. This data is called projection data. The projection data is generated by the simultaneous measurement unit 5 every time two simultaneous gamma ray data are extracted. The projection data is output to the image reconstruction unit 7 by the simultaneous measurement unit 5.

  When the image reconstruction unit 7 collects 360 ° projection data sets from the simultaneous measurement unit 5, the image reconstruction unit 7 reconstructs image data related to the distribution of positron nuclides based on the collected projection data sets. The image data is output to the display unit 9 by the image reconstruction unit 7. In order to reduce the noise of the image data, the image reconstruction may be performed in consideration of the flight time difference (TOF information) to the pair of annihilation gamma ray detectors 12, that is, the light emission time difference between the pair of gamma ray data.

  The display unit 9 displays the image data from the image reconstruction unit 7.

  Thus, the nuclear medicine diagnosis apparatus according to this embodiment can be manufactured at a low manufacturing cost.

  In addition, the structure of the detector 14 is not limited only to the said embodiment, A various modification is possible. Hereinafter, in Modification 1 and Modification 2, modifications of the detector will be described.

(Modification 1)
The detector according to the first modification is a DOI detector intended to use DOI (Depth Of Interest) information. FIG. 8 is a perspective view showing the structure of the detector 50 according to the first modification. As shown in FIG. 8, the detector 50 includes a scintillator group 54 having a plurality of scintillators 52, and photomultiplier tubes 56 attached to both ends of the scintillator group 54 in the X direction. A predetermined number of scintillators 52 are arranged in each of the Y and Z directions. The photomultiplier tube 56 is bonded to the scintillator group 54 with grease or the like. A reflective material is provided between the scintillators 52. Further, a reflective material is also provided on the surface of the scintillator 52 where the photomultiplier tube 56 is not bonded. Each photomultiplier tube 56 receives the scintillation light generated in the scintillator 52 and generates an output signal having an intensity corresponding to the amount of the scintillation light.

(Modification 2)
The detector according to the modified example 2 has a structure for extracting scintillation light to the outside with a light guide. FIG. 9 is a perspective view showing the structure of the detector 60 according to the second modification. As shown in FIG. 9, the detector 60 includes a scintillator 62, light guides 64 attached to both ends of the scintillator in the X direction, and photomultiplier tubes 66 attached to the ride guides. The light guide 64 optically connects the scintillator 62 and the photomultiplier tube 66. As the material of the light guide 64, a plastic having excellent light transmittance is suitable for efficiently transmitting the scintillation light generated from the scintillator 62 to the photomultiplier tube 66.

  Note that the present invention is not limited to the above-described embodiment as it is, and can be embodied by modifying the constituent elements without departing from the scope of the invention in the implementation stage.

In addition, various inventions can be formed by appropriately combining a plurality of components disclosed in the embodiment. For example, some components may be deleted from all the components shown in the embodiment. Furthermore, constituent elements over different embodiments may be appropriately combined.

The cross-sectional view of the detector ring of the nuclear medicine diagnostic apparatus according to the embodiment of the present invention. FIG. 2 is a longitudinal sectional view of the detector ring of FIG. 1. The perspective view of the detector contained in the detector ring of FIG. The figure which shows the XY cross section of the detector of FIG. The figure for demonstrating the principle of the identification process of the light emission position of an annihilation gamma ray by the detector of FIG. The figure which shows the structure of a nuclear medicine diagnostic apparatus provided with the detector ring of FIG. The figure which shows the structure of the signal processing part of FIG. The perspective view of the detector concerning the modification 1 of this embodiment. The perspective view of the detector concerning the modification 2 of this embodiment. The cross-sectional view of the detection ring of the conventional nuclear medicine diagnostic apparatus. FIG. 11 is a longitudinal sectional view of a detector ring of the conventional nuclear medicine diagnosis apparatus of FIG. 10.

  DESCRIPTION OF SYMBOLS 1 ... Detector ring, 3 ... Signal processing circuit, 4 ... Signal processing part, 41 ... Energy calculation part, 43 ... Light emission position calculation part, 45 ... Light emission time calculation part, 47 ... Gamma ray data generation part, 5 ... Simultaneous measurement part , 7 ... Image reconstruction unit, 9 ... Display unit, 10 ... Top plate, 12 ... Detector, 14 ... Scintillator, 16 ... Photomultiplier tube, 18 ... Reflector

Claims (5)

Comprising a plurality of detectors arranged on a circumference around the subject;
Each of the plurality of detectors is
At least one scintillator that receives annihilation gamma rays emitted from radioisotopes in the subject and generates scintillation light;
At least two photomultiplier tubes that are provided at both ends of the scintillator with respect to the central axis direction of the circumference and generate electrical signals by receiving the scintillation light;
A nuclear medicine diagnostic apparatus.   2. The nuclear medicine diagnosis apparatus according to claim 1, wherein each of the at least one scintillator has a strip shape, and is arranged so that a longitudinal direction of the strip shape coincides with the central axis direction.   The nuclear medicine diagnosis apparatus according to claim 2, wherein the at least one scintillator is arranged along a thickness direction of the strip shape.   A calculation unit for calculating a light emission position of the scintillation light in the scintillator with respect to the central axis direction based on generation times of electric signals by the two photomultiplier tubes provided at both ends of the scintillator; The nuclear medicine diagnostic apparatus according to claim 1 provided.   A calculation unit that calculates a light emission position of the scintillation light in the scintillator with respect to the central axis direction based on intensity of electric signals from the two photomultiplier tubes provided at both ends of the scintillator; The nuclear medicine diagnostic apparatus according to claim 1 provided.

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