JP2009257962A - Radiation detector and positron emission tomography apparatus having it - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiation detector that suppresses layer multiplication as much as possible and has high spatial resolution, and to provide a PET device having it. <P>SOLUTION: This radiation detector 1 comprises a scintillator 2 formed of a first scintillator crystal layer (m) and a second scintillator crystal layer (n), and a PMT group 3 that is disposed on the lower surface of the scintillator 2 and detects fluorescence emitted from the scintillator 2. When the thicknesses of both scintillator crystal layers (m) and (n) are compared with each other, the first scintillator crystal layer (m) is thinner than the second scintillator crystal layer (n). <P>COPYRIGHT: (C)2010,JPO&INPIT

Description

  The present invention relates to a radiation detector, and a positron emission tomography apparatus (hereinafter, also referred to as a PET (Positron Emission Tomography) apparatus) including the radiation detector, and in particular, a radiation detector capable of position discrimination in the depth direction, and The present invention relates to a PET apparatus including the same.

  A radiation detector used in a PET apparatus is an optically coupled scintillator that converts incident radiation into fluorescence and a photomultiplier tube (hereinafter referred to as PMT) that senses fluorescence emitted from the scintillator. This radiation detector is provided so as to surround the outer periphery of the gantry opening in the PET apparatus, and forms a detector ring. And the scintillator provided in each radiation detector is formed by arranging prismatic scintillator crystals extending in the radial direction of the gantry opening in a tile shape. In this way, each scintillator crystal extends in the radiation traveling direction, so that radiation such as photons can be detected with high sensitivity.

In order to perform an examination using a PET apparatus equipped with such a radiation detector, first, a radiopharmaceutical labeled with a positron emitting nuclide is injected and administered to a subject. This positron emitting nuclide decays β ⁺ in the subject and generates a positron. The positron immediately collides with the electron in the subject and disappears. At this time, a pair of photons (annihilation photon pair) traveling in opposite directions are generated. The PET apparatus obtains a PET tomographic image showing the radiopharmaceutical distribution in the subject by detecting this annihilation photon pair with a detector ring.

  By the way, since the scintillator crystal provided in the radiation detector has an elongated shape so as to increase the radiation detection sensitivity, conversely, the spatial resolution in the radial direction in the cross section of the gantry decreases during the PET image formation. There is a problem that. First, as shown in FIG. 13, if each of the annihilation photon pairs is incident on the scintillator crystals 100 a and 100 b substantially perpendicularly and converted into fluorescence, it can be seen that the annihilation photon pair is generated in the spatial region 101. In this case, the spatial resolution does not decrease. However, if each of the annihilation photon pairs is incident on the scintillator crystals 100 c and 100 d from an oblique direction and converted into fluorescence, the annihilation photon pair is assumed to have occurred in the spatial region 102. As can be seen by comparing the spatial region 101 and the spatial region 102 in the figure, when an annihilation photon pair generated at the inner peripheral end of the detector ring 103 is incident on the scintillator from an oblique direction, an annihilation photon pair is expected to be generated. The area to be processed becomes wider, and as a result, the spatial resolution of the PET apparatus is lowered.

  In order to solve such a problem, a radiation detector 50 that can acquire information in the depth direction has been developed conventionally. As shown in FIG. 12, in the conventional radiation detector 50, the elongated scintillator crystals 51a, 51b, and 51c are three-dimensionally arranged to form the scintillator 51. The conventional radiation detector 50 has a PMT 52 that receives fluorescence from the scintillator 51 (see, for example, Patent Document 1).

JP 2000-56023 A

  However, the conventional example having such a configuration has the following problems. The conventional scintillator is configured ignoring the characteristic that the fluorescence generated in the scintillator crystal is frequently generated on the side close to the radiation source (that is, the shallow region in the depth direction of the scintillator crystal). In other words, in the conventional technology, the scintillator is multilayered so that the fluorescence emission position can be discriminated in the depth direction, but the length of the scintillator crystal in the depth direction is made constant in each scintillator crystal layer. Therefore, there is a problem that the position discrimination ability of the fluorescence becomes insufficient on the side close to the radiation source where the frequency of fluorescence is frequently generated inside the scintillator.

  However, this does not mean that more scintillator crystal layers constituting the scintillator should be provided. This is because it is very difficult to form a small scintillator crystal with high accuracy. In addition, as the scintillator crystal layer is added, more scintillator crystals are required to construct the scintillator, so that the manufacturing cost of the PET apparatus is greatly increased.

  Furthermore, if the number of layers constituting the scintillator is increased, the number of scintillator crystals increases, so that the fluorescence generated inside the scintillator must pass through more scintillator crystals before entering the PMT. Fluorescence travels through the interface between scintillator crystals, but the intensity of light decreases each time. This leads to a decrease in the number of annihilation photon pairs that can be acquired by the detector ring, and a clear PET tomographic image cannot be obtained. In order to prevent this, there is no choice but to set a longer examination time with the PET apparatus or increase the radiopharmaceutical to be injected and administered to the subject.

  The present invention has been made in view of such circumstances, and an object of the present invention is to provide a radiation detector with excellent spatial resolution and a PET apparatus including the same while suppressing the multilayering of scintillators as much as possible. Is to provide.

In order to achieve such an object, the present invention has the following configuration.
That is, the radiation detector according to the first aspect of the present invention includes a scintillator that converts radiation emitted from a radiation source into fluorescence, and a photomultiplier tube that detects fluorescence from the scintillator, and the scintillator includes a scintillator crystal. Includes a first scintillator crystal layer and a second scintillator crystal layer that are two-dimensionally arranged, and the first scintillator crystal layer and the second scintillator crystal layer are stacked toward the photomultiplier tube. The first scintillator crystal layer and the second scintillator crystal layer have different fluorescence decay time constants, the first scintillator crystal layer is disposed closer to the radiation source than the second scintillator crystal layer, and the first scintillator The layer thickness of the crystal layer is characterized by being thinner than the layer thickness of the second scintillator crystal layer.

  [Operation / Effect] According to the radiation detector of the first aspect, a radiation detector with higher spatial resolution can be provided. That is, the scintillator included in the radiation detector according to the present invention has a multilayer structure. Specifically, the scintillator is provided on the side far from the radiation source and the first scintillator crystal layer provided on the side closer to the radiation source. And a second scintillator crystal layer. Moreover, the thickness of the first scintillator crystal layer is set to be thinner than that of the second scintillator crystal layer. Therefore, a scintillator crystal having a shorter depth is arranged on the side closer to the radiation source of the scintillator. As the scintillator crystal constituting the scintillator becomes finer, the position discrimination ability of the radiation detector is improved, so that the scintillator according to the present invention has enhanced position discrimination ability in the depth direction on the side closer to the radiation source. Will be. Furthermore, since the radiation incident on the scintillator is converted to fluorescence more frequently on the side near the radiation source of the scintillator, the position discrimination capability in the depth direction on the side near the radiation source is enhanced. A radiation detector suitable for position discrimination in the depth direction can be provided. For example, when this is applied to a PET apparatus, a PET apparatus with improved spatial resolution can be provided. In short, even if an annihilation photon pair is generated at the inner peripheral end of the detector ring and radiation is incident on the scintillator from an oblique direction, the first scintillator crystal layer is formed with a small thickness, so that the first If fluorescence is emitted from the scintillator crystal layer, it can be seen that the fluorescence is generated within the limited layer thickness, so that the position where the fluorescence is generated can be distinguished in more detail.

  In addition, according to the above configuration, even if the radiation passes through the first scintillator crystal layer, the second scintillator crystal layer provided on the side farther from the radiation source than the first scintillator crystal layer as viewed from the radiation source is fluorescent. Therefore, sufficient radiation detection sensitivity can be secured. The radiation that has passed through the first scintillator crystal layer then enters the second scintillator crystal layer. Since the thickness of the second scintillator crystal layer is sufficiently thick, the radiation is reliably converted into fluorescence by either the first scintillator crystal layer or the second scintillator crystal layer. That is, even if the thickness of the first scintillator crystal layer is reduced, the radiation detection sensitivity of the radiation detector does not decrease.

  Furthermore, according to said structure, a radiation detector can be manufactured more easily. When the scintillator crystal is bonded to the PMT, the first scintillator crystal layer and the second scintillator crystal layer are both colorless and transparent crystal layers and cannot be distinguished with the naked eye, so there is a risk that both scintillator crystal layers will be mistakenly mistakenly. There is. However, according to the above configuration, since the thicknesses of both scintillator crystal layers can be sufficiently distinguished for the manufacturer, the scintillator is not manufactured with the two scintillator crystal layers mixed up.

  Furthermore, according to said structure, there are few crystal layers which comprise a scintillator. That is, when the fluorescence generated in the first scintillator crystal layer on the side far from the PMT (that is, the side close to the radiation source) goes to the PMT, the interface between the scintillator crystals that pass through is small. Therefore, the PMT detects fluorescence with the light intensity of the fluorescence generated by the scintillator kept as much as possible, and the detection sensitivity of the radiation detector having the above configuration is high. And according to the said structure, the manufacturing cost of a scintillator can be suppressed.

  According to the above configuration, the fluorescence decay times of the first scintillator crystal layer and the second scintillator crystal layer are different from each other. Thus, by measuring the decay time when fluorescence is incident on the PMT, it is possible to distinguish which scintillator crystal layer is the emitted fluorescence.

  According to a second aspect of the present invention, in the radiation detector according to the first aspect, the conversion efficiency, which is the efficiency of converting the radiation in the first scintillator crystal layer into fluorescence, is equal to or higher than the conversion efficiency of the second scintillator. It is characterized by that.

  [Operation / Effect] According to the invention described in claim 2, a radiation detector with higher spatial resolution can be provided. If the first scintillator crystal layer is made of a material with higher conversion efficiency as in the present invention, the radiation incident on the scintillator is more converted to fluorescence in the first scintillator crystal layer than in the second scintillator crystal layer. Will be. As described above, since more radiation is converted into fluorescence in the first scintillator crystal layer of the scintillator, the resolution of the radiation detector provided with such a scintillator becomes higher.

  According to a third aspect of the present invention, there is provided a positron emission tomography apparatus according to the first aspect, wherein the radiation detector according to the first or second aspect is arranged in a ring shape to generate radiation detection data; Simultaneous counting means for performing simultaneous counting of detected data, LOR deriving means for deriving LOR (Line of Response) which is a line segment connecting two scintillator crystals designated by the simultaneous counting means, and radiation based on LOR And re-addressing means for performing further layering processing of the detected data.

  [Operation / Effect] According to the PET apparatus of the fourth aspect, the generation position of the annihilation photon pair can be derived more accurately. In the scintillator according to the present invention, the first scintillator crystal layer having a thin layer thickness is disposed on the side close to the radiation source, and the second scintillator crystal layer having a large layer thickness is disposed on the side far from the radiation source. The discrimination ability in the depth direction of the scintillator is enhanced on the side closer to the radiation source where photons are converted into fluorescence. Therefore, the LOR derived by the LOR deriving unit becomes more accurate, and as a result, a PET apparatus with higher spatial resolution can be provided. Even if an annihilation photon pair is generated near the outer periphery of the detector ring and radiation is incident on the scintillator from an oblique direction, the first scintillator crystal layer is formed thin because the first scintillator crystal layer is thin. If the fluorescence is emitted from the light, it can be seen that the fluorescence is generated within the limited layer thickness, so that the position where the fluorescence is generated can be distinguished in more detail.

  The present specification also discloses an invention relating to the following radiation detector and positron emission tomography apparatus.

(1) In the radiation detector according to any one of claims 1 to 3, the first scintillator crystal layer is made of Lu _{2 (1-X)} Y _2X SiO ₅ and the second scintillator crystal. A radiation detector characterized in that the layer is composed of Gd ₂ SiO ₅ .

According to the configuration as in (1), the conversion efficiency of the first scintillator crystal layer can be increased more reliably. That is, Lu _{2 (1-X)} Y _2X SiO ₅ (LYSO) has a particularly high specific gravity among scintillator crystals that detect radiation, and has a higher effective atomic number. Therefore, if Lu _{2 (1-X)} Y _2X SiO ₅ (LYSO) is arranged in the first scintillator crystal layer that is closer to the radiation source, more radiation is converted into fluorescence in the first scintillator crystal layer. As a result, the spatial resolution of the radiation detector is further improved.

  (2) In the radiation detector according to any one of claims 1 to 3, the thickness of the second scintillator crystal layer is 1.2 times to 10 times the thickness of the first scintillator crystal layer. The radiation detector characterized by being.

  According to the configuration as in (2), a radiation detector with higher spatial resolution can be provided. If the layer thickness of the first scintillator crystal layer is reduced, the structure on the side close to the radiation source in the scintillator becomes finer, but the dose of radiation converted into fluorescence in the first scintillator crystal layer decreases. As a result, the spatial resolution of the radiation detector may rather decrease. That is, a configuration in which the thickness of the first scintillator crystal layer is, for example, about 1 cm is a preferable setting from the viewpoint of improving the spatial resolution. Further, since the layer thickness of the second scintillator crystal layer is sufficient, even if the thickness of the first scintillator crystal layer is reduced, the radiation detection sensitivity does not decrease.

  (3) In the radiation detector according to any one of claims 1 to 3, a first scintillator crystal that forms the first scintillator crystal layer and a second scintillator crystal that forms the second scintillator crystal layer. A radiation detector characterized by being connected in a one-to-one correspondence.

  According to the configuration as in (3), a radiation detector with a simple PMT configuration can be provided. That is, even if the multi-anode PMT is not employed, the discrimination of the fluorescence generation position in the plane direction of each scintillator crystal layer can be performed by using four inexpensive PMTs.

  According to the radiation detector and the positron emission tomography apparatus including the radiation detector according to the present invention, a configuration with high spatial resolution can be provided without reducing the radiation detection sensitivity. That is, the scintillator included in the radiation detector according to the present invention has a multi-layer structure. Specifically, the scintillator is provided on the side far from the radiation source and the first scintillator crystal layer provided on the side closer to the radiation source. And a second scintillator crystal layer. Moreover, the thickness of the first scintillator crystal layer is set to be thinner than that of the second scintillator crystal layer. Therefore, a scintillator crystal having a short depth is arranged on the side of the scintillator close to the radiation source. As the scintillator crystal constituting the scintillator becomes finer, the position discrimination ability of the radiation detector is improved, so that the scintillator according to the present invention has enhanced position discrimination ability in the depth direction on the side closer to the radiation source. Will be. Furthermore, since the radiation incident on the scintillator is converted to fluorescence more frequently on the side near the radiation source of the scintillator, the position discrimination capability in the depth direction on the side near the radiation source is enhanced. A radiation detector suitable for position discrimination in the depth direction can be provided, and when this is applied to a PET apparatus, a PET apparatus with improved spatial resolution can be provided.

  Embodiments of a radiation detector according to the present invention and a positron emission tomography apparatus having the same will be described below with reference to the drawings.

  First, the configuration of the radiation detector according to the first embodiment will be described. FIG. 1 is a perspective view of the radiation detector according to the first embodiment. As shown in FIG. 1, the radiation detector 1 according to the first embodiment includes a scintillator 2 including a first scintillator crystal layer m and a second scintillator crystal layer n, and a fluorescence emitted from the scintillator 2. And a light guide 4 disposed at a position interposed between the scintillator 2 and the PMT group 3. The scintillator crystal layers m and n are optically coupled to form an interface R between the layers. Although not shown in FIG. 1, the scintillator 2 is provided with a reflecting plate 5 (see FIG. 3) extending in the depth direction.

The scintillator 2 is made of a scintillator crystal suitable for photon detection. That is, the first scintillator crystal layer m is composed of Lu _{2 (1-X)} Y _2X SiO ₅ (hereinafter referred to as LYSO) in which Ce is diffused, and the second scintillator crystal layer n is Gd ₂ in which Ce is diffused. It is composed of SiO ₅ (hereinafter referred to as GSO). The fluorescence decay time constants representing the fluorescence decay rates between the scintillator crystal layers m and n are different from each other. The outer surface of the scintillator 2 is covered with a reflection film (not shown).

  The first scintillator crystal layer m is disposed closer to the radiation source than the second scintillator crystal layer n, and the thickness of the first scintillator crystal layer m is thinner than the thickness of the second scintillator crystal layer n. . The first scintillator crystal layer m and the second scintillator crystal layer n have a configuration in which the block-like first scintillator crystals a and second scintillator crystals b are two-dimensionally arranged. Specifically, 90 scintillator crystals a and b are arranged in a matrix in which nine scintillator crystals a and b are arranged in the X direction and nine in the Y direction per one of the scintillator crystal layers m and n. Next, scintillator crystals a and b will be described. FIG. 2 is a diagram for explaining the coupling between the first scintillator crystal and the second scintillator crystal according to the first embodiment. As shown in FIG. 2, the first scintillator crystal a has an incident surface J for allowing photons to enter the upper surface, and a bottom surface K optically coupled to the second scintillator crystal b on the lower surface on the opposite side. Have On the other hand, the second scintillator crystal b has a light guide coupling surface Q that is optically coupled to the light guide 4 on the upper surface P coupled to the above-described bottom surface K and the lower surface on the opposite side. Further, the surfaces J, K, P, and Q are all substantially the same shape. The interface between the light guide coupling surface Q and the light guide 4 is filled with grease such as silicone rubber. The photon corresponds to the radiation in the present invention.

  Next, the arrangement of the first scintillator crystal a and the second scintillator crystal b will be described. The bottom surface J of the first scintillator crystal a having substantially the same shape and the top surface P of the second scintillator crystal b are joined together so that their four sides coincide. That is, each first scintillator crystal a and each second scintillator crystal b are coupled in a one-to-one correspondence. Note that the interface R formed by the bottom surface J and the top surface P may be filled with grease such as silicone rubber, or an air layer may be interposed.

  Further, as described above, when the layer thicknesses of both scintillator crystal layers m and n are compared, the first scintillator crystal layer m is thinner than the second scintillator crystal layer n. That is, the thickness of the first scintillator crystal layer m is set to 1 cm, for example. On the other hand, the thickness of the second scintillator crystal layer n is set to 1.2 cm to 10 cm. In the present invention, the layer thickness direction is the same as the depth direction Z of the first scintillator crystal a and the second scintillator crystal b. When the total thickness of both scintillator crystal layers m and n is 3 cm, the thickness of the first scintillator crystal layer m is set between 13.6 mm and 2.7 mm, and the second scintillator crystal layer The thickness of n is set between 16.4 mm and 27.3 mm.

  The PMT group 3 is formed by arranging four PMTs 3a, 3b, 3c, and 3d in a matrix form, two in the X direction and two in the Y direction. The PMT group 3 is optically coupled to a light guide 4 that transmits fluorescence. Therefore, the fluorescence emitted from both scintillator crystal layers m and n enters the PMT group 3 through the light guide 4. The PMTs 3a, 3b, 3c, and 3d detect the intensity of fluorescence and the decay time. Each of the PMTs 3a, 3b, 3c, and 3d does not necessarily need a function of discriminating the position with respect to the X direction and the Y direction that are the arrangement directions of the PMTs 3a, 3b, 3c, and 3d. The PMTs 3a, 3b, 3c, and 3d correspond to the photomultiplier tube of the present invention.

  Next, the reflecting plate 5 will be described. FIG. 3 is a diagram illustrating the configuration of the reflector according to the first embodiment. The first scintillator crystals a are arranged in a matrix and form a first scintillator crystal layer m. Between the first scintillator crystals a adjacent to each other, a reflecting plate 5 that reflects fluorescence is provided. As shown in FIG. 3, the reflecting plate 5 has a flat plate shape extending along the depth direction Z of the scintillator 2. The reflector 5 is provided across both scintillator crystal layers m and n. The scintillator 2 has a plurality of reflection plates 5 and extends in the direction of the light guide 4 when viewed as a whole of the scintillator 2 to form a reflection plate lattice 6. Moreover, the length L of the reflecting plate 5 extending in the direction of the light guide 4 varies depending on the scintillator 2 portion. That is, the length L is set shorter as it goes from the corner of the scintillator 2 to the center. Therefore, when the reflector grating 6 is viewed from the light guide 4 side, the scintillator 2 has a structure that is recessed in the depth direction Z as if a spherical body can be held.

  Next, a method for discriminating the fluorescence emission positions in the X, Y, and Z (depth) directions of the radiation detector 1 according to the first embodiment will be described. Photons incident on the scintillator 2 are converted into fluorescence having a wavelength of 440 nm to 420 nm in either of the scintillator crystal layers m and n forming the scintillator 2. Fluorescence emitted from one of the scintillator crystals a and b enters the light guide 4 while being reflected by the reflection plate 5 described above. The reflector 5 has a length L set shorter as it goes from the corner of the scintillator 2 to the center, so that when the fluorescent light is emitted from the center of the scintillator 2, the reflector 5 obstructs the progress. Without going to PMT group 3. Therefore, the fluorescence spreads in the XY direction from the crystal in which it is generated, becomes wide light, and is detected by the four PMTs 3a, 3b, 3c, and 3d. On the other hand, when fluorescence is generated at the corners of the scintillator 2, the light travels toward the PMT group 3 while being repeatedly reflected by the reflector 5. As a result, the fluorescence is hardly spread and is detected by any one of the four PMTs 3a, 3b, 3c, and 3d constituting the PMT group 3. That is, in the radiation detector 1 according to the first embodiment, the distribution of the fluorescence to the four PMTs 3 a, 3 b, 3 c, and 3 d changes stepwise according to the fluorescence generation position in the scintillator 2. Using this, the emission position of the fluorescence in the X and Y directions in the scintillator 2 is calculated from the distribution ratio of the fluorescence to the PMTs 3a, 3b, 3c, and 3d.

  Discrimination of the fluorescence generation position in the Z direction utilizes the fact that the materials constituting the scintillator crystal layers m and n are different. In the scintillator 2 according to the first embodiment, LYSO is used for the first crystal layer 2m, and GSO is used for the second crystal layer 2n. Therefore, the fluorescence decay times are different between the scintillator crystal layers m and n. Therefore, if the decay time of fluorescence observed in the PMT group 3 is distinguished, it can be determined whether the fluorescence is emitted from the first scintillator crystal layer m or the second scintillator crystal layer n. Thus, the fluorescence emission position in the Z direction is discriminated. The decay time constant of fluorescence can be adjusted by changing the Ce concentration diffused in the scintillator crystals constituting both scintillator crystal layers m and n.

  As described above, the radiation detector 1 according to the first embodiment can provide a configuration with high spatial resolution without reducing the photon detection sensitivity. That is, the scintillator 2 included in the radiation detector 1 according to the first embodiment includes a first scintillator crystal layer m provided on the side close to the radiation source and a second scintillator crystal layer n provided on the side far from the radiation source. It has become. Moreover, the thickness of the first scintillator crystal layer m is set to be thinner than that of the second scintillator crystal layer n. Therefore, the position discrimination capability in the depth direction Z is enhanced on the side closer to the photon source that generates fluorescence more frequently. Therefore, a radiation detector suitable for position discrimination in the depth direction can be provided.

  Moreover, according to Example 1, since the 1st scintillator crystal layer m is comprised by LYSO with high conversion efficiency to the light of a photon, it can increase the fluorescence emitted by the 1st scintillator crystal layer m, A radiation detector capable of accurately discriminating the position in the depth direction Z can be provided. The second scintillator crystal layer n is formed of GSO. Thereby, a cheaper radiation detector can be provided. Further, the second scintillator crystal layer n does not contain Lu. Thereby, the dose of self-radiation generated from Lu and its influence can be suppressed as much as possible.

Subsequently, a positron emission tomography apparatus (hereinafter referred to as a PET apparatus) including the radiation detector described in the first embodiment will be described. FIG. 4 is a block diagram illustrating the configuration of the PET apparatus according to the second embodiment. As shown in FIG. 4, the PET apparatus 10 according to the second embodiment includes a gantry 11, a detector ring 12 provided inside the gantry 11, and ¹³⁷ Cs provided on the inner surface side of the detector ring 12. A photon point source 13 for irradiating a photon fan beam, a photon point source driving unit 14 for driving the photon fan beam, a bed 16 having a top plate 15 on which the subject M is placed, and a ceiling on which the top plate 15 is slid. A plate driving unit 17. The photon point source drive unit 14 is controlled according to the photon point source control unit 18, and the top plate drive unit 17 is controlled according to the top plate control unit 19. The PET apparatus 10 is further provided with various units for acquiring a tomographic image of the subject M. Specifically, the PET apparatus 10 receives a photon detection signal representing a detection position, detection intensity, and detection time of a photon from the detector ring 12 and simultaneously performs a coincidence counting unit 20 that simultaneously counts annihilation photon pairs. Referring to the incidence direction of the annihilation photon pair derived by the LOR deriving unit 21 and the LOR deriving unit 21 that specifies the incidence direction of the annihilation photon pair from the two photon detection data recognized as the annihilation photon pair by the counting unit 20 A readdress unit 22 is provided for further processing the position information of the scintillator 2 in the depth direction in the photon detection signal. In addition, the rear stage of the re-address unit 22 includes an absorption correction unit 23 that performs photon absorption correction with reference to transmission data to be described later, and an image forming unit 24 that forms a PET image of the subject M. The data transmitted from the detector ring 12 is packet data.

  The PET apparatus 10 according to the second embodiment further includes a main control unit 25 that controls the control units 18 and 19 in an integrated manner, and a display unit 26 that displays a PET image. The main control unit 25 is constituted by a CPU, and executes various programs to control the units 18 and 19, the coincidence counting unit 20, the LOR deriving unit 21, the readdressing unit 22, the absorption correcting unit 23, and the image formation. The unit 24 is realized. The coincidence counting unit 20, the LOR deriving unit 21, and the readdressing unit 22 correspond to the coincidence counting unit, the LOR deriving unit, and the readdressing unit in the present invention, respectively.

  The detector ring 12 is configured by arranging the radiation detectors according to the first embodiment in a block shape in a ring shape.

  Next, the LOR deriving unit 21 included in the PET apparatus 10 according to the second embodiment will be further described. LOR is an abbreviation of Line Of Responses, and means a line segment connecting two scintillator crystals on which an annihilation photon pair is incident. The paired photons are converted to fluorescence by the two scintillator crystals of the detector ring 12, so if the two scintillator crystals are connected by a straight line, the annihilation photon pair is generated from one of the straight lines. . Using this, the PET apparatus 10 according to the second embodiment knows the distribution information of the radiopharmaceutical of the subject M.

  A method for deriving LOR by the LOR deriving unit 21 will be described. FIG. 5 is a diagram for explaining a LOR derivation method according to the second embodiment. As an example, assume that fluorescence is emitted from the scintillator crystals 121 and 122. In order to easily obtain the LOR of the scintillator crystals 121 and 122, it is effective to use the surface where the fluorescence is emitted as the surface layer of the scintillator crystals 121 and 122.

  Prior to the derivation of the LOR, as shown in FIG. 5, a reference line connecting each of the centroids of the scintillator crystals 121 and 122 and ζ sub-LORs penetrating each of the scintillator crystals 121 and 122 parallel to the reference line are obtained. prepare.

  Next, the detection probability in each sub-LOR is obtained. The detection probability is an index of how much the photon is converted to fluorescence. The higher the probability, the more easily the photon incident by the scintillator crystal located on the sub-LOR is converted to fluorescence. As an example, the detection probability for the η th sub-LOR will be described. If a photon reaches the scintillator crystal 121 along the η-th sub-LOR, the photon follows three fate. First, it may be converted into fluorescence by the scintillator crystal 121. Second, before entering the scintillator crystal 121, the scintillator crystal 123 to the scintillator crystal 125 on the side close to the photon source (the subject M), etc. May be converted into fluorescence. Third, any scintillator crystal may penetrate and fly away from the detector ring 12. Here, considering the detection probability in the scintillator crystal 121, it is the probability that the first case will occur. Therefore, in order to obtain this detection probability, the scintillator crystal 123 or scintillator disposed on the front side (closer to the photon source) of the photon traveling direction as viewed from the scintillator crystal 121 from the probability that the entire scintillator converts photons into fluorescence. This means that the probability that the crystal 125 converts photons to fluorescence should be subtracted. This will be referred to as a detection probability related to the scintillator crystal 121. Further, the detection probability related to the scintillator 122 can be obtained in the same manner as the scintillator crystal 121. That is, the probability that the entire scintillator converts photons into fluorescence is subtracted from the probability that the scintillator crystal 126 arranged on the near side of the photon traveling direction (closer to the photon source) as viewed from the scintillator crystal 122 converts the photons into fluorescence. do it. From these facts, the probability that fluorescence is detected by the scintillator crystals 121 and 122 is the product of the detection probability related to the scintillator crystal 121 and the detection probability related to the scintillator crystal 122. The detection probability in the sub-LOR increases according to the length that the sub-LOR penetrates the scintillator crystal that emits fluorescence, and conversely, the traveling direction of the photons as seen from the scintillator crystal in which the sub-LOR emits fluorescence. It becomes small according to the length which penetrates the scintillator crystal arranged on the near side.

  Similarly, the detection probabilities in each sub-LOR from the first to the ζth are obtained in order. When comparing the detection probabilities, it is noticed that the detection probabilities are different from each other. That is, it can be said that there is an actual LOR in the vicinity of the sub-LOR having a higher detection probability, and the annihilation photon pair has entered the scintillator crystals 121 and 122 along the actual LOR. That is, if the detection probability is derived for each of the sub-LORs, and weighted and added according to the detection probability, the LOR in the scintillator crystals 121 and 122 can be obtained.

  The LOR obtained in this way is sent to the subsequent re-address unit 22. In the re-address unit 22, information processing for unifying the position information of the scintillator 2 in the depth direction is performed. (Such information processing is hereinafter referred to as a single layer processing.) In other words, if this single layer processing is performed, in the subsequent information processing, the structure of the scintillator 2 is virtually composed of one scintillator crystal layer. Thus, compression of position information in the depth direction in the scintillator 2 and high speed processing can be realized simultaneously. FIG. 6 is a diagram for explaining a single layer process in the PET apparatus according to the second embodiment. It is assumed that the annihilation photon pair is incident on the radiation detectors 131a and 131b constituting the detector ring 12. At this time, the radiation detector 131a discriminates the position where the fluorescence is emitted from the scintillator provided in the radiation detector 131a. Thus, the position in the X, Y, Z (depth) direction is discriminated, and the scintillator crystal 132a is specified as the fluorescence generation source. The radiation detector 131b is similarly discriminated for the position of the fluorescence. For example, the scintillator crystal 132a that is the first scintillator crystal and the scintillator crystal 132b that is the second scintillator crystal are respectively distinguished by the position discrimination. Is identified as the source of Based on this, the LOR deriving unit 21 derives an LOR connecting the scintillator crystals 132a and 132b. Note that the LOR deriving method has already been described.

  As in the example, when the second scintillator crystal emits fluorescence, the layering process in which the fluorescence generated in the second scintillator crystal is virtually emitted by the first scintillator crystal in the upper layer is readdressed. 22 is performed. FIG. 7 is a diagram for explaining a single layer process in the PET apparatus according to the second embodiment. First, as shown in FIG. 7, the re-address unit 22 acquires positional information of the scintillator crystal 132 c that is the first scintillator crystal through which the LOR obtained by the LOR deriving unit 21 passes. If this is overwritten on the position information of the scintillator crystal 132b which is the second scintillator crystal, the further layering process is completed. In FIG. 7, the stratification process is performed on the scintillator crystal 132b. However, when both of the two scintillator crystals that emit fluorescence are the second scintillator crystals, the re-address unit 22 further increases both of the scintillator crystals. Execute the conversion process. The photon detection data that has undergone the layering process is assembled into a PET image by the subsequent image forming unit 24.

  Next, the spatial resolution of the PET apparatus according to the second embodiment will be described. FIG. 8 is a diagram illustrating the spatial resolution of the PET apparatus according to the second embodiment. FIG. 8A shows a configuration of the scintillator in which the first scintillator crystal layer m and the second scintillator crystal layer n have the same thickness. Since the region 141 is located on the side closest to the radiation source, fluorescence is most easily generated. The region 142 is a region that does not reach the region 141 but has a relatively large amount of fluorescence when viewed as a whole scintillator. In the example of FIG. 8A, both regions 141 and 142 are disposed in the first scintillator crystal layer m. Therefore, the fluorescence generated in either of the regions 141 and 142 is not discriminated from which region. By the way, according to the above-described LOR derivation method, the LOR is further provided on the region 141 side. Therefore, when fluorescence is generated in the region 142, the LOR is shifted from the region 142, resulting in a decrease in spatial resolution. To do.

  On the other hand, as shown in FIG. 8B, the thickness of the first scintillator crystal layer m provided in the detector ring 12 according to the second embodiment is set to be thinner than that of the second scintillator crystal layer n. The Therefore, the region 142 now extends to the second scintillator crystal layer n. With this setting, the LOR is derived from the position of the second scintillator crystal layer n. Since the fluorescence generated in the region 142 is the largest in the second scintillator crystal layer n, the LOR derived by the LOR deriving unit 21 is provided on the region 142 side. Therefore, the fluorescence generated in the region 142 represents the emission position of the fluorescence more accurately without further separation from the LOR. As described above, in the scintillator according to the second embodiment, the position discrimination capability in the depth direction is further enhanced in the regions 141 and 142 on the side close to the radiation source.

  In addition, regarding the example 2, the inventors obtained the suitable conditions of the layer thickness of the first scintillator crystal layer m and the layer thickness of the second scintillator crystal layer n by simulation. More specifically, the central portion of the detector ring inside the detector ring and detection under a plurality of conditions in which the layer thickness of the first scintillator crystal layer m and the layer thickness of the second scintillator crystal layer n are changed. The spatial resolution of the PET device was compared at the outer periphery of the vessel ring. Each calculation result is obtained from the full width at half maximum (FWHM) of the spatial intensity distribution function of the fluorescence emitted by the scintillator.

  The full width at half maximum will be described. FIG. 9 is a diagram for explaining the full width at half maximum according to the second embodiment. The radiation detector 1 detects fluorescence in a wide area. The spatial intensity distribution function of the fluorescence is a function related to the fluorescence intensity I and the position, and represents the spatial spread of the fluorescence. More specifically, FIG. 9 schematically shows a spatial intensity distribution function of fluorescence as p and a fluorescence intensity I related to the position in the direction P in which the fluorescence spreads. That is, the fluorescence spreads with a predetermined distribution in the direction P. Specifically, the fluorescence intensity I is a maximum max at the position pc, and the fluorescence intensity I monotonously decreases as the position is farther from the pc.

  The full width at half maximum is a spread index for the position p of the fluorescence intensity I constituting the fluorescence. Specifically, first, max / 2 which is a half value of the maximum value max of the fluorescence intensity I in the fluorescence spatial intensity distribution function is obtained. Then, in the fluorescence spatial intensity distribution function, the two positions pa and pb corresponding to max / 2 are read, and the region H sandwiched between them is defined as the full width at half maximum.

  FIG. 10 is a diagram for explaining the simulation conditions in the second embodiment. In the simulation, the radius of the detector ring is 30 cm. Point E is 1 cm away from the center point O of the detector ring (distance D1 = 1 cm from center point O). The point F is 25 cm away from the center point O of the detector ring (distance D2 = 5 cm from the detector ring). The point F is an example of an inner peripheral side end of the detector ring 12 according to the second embodiment. The point E is included in the line segment OF. The extending direction of the line segment OF is defined as direction Q, and the axial direction of the detector ring 12 is defined as direction R.

  This simulation was performed on the assumption that the annihilation photon pair generated at the points E and F travels along the direction S orthogonal to the direction Q and the direction R and enters the detector ring 12. It is. The above-described direction P in which the fluorescence at the full width at half maximum spreads corresponds to the direction S in FIG. 10 and the axial direction R of the detector ring 12.

  Under the condition of a single scintillator crystal layer having a layer thickness of 30 mm (hereinafter referred to as condition 1), the spatial resolution decreases as the distance from the center of the detector ring increases. Note that E and F in the table correspond to point E and point F in FIG. 10, and E in the table is an example of the central portion of the detector ring according to Example 2, F is an example of an inner circumferential end of the detector ring according to the second embodiment.

  Next, in Table 1, the scintillator is a scintillator having a two-layer structure of a first scintillator crystal layer made of 15 mm LYSO and a second scintillator crystal layer made of 15 mm GSO (hereinafter referred to as condition 2). As compared with the case of Condition 1, a decrease in spatial resolution at the inner peripheral side end of the detector ring is suppressed. However, since the first scintillator crystal layer is thick, the spatial resolution at the outer peripheral portion of the detector ring is insufficient.

  Under the condition of a scintillator having a two-layer structure of a first scintillator crystal layer made of 10 mm LYSO and a second scintillator crystal layer made of 20 mm GSO (hereinafter referred to as condition 3), detection is further performed than condition 2. The reduction in spatial resolution at the inner peripheral end of the vessel ring is suppressed. The scintillator according to Example 2 adopting such conditions has enhanced position discrimination capability in the depth direction on the side close to the radiation source, and has a sufficiently high spatial resolution at the outer periphery. .

  Next, the operation of the PET apparatus according to the second embodiment will be described with reference to FIG. 4 again. In order to perform an examination with the PET apparatus 10 according to the second embodiment, first, the subject M, to which the radiopharmaceutical is injected and administered in advance, is laid on the top 16 and the top 16 is raised to a predetermined position. Then, the top 16 is slid to introduce the subject M into the gantry 11, and then transmission data indicating the photon absorption distribution inside the subject M is acquired. That is, a fan-like photon fan beam is irradiated from the photon point source 13 toward the subject M. This photon beam passes through the subject M and is detected by the detector ring 12. Then, while rotating the photon point source 13 along the inner peripheral surface of the detector ring 12, such detection is performed over the entire circumference of the subject M, and based on this, photon absorption in the cross section of the subject M is performed. Get the map. Then, the top plate 16 is slid again to sequentially change the position of the subject M, and the above-described acquisition of the photon absorption coefficient map is repeated each time. In this way, a photon absorption coefficient map of the entire head of the subject M is obtained.

  Subsequent to the acquisition of transmission data as described above, emission data for detecting annihilation photon pairs released from the radiopharmaceutical administered to the subject M is acquired. Prior to that, the photon point source 13 that has hindered the acquisition of the emission data is moved in the direction of the body axis of the subject M, and stored in a source shield not shown.

  Subsequently, emission data is acquired. That is, the annihilation photon pair whose traveling direction emitted from the inside of the subject M is opposite by 180 ° is detected by the detector ring 12. The photon detection signal detected by the detector ring 12 is sent to the coincidence counting unit 20, and only when two photons are detected at different positions on the detector ring 12 at the same time, one count is made and subsequent data processing is performed. It has come to be. Then, by continuing to acquire such emission data while sliding the top 16 and sequentially changing the position of the subject M, it is sufficient to image the internal distribution of the subject M of the radiopharmaceutical. Get the emission data of the count. Finally, the top plate 16 is slid again to move the subject M away from the gantry 11, and then the top plate 16 is lowered in order to move the subject M away from the top plate 16, and the examination is completed.

  Next, data processing in the PET apparatus according to the second embodiment will be described with reference to FIG. The transmission data and emission data output from the detector ring 12 have already been identified as to which scintillator crystal has been sensed. First, transmission data processing will be described. As shown in FIG. 4, the transmission detection data Tr is sent from the detector ring 12 to the readdress unit 22. Here, of the scintillators constituting the detector ring 12, information processing is performed to rewrite the position information so that the photon detection signal detected by the second scintillator layer is virtually detected by the first scintillator layer. More specifically, a line connecting the position of the scintillator crystal that is the source of fluorescence and the position of the photon point source 13 at the time of the occurrence of the fluorescence is provided, and the data detected by the second scintillator crystal The detection position is rewritten to the position of the first scintillator crystal through which this line passes. In this way, transmission data is acquired and sent to the absorption correction unit 23 at the subsequent stage.

  On the other hand, the emission detection data Em is output by the coincidence counting unit 20 to two photon detection signals sensed by the detector ring 12 within a predetermined time width by an annihilation photon pair and output to the LOR deriving unit 21. The LOR deriving unit 21 derives an LOR that connects two scintillator crystals that have detected the annihilation photon pair. The emission detection data Em is further layered by the re-address unit 22 based on the LOR, and is sent to the absorption correction unit 23 at the subsequent stage.

  In the absorption correction unit 23, the emission detection data Em is subjected to absorption correction that excludes the influence of the photon absorption distribution of the subject M superimposed on the emission detection data, based on the transmission detection data described above. In this way, detection data representing the radiopharmaceutical distribution in the subject M more accurately is sent to the image forming unit 24, where a PET image is reconstructed. Finally, it is displayed on the display unit 26.

  Next, a PET image under each condition shown in Table 1 is exemplified to emphasize the effect of the invention in Example 2. FIG. 11 is a diagram illustrating a PET image according to the second embodiment. These PET images are actual measurement results obtained by actually assembling the above-described simulation configuration and conducting experiments. First, FIGS. 11 (a) and 11 (b) show a PET image under Condition 3 according to the configuration of Example 2. FIG. FIG. 11A is a PET image when an annihilation photon pair is present at point E in FIG. As shown to Fig.11 (a), the image which appears is dot-like, and the generation | occurrence | production position of fluorescence is discriminated correctly. When this is compared with FIG. 11 (b) showing a PET image when an annihilation photon pair exists at point F in FIG. 10, the point F is somewhat blurred, but there is no significant change. Therefore, according to the PET apparatus 10 according to the second embodiment, even if an annihilation photon pair is generated at the inner peripheral side end of the detector ring 12, the position where the fluorescence is generated can be discriminated faithfully.

  Further explanation will be given focusing on the case where an annihilation photon pair exists at this point F. FIG. 11C is a PET image under the condition 2 shown in Table 1, and FIG. 11D is a PET image under the condition 1 shown in Table 1. FIGS. 11C and 11D show PET images when an annihilation photon pair exists at the point F in FIG. Comparing each of FIGS. 11B, 11C, and 11D, it can be seen that Condition 3 according to Example 2 is a configuration that most faithfully discriminates the position where the fluorescence is generated. In particular, the image in FIG. 11D is elongated, and it can be seen that the image is considerably blurred.

  As described above, according to the above-described PET apparatus 10 according to the second embodiment, the generation position of the annihilation photon pair can be derived more accurately. In the scintillator according to Example 2, the first scintillator crystal layer m having a thin layer thickness is disposed on the side close to the photon beam source, and the second scintillator crystal layer n having a large layer thickness is disposed on the far side. In the entire scintillator, the discrimination capability in the depth direction of the scintillator is enhanced on the side closer to the photon beam source where photons are converted into fluorescence. Therefore, the LOR formed by this configuration becomes more accurate, and as a result, the PET apparatus 10 with higher resolution can be provided. Even if an annihilation photon pair is generated at the inner peripheral end of the detector ring 12 and a photon is incident on the scintillator from an oblique direction, the first scintillator crystal layer m is formed with a thin layer thickness. If fluorescence is emitted from the scintillator crystal layer m, it can be seen that the fluorescence is generated within the limited layer thickness, so that the position where the fluorescence is generated can be distinguished in more detail.

  The present invention is not limited to the above embodiments, and can be modified as follows.

  (1) In each embodiment described above, a multi-anode type photomultiplier tube can be used as the PMT. As a result, a radiation detector capable of more accurately discriminating the position in the XY directions and a PET apparatus including the radiation detector can be provided.

(2) In Example 2 mentioned above, it is good also as a structure using ⁶⁸ Ge- ⁶⁸ Ga which is a nuclide which discharge | releases a positron as the photon point source 13. FIG. In this case, the energy of the photons emitted from the photon point source 13 is the same as the energy of the annihilation photon pair emitted from the radiopharmaceutical labeled with the positron emitting nuclide injected into the subject. Therefore, according to this modification, a photon absorption coefficient map that more accurately reflects the in-subject absorption of the annihilation photon pair can be acquired from the transmission data.

  (3) In each of the above-described embodiments, the first scintillator crystal layer m is formed of LYSO, but may be formed of, for example, GSO. In this case, the conversion efficiency, which is the efficiency of converting the radiation into fluorescence, is approximately the same for both scintillator crystal layers m and n. In addition, it is possible to appropriately select materials constituting both scintillator crystal layers m and n.

1 is a perspective view of a radiation detector according to Embodiment 1. FIG. It is a figure explaining the coupling | bonding of the 1st scintillator crystal which concerns on Example 1, and a 2nd scintillator crystal. 6 is a diagram illustrating a configuration of a reflector according to Example 1. FIG. It is a block diagram explaining the structure of the PET apparatus which concerns on Example 2. FIG. FIG. 10 is a diagram for explaining a LOR derivation method according to the second embodiment. It is a figure explaining the single layer process in the PET apparatus which concerns on Example 2. FIG. It is a figure explaining the single layer process in the PET apparatus which concerns on Example 2. FIG. It is a figure explaining the spatial resolution of the PET apparatus in Example 2. FIG. It is a figure explaining the full width at half maximum in Example 2. FIG. It is a figure explaining the conditions of the simulation in Example 2. FIG. FIG. 6 is a diagram showing a PET image according to Example 2. It is a figure explaining the structure of the conventional radiation detector. It is a figure explaining the fall of the spatial resolution in the conventional positron emission tomography apparatus.

Explanation of symbols

m ... first scintillator crystal layer n ... second scintillator crystal layer 2 ... scintillators 3a, 3b, 3c, 3d ... PMT (photomultiplier tube)
12 ... detector ring 20 ... coincidence counting unit (simultaneous counting means)
21 ... LOR deriving section (LOR deriving means)
22 ... Re-address part (re-address means)

Claims (3)

A scintillator that converts the radiation emitted from the radiation source into fluorescence;
Having a photomultiplier tube for detecting fluorescence from the scintillator;
The scintillator includes a first scintillator crystal layer and a second scintillator crystal layer configured by arranging scintillator crystals two-dimensionally,
The first scintillator crystal layer and the second scintillator crystal layer are laminated toward the photomultiplier tube,
The fluorescence decay time constants of the first scintillator crystal layer and the second scintillator crystal layer are different from each other,
The first scintillator crystal layer is disposed closer to the radiation source than the second scintillator crystal layer, and the layer thickness of the first scintillator crystal layer is smaller than the layer thickness of the second scintillator crystal layer. Characteristic radiation detector.   The radiation detector according to claim 1, wherein a conversion efficiency, which is an efficiency of converting the radiation in the first scintillator crystal layer into fluorescence, is equal to or higher than the conversion efficiency of the second scintillator. vessel. A detector ring in which the radiation detector according to claim 1 or 2 is arranged in a ring shape to generate radiation detection data;
Coincidence counting means for simultaneously counting the radiation detection data;
LOR deriving means for deriving LOR that is a line segment connecting two scintillator crystals designated by the coincidence means with a straight line;
A positron emission tomography apparatus comprising: a re-addressing unit configured to further process radiation detection data based on the LOR.