JP5024182B2 - Tomography equipment - Google Patents

Description

  The present invention relates to a tomographic apparatus, and more particularly to a technique for performing time correction.

  As a tomography apparatus, a nuclear medicine diagnosis apparatus, that is, an ECT (Emission Computed Tomography) apparatus will be taken as an example, and as a nuclear medicine diagnosis apparatus, a PET (Positron Emission Tomography) apparatus will be explained as an example. In addition, a DOI-PET apparatus will be described as an example of a device having a DOI detector for discriminating a depth-of-interaction light source position (DOI: Depth of Interaction) (for example, Patent Document 1). -3, see Non-Patent Documents 1-3). The PET device is configured to reconstruct a tomographic image of a subject only when a plurality of photons generated by the annihilation of positrons (Positrons), ie positrons, are detected and detected simultaneously by a plurality of detectors. ing.

  Specifically, a radiopharmaceutical containing a positron emitting nuclide is administered into a subject, and a 511 KeV pair annihilation photon released from the administered subject is detected by a group of detection elements (for example, scintillators). Detect with instrument. And if two detectors detect photons within a certain period of time, they are detected at the same time, and are counted as a pair of annihilation photons. Further, the point of occurrence of annihilation is identified as a straight line of the detected detector pair. To do. By accumulating such coincidence information and performing reconstruction processing, a positron emitting nuclide distribution image (ie, a tomographic image) is obtained.

  At this time, the time until the photon is detected is slightly different for each detector. Therefore, by correcting the time required for the detection (this time correction is also called “timing calibration”), Obtain accurate coincidence information. By performing such time correction (timing calibration) to obtain coincidence count information, a more accurate positron emission nuclide distribution image can be obtained.

Thus, in an apparatus that obtains light source information from position information of a detector, obtains an image from the information, and performs imaging (ie, tomography), the image quality is improved by obtaining detailed position information of the detector. It is improving. For example, a method using a multi-anode photomultiplier tube (PMT: Photo Multiplier Tube) obtained by subdividing the detector as a planar detector as in Patent Document 3 described above, as in Patent Document 2 described above. A method for adjusting the depth direction of a partition wall made of a light reflecting material or a light shielding material embedded in a light guide interposed between a scintillator and a photomultiplier tube, and a detection element as in Non-Patent Document 1 described above A method of multi-layering in the depth direction, multi-layering in the depth direction as described above, and further combining a plurality of detection elements having different attenuation times in the depth direction as in Patent Document 1 and Non-Patent Document 2 described above. There are techniques for constructing a detector. In addition, a method using time difference information (TOF: Time Of Flight) that limits the point of occurrence of annihilation photons has been developed.
JP 2000-56023 A Japanese Patent Laid-Open No. 2004-245592 JP 2006-522925 A Naoko Inama, “Chapter 4 DOI Measuring Device”, [online], National Institute of Radiological Sciences, Internet <URL: http://www.nirs.go.jp/usr /medical-imaging/en/study/nextgeneration-pet2000/ninadama.html> Eiichi Tanaka (Hamamatsu Photonics, former National Institute of Radiological Sciences), Atsushi Kanno (Akita Brain Research Institute), “MIT Virtual Museum PET text14”, [online], ACADEMIA, Internet <URL: http://www.ricoh.co.jp/net- messena / ACADEMIA / JAMIT / MITVM / PET / TANAKA04 / text14.html> A GSO depth of interaction detector for PET: IEEE Trans. Nucl.Sci., 45: 1078-1082, 1998.

However, such a conventional example has the following problems. First, when a crystal element containing a natural radioactive element such as LSO (that is, having self-radioactivity) is not used as a photon detection element as a detection element, an external radiation source (for example, ¹⁸ ) is used to perform timing calibration. F pool phantom) is required. When an external source is used, timing calibration can be performed with a single measurement for a large number of detector pairs. As described above, the timing calibration is an important adjustment for the PET apparatus in order to obtain more accurate data. However, since a lot of statistics of such adjustment data are required, an external radiation source of about 1 mCi may be made, and there is a lot of labor and exposure. Further, since it is preferable that the timing calibration has less scattered radiation, a point source is ideally ideal, but it is not easy to hold such a strong point source. In addition, when using a point source, in order to obtain correction values for all detector pairs, for example, as described in the following reference, it is necessary to prepare a reference detector outside. In this case, the accuracy of the location of the point source and reference detector is important and not a simple method (Reference: Nuclear Science Symposium Conference Record, 2004 IEEE Volume 4, Issue, 16-22 Oct. 2004 Page ( s): 2361-2365 Vol.4).

When a detection element containing a natural radioactive element ¹⁷⁶ Lu such as LSO is used as the photon detection element, a certain amount of Back Ground is further formed in the adjustment data. Therefore, it is possible to eliminate many of such random components by performing coincidence, but when performing timing calibration, in order to reduce the influence of self-radioactivity, some strong external radiation source It is preferable to use Use of such a strong external radiation source will increase the exposure of the adjuster.

  The present invention has been made in view of such circumstances, and an object thereof is to provide a tomographic apparatus capable of performing time correction without using an external radiation source.

In order to achieve such an object, the present invention has the following configuration.
That is, the invention described in claim 1 is a tomography apparatus including a detector including a crystal element having self-radiation, and includes time correction means for performing time correction using the self-radioactivity. It is characterized by.

  [Operation / Effect] When a crystal element having self-radiation is used as a detection element, Back Ground can be obtained as described above. We have found that this Back Ground can be used for time correction. Therefore, according to the first aspect of the present invention, the time correction can be performed without using an external radiation source by providing time correction means for performing time correction using self-radioactivity.

  In the above-described invention, it is preferable that an energy width setting unit for setting a photon energy width is provided, and the time correction unit performs time correction using self-radiation within the photon energy width set by the energy width setting unit ( Invention of Claim 2). Only photons having a desired energy width can be extracted by setting with the energy width setting means, and a time distribution (timing histogram) with a good S / N ratio can be obtained. As a result, more accurate time correction can be performed.

  In addition, in a preferable example of these inventions described above, the detector is a detector configured by stacking crystal elements in the depth direction in order to discriminate the light source position in the depth direction in which the interaction has occurred (that is, DOI). Detector) (the invention according to claim 3). By using such a detector, the resolution in the depth direction can be improved and time correction can be performed.

  In addition, in a preferable example of these inventions described above, the crystal element having self-radiation includes a nuclide that undergoes α or β-decay and emits γ rays accompanying the α or β-decay. Alternatively, a detector that has caused β-decay detects α-rays or β-rays, and another detector detects γ-rays associated with α- or β-decay, and α- or β-ray detection events and γ The time correction means performs time correction using the detection event of the line (the invention according to claim 4).

  In addition, the above-described detectors may all be composed of crystal elements having self-radiation, or may be composed of crystal elements having self-radiation and crystal elements not having self-radiation ( Invention of Claim 5). In the case of including a crystal element that does not have self-radiation, such as the latter, self-radiation is performed based on a detection event by a detector composed of a crystal element that has self-radiation and a crystal element that does not have self-radiation. The time correction means can perform time correction between detectors that do not have the ability (the invention according to claim 6). When a detector composed of crystal elements with self-radioactivity is used in combination with another detector, correction data between the detectors with self-radioactivity and one with self-radioactivity and the other with self-radioactivity Correction data for detectors that do not have an error can be obtained, and by using these correction data, it is possible to perform time correction between detectors that do not have self-radiation.

  According to the tomographic apparatus of the present invention, the time correction can be performed without using an external radiation source by providing the time correction means for performing the time correction using the self-radioactivity.

Embodiment 1 of the present invention will be described below with reference to the drawings.
FIG. 1 is a side view of the PET-CT apparatus according to the first embodiment, and FIG. 2 is a block diagram of the PET-CT apparatus according to the first embodiment. In the first embodiment, a nuclear medicine diagnosis apparatus will be described as an example of the tomography apparatus, and a PET-combined PET (Positron Emission Tomography) apparatus and an X-ray CT apparatus will be used as the nuclear medicine diagnosis apparatus. The CT apparatus will be described as an example.

  As shown in FIG. 1, the PET-CT apparatus 1 according to the first embodiment includes a top plate 2 on which a subject M in a horizontal posture is placed. The top plate 2 is configured to move up and down and translate along the body axis of the subject M. The PET-CT apparatus 1 includes a PET apparatus 3 and an X-ray CT apparatus 4 for diagnosing the subject M placed on the top 2. The PET-CT apparatus 1 corresponds to the tomography apparatus in this invention.

  The PET apparatus 3 includes a gantry 31 having an opening 31a and a photon detector 32 that detects photons generated from the subject M. The photon detector 32 is arranged in a ring shape so as to surround the body axis of the subject M, and is embedded in the gantry 31. The photon detector 32 includes a scintillator block 32a, a light guide 32b, and a photomultiplier tube (PMT) 32c (see FIG. 3). The scintillator block 32a includes a plurality of scintillators. Photons generated from the subject M to which the radiopharmaceutical is administered are converted into light by the scintillator block 32a, and the converted light is guided by the light guide 32b, and the photomultiplier tube 32c photoelectrically converts the electric signal. Output to. The photon detector 32 corresponds to the detector in the present invention. A specific configuration of the photon detector 32 will be described later with reference to FIG.

  On the other hand, the X-ray CT apparatus 4 includes a gantry 41 having an opening 41a. In the gantry 41, an X-ray tube 42 for irradiating the subject M with X-rays and an X-ray detector 43 for detecting X-rays transmitted through the subject M are disposed. The X-ray tube 42 and the X-ray detector 43 are arranged so as to face each other, and the X-ray tube 42 and the X-ray detector 43 are covered in the gantry 41 by driving a motor (not shown). Rotate around the body axis of the specimen M. In the first embodiment, a flat panel X-ray detector (FPD) is employed as the X-ray detector 43. Of course, an X-ray detector other than the flat panel X-ray detector (FPD) may be used.

  In FIG. 1A, the gantry 31 of the PET apparatus 3 and the gantry 41 of the X-ray CT apparatus 4 are separated from each other. However, as shown in FIG.

  Subsequently, a block diagram of the PET-CT apparatus 1 will be described. As shown in FIG. 2, the PET-CT apparatus 1 includes a console 6 in addition to the top plate 2, the PET apparatus 3, and the X-ray CT apparatus 4 described above. The PET apparatus 3 includes a PET data collection unit 5 in addition to the gantry 31 and the photon detector 32 described above. The X-ray CT apparatus 4 includes a CT data collection unit 44 in addition to the gantry 41, the X-ray tube 42, and the X-ray detector 43 described above.

  The PET data collection unit 5 collects PET data (nuclear medicine data) based on the photons detected by the photon detector 32, and calculates the time difference between the coincidence circuit 51, the amplifier 52, the AD converter 53, the energy window unit 54, and the like. An information calculation unit 55 and a timing calibration unit 56 are provided. The energy window unit 54 corresponds to the energy width setting unit in the present invention, and the timing calibration unit 56 corresponds to the time correction unit in the present invention.

  The console 6 includes a data collection unit 61, an image reconstruction unit 62, a memory unit 63, an input unit 64, an output unit 65, and a controller 66.

  The coincidence circuit 51 determines whether or not photons are simultaneously detected (that is, coincidence) by the photon detector 32. The PET data simultaneously counted by the coincidence counting circuit 51 is sent to the data collecting unit 61. The amplifier 52 amplifies the electric signal detected and output by the photon detector 32. The AD converter 53 converts the analog value of the electric signal amplified by the amplifier 52 into a digital value and outputs the digital value.

  The energy window unit 54 extracts only photons having a desired energy width from the photons detected by the photon detector 32. In the first embodiment, since simultaneous counting of β rays (maximum energy 580 KeV) and 307 KeV γ rays is required, β-ray energy window (energy width) EW_1: 400 to 800 KeV, γ-ray energy window ( Energy window) EW_2: The energy window unit 54 sets 200 to 400 KeV.

  The time difference information calculation unit 55 obtains the time difference between the detector pairs based on the photons having the desired energy width extracted in the energy window unit 54 (in the first embodiment, β rays and 307 KeV γ rays). (Timing Histogram) is obtained. The timing calibration unit 56 performs timing calibration (time correction) based on the time difference distribution obtained by the time difference information calculation unit 55. The coincidence counting circuit 51 performs coincidence counting of photons at timing after timing calibration.

  On the other hand, the CT data collection unit 44 collects projection data as CT data (data for X-ray CT) based on the X-rays detected by the X-ray detector 43. The CT data collected by the CT collection unit 44 is sent to the data collection unit 61.

  The data collection unit 61 superimposes the PET data collected by the coincidence counting by the coincidence circuit 51 and the CT data collected by the CT data collection unit 44. Further, the CT data collected by the CT data collection unit 44 may be made to act on the PET data as transmission data to perform absorption correction of the PET data. The data collection unit 61 sends the superimposed projection data to the image reconstruction unit 62. The image reconstruction unit 62 reconstructs the projection data superimposed by the data collection unit 61 to generate a tomographic image.

  The memory unit 63 is connected to the PET data collection unit 5, the CT data collection unit 44, and the data collection unit 61 via the controller 66 and data such as tomographic images reconstructed by the image reconstruction unit 62. Is written and stored, read out as necessary, and each data is sent to the output unit 65 via the controller 66 and output. The memory unit 63 is configured by a storage medium represented by ROM (Read-only Memory), RAM (Random-Access Memory), and the like.

  The input unit 64 sends data and commands input by the operator to the controller 66. The input unit 64 includes a pointing device represented by a mouse, a keyboard, a joystick, a trackball, a touch panel, and the like. The output unit 65 includes a display unit represented by a monitor, a printer, and the like.

  The controller 66 performs overall control of each part constituting the PET-CT apparatus 1 according to the first embodiment. The controller 66 includes a central processing unit (CPU). The data collected by the PET data collection unit 5, the CT data collection unit 44, the data collection unit 61, and the data such as the tomographic image reconstructed by the image reconstruction unit 62 are sent to the memory unit 63 via the controller 66. Is written and stored, or sent to the output unit 65 for output. When the output unit 65 is a display unit, output is displayed. When the output unit 65 is a printer, output printing is performed.

  Photons generated from the subject M to which the radiopharmaceutical is administered are converted into light by the scintillator block 32a (see FIG. 3) of the corresponding photon detector 32 among the photon detectors 32, and the converted light is photon. The photomultiplier tube 32c (see FIG. 3) of the detector 32 performs photoelectric conversion and outputs an electrical signal. The electric signal is sent to the amplifier 52 together with the coincidence circuit 51 as image information (pixel value).

  Specifically, when a radiopharmaceutical is administered to the subject M, the positron emission RI positron disappears and two photons are generated. The coincidence circuit 51 checks the position of the scintillator block 32a (see FIG. 3) of the photon detector 32 and the incident timing of the photon, and the photon is detected by the two scintillator blocks 32a that are opposed to each other with the subject M interposed therebetween. The image information sent in is determined to be appropriate data only when the images are incident simultaneously (that is, when simultaneous counting is performed). When a photon is incident only on one scintillator block 32a, the coincidence counting circuit 51 treats it as noise instead of a photon generated by annihilation of the positron, and determines that the image information sent at that time is also noise and rejects it. .

  When simultaneous counting is performed by the simultaneous counting circuit 51, if there is a difference between the detection time of one detector and the detection time of the other detector, simultaneous counting cannot be performed accurately. Therefore, in order to eliminate such a deviation, timing calibration is performed by the timing calibration unit 56 described above, and the timing of the detection time between the detectors is matched. Therefore, when the coincidence counting is performed by the coincidence counting circuit 51, it is performed at the timing after the timing calibration. Specific functions of the above-described time difference information calculation unit 55 and timing calibration unit 56 will be described in detail later.

  The image information sent to the coincidence circuit 51 is sent to the data collecting unit 61 as projection data (PET data). On the other hand, while rotating the X-ray tube 42 and the X-ray detector 43, the subject M is irradiated with X-rays from the X-ray tube 42, and the X-rays irradiated from the outside of the subject M and transmitted through the subject M are emitted. The X-ray detector 43 detects an X-ray by converting it into an electric signal. The electrical signal converted by the X-ray detector 43 is sent to the CT data collection unit 44 as image information (pixel value). The CT data collection unit 44 collects the distribution of the sent image information as projection data (CT data) projected on the projection plane of the X-ray detector 43 and sends it to the data collection unit 61.

  The data collection unit 61 performs absorption correction of PET data and superimposition of PET data and CT data, and sends them to the image reconstruction unit 62. The image reconstruction unit 62 reconstructs the projection data that has been sent in, and generates a tomographic image. Generate.

Next, a specific configuration of the photon detector 32 according to the first embodiment, including a second embodiment described later, will be described with reference to FIG. FIG. 3 is a schematic diagram of a specific configuration of the photon detector. FIG. 4 is a collapse view of ¹⁷⁶ Lu, and FIG. 5 is a schematic diagram showing a combination of detectors having the same line (LOR: Line Of Response) connecting the simultaneously counted detectors.

  The photon detector 32 includes a scintillator block 32a configured by combining a plurality of scintillators as detection elements having different attenuation times in the depth direction, a light guide 32b optically coupled to the scintillator block 32a, and a light guide 32b. And a photomultiplier tube 32c optically coupled to each other. Each scintillator in the scintillator block 32a detects γ-rays by emitting light by incident photons and converting it into light.

  In Example 1, including Example 2 described later, two scintillator groups having different attenuation times in the depth direction (r in FIG. 3) are combined, and the incident side with respect to the depth direction is combined. A scintillator group A and a scintillator group B in order. In this way, the light source position in the depth direction is discriminated by combining a plurality (two in FIG. 3) of scintillators that are detection elements having different decay times in the depth direction. Further, in this embodiment, a plurality of scintillators are combined with respect to a plane orthogonal to the depth direction, and the light source position is also distinguished from the plane orthogonal to the depth direction.

  In addition, in Example 1, including Example 2 described later, the scintillator group A on the incident side is formed of LYSO (LuYSiO), and the scintillator group B on the photomultiplier tube side is formed of GSO (GdSiO). Form. That is, in each embodiment, a two-layer DOI detector composed of scintillator group A: LYSO and scintillator group B: GSO will be described as an example. The LYSO forming the scintillator group A is a crystal element having self-radiation, and the GSO forming the scintillator group B is a crystal element having no self-radiation.

Lu has a natural abundance of 2.59% and has ¹⁷⁶ Lu (half-life of 37.8 billion years), and emits electrons due to β-decay (maximum energy of 580 KeV), and then immediately 307 KeV (94%), 202 KeV (86%), 88 KeV ( 13.3%) and other gamma rays are emitted. The collapse view is shown in FIG. “3.78 × 10 ¹⁰ y” in FIG. 4 indicates a half-life of 37.8 billion years, and represents a state in which energy transitions while emitting γ rays. In FIG. 4, γ rays are emitted at 306.78 KeV and 201.83 KeV, respectively. However, the following explanation will be made assuming that 306.78 KeV is set to 307 KeV by rounding off one decimal place, and 201.83 KeV is set to 202 KeV.

In addition, in Example 1, including Example 2 to be described later, ¹⁷⁶ Lu is taken as an example of a crystal element having self-radiation, but Pb, Tl which are natural radioactive elements other than ¹⁷⁶ Lu are described. , Bi, K, lanthanoid elements, actinoid elements, and the like. The reason why ¹⁷⁶ Lu is adopted is that the self-radiation Lu decay frequency is extremely low.

That is, the ¹⁷⁶ Lu volume of at most including apparatus systemic PET scanner at ¹⁷⁶ Lu and 5230cm ^{3. 176} Lu has a radioactivity of 300 Bq / cm ³ . Therefore, 5230 cm ³ × 300 Bq / cm ³ = 1.569 MBq. On the other hand, the radiopharmaceutical administered to the subject is generally 185 MBq. Even assuming that all of the detector pairs collapse (in this case β-decay) and all of them are detected, the frequency is approximately 1/118 at 1.569 MBq / 185 MBq≈ 1/118. Only detected in

Further, the number of events detected by collapsing two Lu included in the detector pair during 128 ns is represented by Bq = count / sec because Bq is a count per unit time, and 1 × {1 .569 MBq × (128 ns / 1 sec)} = 0.2 count / sec. Actually, since most of the γ rays from Lu interact in the crystal element in which they are collapsed, the probability that they are accidentally detected at the same time (simultaneously counted) decreases. Thus, when ¹⁷⁶ Lu is adopted, the decay frequency of self-radiated Lu is extremely low, and the data of the timing histogram of FIG. 7 described later can be collected for each detector pair.

When ¹⁷⁶ Lu is used, if the refractive index of the crystal element is n, the transmission of γ rays remains at the light velocity c, whereas the transmission of visible light or ultraviolet light is c / n. The deceleration causes a fundamental TOF time difference (also referred to as “detection time difference” as will be described later) depending on the depth position interacting with the crystal element. Therefore, it is useful when timing calibration in the depth direction employs ¹⁷⁶ Lu. Furthermore, when ¹⁷⁶ Lu is employed, a high-density LYSO or an LSO crystal element described later can be used, and the sensitivity is high and the spatial resolution can be improved. Therefore, it is useful for an apparatus that pursues spatial resolution such as a PET apparatus for small animals.

  In the case of a two-layer DOI detector, as shown in FIG. 5, there are four combinations that take the same LOR in a line (LOR: Line Of Response) connecting the simultaneously counted detectors. A method for performing timing calibration for each combination of the four LORs will be described below. For the determination of each DOI, a waveform discrimination method using a difference in light emission decay time may be used, or a discrimination method using two-dimensional position information using a difference in light path of light emission may be used ( Reference 1: Eiichi Tanaka (Hamamatsu Photonics, former National Institute of Radiological Sciences), Atsushi Kanno (Akita Brain Research Institute), “MIT Virtual Museum PET text14”, [online], ACADEMIA, Internet <URL: http://www.ricoh.co. jp / net-messena / ACADEMIA / JAMIT / MITVM / PET / TANAKA04 / text14.html>, Reference 2: Hideo Murayama (National Institute of Radiological Sciences, Molecular Imaging Center), “Development of Next-Generation PET Equipment—Current Status and Future development-", [online], Japan Society for Radiological Technology, Journal of Japanese Society for Radiological Technology, Internet <http://www.jstage.jst.go.jp/article/jjrt/62/6/62_786/_article / -char / en>).

  Next, in the method of performing timing calibration for each combination of four LORs, specific functions of the time difference information calculation unit 55 and the timing calibration unit 56 will be described with reference to FIGS. 6 is a flowchart showing a flow until a timing histogram (time difference distribution) is obtained, FIG. 7 is a schematic diagram and an actual measurement diagram of a timing histogram between scintillator groups A and A, and FIG. It is an example of setting of a schematic diagram, an actual measurement figure, and an energy window.

Among the photon detectors 32 to be subjected to coincidence counting, one is “detector 1” and the other is “detector 2”. In Example 1, including Example 2 to be described later, as the coincidence event, ¹⁷⁶ Lu collapses in “Detector 1”, and “Detector 1” detects β rays emitted by itself, In the “detector 2”, ¹⁷⁶ Lu is collapsed in the “detector 2” in the event of detecting γ-rays generated by the decay, and in the “detector 2”, the β-rays released by itself are detected, “Detector 1” uses an event when γ-rays generated by the decay are detected.

  In order to simplify the explanation, it is assumed that “detector 1” and “detector 2” are opposed to each other and are separated by 60 cm. The detection time when detected by “detector 1” is T1, and the detection time when detected by “detector 2” is T2. In particular, in the scintillator group A having self-radioactivity, the detection time when detected by the “detector 1” is TA1, and the detection time when detected by the “detector 2” is TA2, and the self-radioactivity is detected. In the scintillator group B that does not have the “detector 1”, the detection time when detected by the “detector 1” is TB1, and the detection time when detected by the “detector 2” is TB2.

(Step S11) Output of Detector 1 (Step S21) Output of Detector 2 The case where photons are respectively detected by “detector 1” and “detector 2” between scintillator groups A-A will be described. ¹⁷⁶ Lu collapses in the scintillator group A of the “detector 1” (step S11), the scintillator group A of the “detector 1” detects the β rays emitted by itself, and the scintillator group A of the “detector 2” Detect broken gamma rays. On the contrary, ¹⁷⁶ Lu collapses in the scintillator group A of “detector 2” (step S12), and the scintillator group A of “detector 2” detects the β ray emitted by itself, and the scintillator of “detector 1”. In group A, the broken gamma rays are detected.

(Step S12) Detection time / (Step S22) Detection time
Assuming that the time when ¹⁷⁶ Lu decays and the detection time when detecting the β-rays emitted by the ¹⁷⁶ Lu are almost the same, that is, if the β-rays emitted by the ¹⁷⁶ Lu itself are detected almost simultaneously, the detection time ¹⁷⁶ Lu collapses in the scintillator group A of “detector 1” at TA1, and the β-rays emitted by itself are detected in the scintillator group A of “detector 1” at detection time TA1 (step S12). Since the emitted photon flies 30 cm in 1 ns, it takes about 2 ns to reach the scintillator group A of “detector 2” and is detected by scintillator group A of “detector 2” at detection time TA2. The On the other hand, ¹⁷⁶ Lu collapses in scintillator group A of “detector 2” at detection time TA2, and β-rays emitted by itself are detected in scintillator group A of “detector 2” at detection time TA2 (step S22). The emitted photons reach the scintillator group A of “detector 1” in about 2 ns, and are detected by the scintillator group A of “detector 1” at the detection time TA1.

(Step S3) Calculation of Detection Time Difference A detection time difference (TA2−TA1) of the detection time TA2 with respect to the detection time TA1 is obtained. When “detector 2” detects slower by α than “detector 1”, ¹⁷⁶ Lu collapses in scintillator group A of “detector 1” and releases itself in scintillator group A of “detector 1” In the event of detecting β rays and detecting the γ rays generated by the decay in the scintillator group A of “detector 2”, the detection time difference (TA2−TA1) is (2 ns + α). On the contrary, ¹⁷⁶ Lu collapses in scintillator group A of “detector 2”, β-rays emitted by scintillator group A of “detector 2” are detected, and in scintillator group A of “detector 1”, In the event of detecting γ-rays generated by decay, the detection time difference (TA2−TA1) is (−2 ns + α). Here, the detection time difference (TA2−TA1) of the detection time TA2 based on the detection time TA1 is obtained, but the detection time difference (TA1−TA2) of the detection time TA1 based on the detection time TA2 may be obtained.

(Step S4) Accumulation of Detection Time Difference Data Actually, the detection time difference (TA2−TA1) has a spread. Therefore, the detection time difference data is accumulated by increasing the number of samples for the above-described event.

(Step S5) Creation of Timing Histogram A time difference distribution (timing histogram) with the horizontal axis as the detection time difference (TA2−TA1) and the vertical axis as the count (count) is as shown in FIG. 7A is a schematic diagram, and FIG. 7B is an actual measurement diagram (detection time difference (TA2−TA1) = Δt [ns]). The vertical axis count is represented by “Count” in FIG.

  The processes from step S11, step S12, step S3, or step S21, step S22, and step S3 are performed from the photon detector 32 through the AD converter 53 and the energy window unit 54 (see FIG. 2). Moreover, about the process of step S3-S5, the time difference information calculation part 55 (refer FIG. 2) performs.

  As is clear from this timing histogram, some photons are (2 ns + α) (in FIG. 7 (a), it is expressed as [ns], so “2 + α”) or (−2 ns + α) (FIG. 7 (a)). In [ns], it is detected with a detection time difference greatly deviated from “−2 + α”), but most photons are detected with a detection time difference in the vicinity of (2 ns + α) or (−2 ns + α). . Therefore, a histogram having two peaks at (2 ns + α) and (−2 ns + α) is obtained.

  In order to obtain α which is a temporal shift, the above-described two peaks are detected (measured), and the time differences (2 ns + α) and (−2 ns + α) at the peaks are detected (measured). Then, from the time difference (2 ns + α) and (−2 ns + α) at each peak, which is the detection result, α can be obtained by a temporal average α (= {(2 ns + α) + (− 2 ns + α)} / 2). In the case where the peaks are close to each other, a method of calculating the center of gravity of the histogram in FIG. 7 may be employed.

  With the above method, it is possible to obtain the temporal deviation α when the photon is detected by the “detector 1” and the “detector 2” between the scintillator groups AA. Therefore, when the photon is detected by the “detector 1” and the “detector 2” between the scintillator groups A-A, the timing calibration unit 56 (see FIG. 1) has a temporal shift α. Perform timing calibration (time correction) to set “0” to “0”. When obtaining the detection time difference (TB2-TA1) when the photons are detected by the “detector 1” and the “detector 2” between the scintillator groups A and B, the “detector 1” is obtained between the scintillator groups B and A. ”And“ detector 2 ”, the detection time difference (TA2−TB1) in the case of detecting the photons respectively can also be obtained in the same manner, so that a temporal deviation can be obtained, and timing calibration is performed. be able to.

  In the first embodiment, the photon detector 32 includes self-radiating crystal elements (scintillator group A formed of LYSO) and non-self-emitting crystal elements (scintillator group B formed of GSO). Is a two-layer DOI detector. In the case of such a crystal element having no self-radiation, photons are detected between detectors having no self-radiation (that is, “detector 1” and “detector” between scintillator groups BB layers). In the case of detecting each photon with “2”), β-decay occurs only in the scintillator group A, and therefore, the temporal deviation between the scintillator groups BB cannot be directly obtained.

  However, the detection time difference between the scintillator groups A-A (TA2-TA1), the detection time difference between the scintillator groups A-B (TB2-TA1), and the detection time difference between the scintillator groups B-A (TA2-TB1). ), The detection time difference (TB2−TB1) between the scintillator groups BB is indirectly determined by setting (TB2−TA1) + (TA2−TB1) − (TA2−TA1) = TB2−TB1. Therefore, it is possible to indirectly determine the time shift between the scintillator groups BB.

  Naturally, even when the scintillator group B has self-radioactivity, the detection time difference and the temporal deviation can be indirectly obtained using the above-described known detection time difference. Of course, the temporal deviation can be obtained from the timing histogram. It can also be obtained directly.

  Furthermore, in order to obtain a timing histogram with high accuracy, the above-described energy window unit 54 (see FIG. 2) may be provided between the AD converter 53 and the time difference information calculation unit 55. What I want in this measurement is the coincidence of β rays (maximum energy 580 KeV) and 307 KeV γ rays as mentioned above. Therefore, as an example, the energy window unit 54 sets two energy windows (energy widths) as shown in FIG. FIG. 8A is a schematic diagram, and FIG. 8B is an actual measurement diagram and an example of setting an energy window. One is an energy window (energy width) EW_1: 400 to 800 KeV for β rays, and the other is an energy window (energy width) EW_2 for γ rays: 200 to 400 KeV. As shown in FIG. 8, EW_1 is an energy window for extracting γ rays (307 KeV) from other detectors, and EW_2 is a self-luminous component (β rays and γ rays), that is, detection with self-radiation. It is an energy window for components that detect photons emitted from the device itself. At this time, since the energy window EW_2 is 200 to 400 KeV, 202 KeV gamma rays are included in the timing histogram in addition to the 307 KeV gamma rays, but there is no problem.

  By using only the coincidence of EW_1 and EW_2 as true data (True), a timing histogram with a good S / N ratio can be obtained, and more accurate timing calibration (time correction) is performed. be able to.

  Timing calibration is performed by the above-described timing calibration unit 56 (see FIG. 2), and occurs between the detection time of one “detector 1” and the detection time of the other “detector 2”. After eliminating the time lag, the radiopharmaceutical is administered to the subject M, and simultaneously counted by the coincidence circuit 51, and the nuclear medicine diagnosis by the PET-CT apparatus 1 is performed. Since there is no time lag after calibration, the coincidence circuit 51 can accurately perform coincidence and accurately perform nuclear medicine diagnosis.

  When a crystal element having self-radiation is used as a detection element, Back Ground is obtained as described above. We have found that this Back Ground can be used to perform timing calibration (time correction). Therefore, according to the PET-CT apparatus 1 according to the first embodiment having the above-described configuration, an external radiation source is not used by including the timing calibration unit 56 that performs time correction using self-radioactivity. Can perform timing calibration (time correction).

In addition, since the timing calibration can be performed only without preparing anything in the time zone where there is no measurement such as at night, there is also an effect that it can be carried out very easily as compared with the conventional case. ^{Since 176} Lu is a long-lived nucleus with a half-life of 37.8 billion years, it is a stable radiation source with negligible aging, and can also be used for detector adjustment (eg timing calibration) and detector status measurement. It also has the effect of being suitable. In addition, since the detector geometry (geometric detector arrangement) and radiation intensity do not change, the conditions are extremely reproducible except for measurement errors. There is also an effect that it is easy to perform.

  In the first embodiment, preferably, an energy window portion 54 for setting a photon energy width (energy window) is provided, and timing calibration is performed using self-radiation within a photon energy width set by the energy window portion 54. The unit 56 performs timing calibration. Only photons having a desired energy width (energy window) (in the first embodiment, β rays (maximum energy 580 KeV) and 307 KeV γ rays) can be extracted by setting the energy window 54, and S A time distribution (timing histogram) with a good / N ratio can be obtained. As a result, more accurate timing calibration (time correction) can be performed.

  In the first embodiment, the photon detector 32 is a DOI detector configured by stacking crystal elements in the depth direction in order to discriminate the position of the light source in the depth direction where the interaction has occurred. By using such a detector, the resolution in the depth direction can be improved and timing calibration can be performed.

  In the present Example 1, the crystal element (scintillator group A formed of LYSO) having self-radiation includes a nuclide that causes β-decay and emits γ-rays accompanying the β-decay. -Β-rays are detected by the detector that caused the decay itself, and γ-rays associated with β-decay are detected by another detector, and the timing calibration is performed using the β-ray detection event and the γ-ray detection event. The calibration unit 56 performs timing calibration.

  In the first embodiment, the photon detector 32 includes self-radiating crystal elements (scintillator group A formed of LYSO) and non-self-emitting crystal elements (scintillator group B formed of GSO). It is configured. When a crystal element that does not have self-radiation is included, it does not have self-radioactivity based on a detection event by a detector composed of a crystal element that has self-radiation and a crystal element that does not have self-radiation. The timing calibration unit 56 can perform timing calibration between the detectors (in this embodiment, the scintillator group B-B). When a detector composed of a crystal element having self-radioactivity is used in combination with another detector, correction data between the detectors having self-radioactivity (in this case, a detection time difference (TA2−TA1)), and Correction data (detection time difference (TB2-TA1) and (TA2-TB1) in this case) of a detector having one self-radioactivity and the other not self-radioactive can be obtained. Can be used to perform timing calibration between detectors that do not have self-radioactivity.

Next, Embodiment 2 of the present invention will be described with reference to the drawings.
FIG. 9 is a side view of a transmission-type PET apparatus according to the second embodiment, and FIG. 10 is a block diagram of the transmission-type PET apparatus according to the second embodiment. In the second embodiment, a nuclear medicine diagnosis apparatus will be described as an example of a tomography apparatus, and a transmission type PET apparatus in which a PET (Positron Emission Tomography) apparatus and a transmission apparatus are combined as a nuclear medicine diagnosis apparatus. An example will be described.

  As illustrated in FIG. 9, the transmission-type PET apparatus 1 according to the second embodiment includes a bed 2 and a PET apparatus 3 as in the first embodiment. In the second embodiment, a transmission device 7 is provided instead of the X-ray CT apparatus 4 of the first embodiment described above. Since the PET apparatus 3 is the same as that in the first embodiment, the description thereof is omitted.

  The transmission device 7 includes a gantry 71 having an opening 71a. In the gantry 71, a radiopharmaceutical to be administered to the subject M, that is, a radiation source 72 for irradiating the same kind of radiation as the radioisotope (RI) (photon in this embodiment), and a photon transmitted through the subject M are contained. A transmission detector 73 for detection is provided. The radiation source 72 is rotated around the body axis of the subject M in the gantry 71 by driving a motor (not shown). The transmission detector 73 is arranged in a ring shape around the body axis of the subject M and is stationary. Of course, similarly to the radiation source 72, the transmission detector 73 may be rotated around the body axis of the subject M.

  In FIG. 9A, the gantry 31 of the PET apparatus 3 and the gantry 71 of the transmission apparatus 7 are separated from each other. However, as in the first embodiment, as shown in FIG. You may comprise.

  Next, a block diagram of the transmission type PET apparatus 1 will be described. As shown in FIG. 10, the transmission-type PET apparatus 1 includes a console 6 in addition to the bed 2, the PET apparatus 3, and the transmission apparatus 7 described above. The block diagram of the PET apparatus 3 and the console 6 is the same as that of the first embodiment except for the data collection unit 61, and thus the description thereof is omitted.

  The transmission data collection unit 74 collects photon absorption coefficient distribution data as transmission data (absorption correction data) based on the photons detected by the transmission detector 73. Transmission data collected by the transmission data collection unit 74 is sent to the data collection unit 61.

  The data collection unit 61 causes the transmission data collected by the transmission data collection unit 77 to act on the PET data that is simultaneously counted by the coincidence circuit 51 and collected to take into account photon absorption in the body of the subject M. The projection data is corrected. That is, the transmission data is applied to the PET data to correct the PET data. The data collection unit 61 reconstructs the projection data subjected to the absorption correction and generates a tomographic image.

  As described in the first embodiment, the radiopharmaceutical is administered to the subject M, and the coincidence counting circuit 51 collects data using the image information based on the photons detected by the photon detector 32 as projection data (PET data). Send to part 61. On the other hand, a photon is irradiated from the source 72 to the subject M while rotating the source 72, and the transmission detector 73 converts the photon irradiated from the outside of the subject M and transmitted through the subject M into an electrical signal. The photon is detected. The electric signal converted by the transmission detector 73 is sent to the transmission data collection unit 74 as image information (pixel value). The transmission data collection unit 74 obtains transmission data (absorption correction data) based on the sent image information. The transmission data collection unit 74 uses the calculation representing the relationship between the photon or X-ray absorption coefficient and energy to convert the CT projection data, that is, the X-ray absorption coefficient distribution data into the photon absorption coefficient distribution data. Then, the distribution data of the photon absorption coefficient is collected as transmission data (absorption correction data). The transmission data collection unit 77 sends transmission data to the data collection unit 61.

  The data collection unit 61 performs the absorption correction of the PET data and sends it to the image reconstruction unit 62. The image reconstruction unit 62 reconstructs the projection data after the absorption correction that has been sent into the body of the subject M. A tomographic image taking into account photon absorption is generated.

  The energy window unit 54, the time difference information calculation unit 55, and the timing calibration unit 56 are the same as those described above except for the configuration in which the transmission device 7 is provided in the second embodiment instead of the X-ray CT apparatus 4 of the first embodiment. Since this is the same as the first embodiment, the description thereof is omitted.

  Since the operation and effect in the second embodiment are the same as those in the first embodiment except that the transmission device 7 is provided in the second embodiment instead of the X-ray CT apparatus 4 in the first embodiment, The description is omitted.

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

  (1) In the first embodiment described above, a description will be given by taking as an example a device that combines a PET device and an X-ray CT device. In the second embodiment described above, a device that combines a PET device and a transmission device will be described as an example. However, the present invention may be applied to a single PET apparatus. In particular, the present invention has an effect that the timing calibration can be performed without the need for an external radiation source. Therefore, the external radiation source as in the second embodiment described above is not necessary. Of course, an external radiation source may be provided as in the second embodiment.

  (2) In each of the above-described embodiments, the nuclear medicine diagnosis apparatus is described as an example of the tomography apparatus, but the present invention may be applied to a tomography apparatus that performs only timing calibration.

  (3) Each embodiment described above includes energy width setting means (energy window portion 54 in each embodiment) for setting a photon energy width (energy window), and is set by the energy width setting means (energy window portion 54). The time correction means (timing calibration unit 56 in each embodiment) performed time correction (timing calibration) using the self-radioactivity within the photon energy width (energy window), but the S / N ratio Etc. or the S / N ratio is sufficiently high, the energy width setting means (energy window part 54) is not necessarily provided.

  (4) In each of the above-described embodiments, the detector is a detector (that is, a DOI detector) configured by stacking crystal elements in the depth direction in order to discriminate the position of the light source in the depth direction in which the interaction has occurred. However, if the resolution in the depth direction is not considered, or if the resolution in the depth direction is sufficiently high, it is not always necessary to be a DOI detector. Therefore, it may be a detector composed of a crystal element having one kind of self-radiation.

  (5) In each of the embodiments described above, the self-radiating crystal element includes β-decay and includes a nuclide that emits γ-rays accompanying the β-decay. The detector itself detects β-rays, and another detector detects γ-rays associated with β-decay, and uses a β-ray detection event and a γ-ray detection event for time correction means (in each example, timing). The calibration unit 56) performs time correction (timing calibration), but is not limited to β-decay. It may be applied to α-decay. For example, the present invention may be applied to a detector made of PWB containing lead (Pb) that causes α-decay.

  (6) In each of the embodiments described above, the detector includes a crystal element having self-radiation (in each embodiment, a scintillator group A formed of LYSO) and a crystal element having no self-radiation (GSO in each embodiment). However, any detector including a self-radiating crystal element may be composed of, for example, a self-radiating crystal element.

  (7) In each of the above-described embodiments, the scintillator group which is a crystal element having self-radiation is provided on the incident side, and the scintillator group which is a crystal element having no self-radiation is provided on the photomultiplier tube side. It may be. It can also be applied to a DOI detector having three or more layers. In addition, a crystal element having self-radiation and a crystal element having no self-radiation are formed separately for each layer. You may mix an element.

(8) In each of the above-described embodiments, LYSO is taken as an example of a crystal element having self-radiation, and GSO is taken as a crystal element having no self-radiation, but the present invention is not limited thereto. For example, as a crystal element containing ¹⁷⁶ Lu, LSO or LGSO may be used, or GSO containing Lu may be used. As for the form of the crystal element having self-radioactivity, a thin film tape made of a substance to which self-radioactivity (for example, Lu) is added, such as GSO containing Lu, is used as a crystal having no self-radioactivity. It may be in a form to be affixed to an element (for example, GSO) or a form in which a coating agent made of a substance to which self-radioactivity (for example, Lu) is added is affixed to a crystal element (for example, GSO) that does not have self-radiation. It may be.

1 is a side view of a PET-CT apparatus according to Example 1. FIG. 1 is a block diagram of a PET-CT apparatus according to Example 1. FIG. It is the schematic of the specific structure of a photon detector. ^{176 is} a decay view of Lu. It is a schematic diagram which shows the combination of the detector in which the line (LOR: Line Of Response) which connects the detectors simultaneously counted is the same. It is a flowchart which shows the flow until a timing histogram (time difference distribution) is obtained. (A), (b) is the schematic diagram and measurement figure of a timing histogram between scintillator group AA layers. (A), (b) is a schematic diagram of an energy spectrum, an actual measurement diagram, and an example of setting an energy window. 6 is a side view of a transmission type PET apparatus according to Embodiment 2. FIG. 6 is a block diagram of a transmission type PET apparatus according to Embodiment 2. FIG.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 ... PET-CT apparatus, transmission type PET apparatus 3 ... PET apparatus 32 ... Photon detector 54 ... Energy window part 56 ... Timing calibration part

Claims (6)

  A tomographic apparatus comprising a detector including a crystal element having self-radiation, comprising a time correction means for performing time correction using the self-radioactivity.   2. The tomography apparatus according to claim 1, further comprising an energy width setting means for setting a photon energy width, wherein the time correction means uses the self-radiation within the photon energy width set by the energy width setting means. A tomography apparatus characterized by performing time correction.   3. The tomography apparatus according to claim 1, wherein the detector is configured by stacking crystal elements in the depth direction in order to discriminate a light source position in the depth direction in which interaction has occurred. A tomography apparatus characterized by being a detector.   4. The tomography apparatus according to claim 1, wherein the crystal element having self-radiation causes α or β-decay and emits γ rays accompanying the α or β-decay. The detector containing the nuclide and detecting the α or β-decay itself detects the α-ray or β-ray, and another detector detects the γ-ray accompanying the α- or β-decay, and α A tomography apparatus wherein the time correction means performs time correction using a detection event of a line or β ray and a detection event of a γ ray.   5. The tomography apparatus according to claim 1, wherein the detector includes the crystal element having self-radiation and the crystal element having no self-radiation. Tomography equipment.   6. The tomography apparatus according to claim 5, wherein detection without self-radiation is performed based on a detection event by the detector composed of the crystal element having self-radioactivity and a crystal element having no self-radioactivity. A tomography apparatus characterized in that the time correction means performs time correction between instruments.

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