JP5024566B2 - Method for obtaining timing correction value for coincidence of PET apparatus - Google Patents

Description

  According to the present invention, the timing of coincidence of a PET apparatus that generates a functional image of a subject by detecting radiation emitted in directions opposite to each other by 180 ° from within the subject who has received administration of a medicine including an RI radiation source. The present invention relates to a method for obtaining a correction value.

  A PET (Positron Emission Tomography) apparatus is known as one of nuclear medicine diagnosis apparatuses. The PET apparatus simultaneously counts the pairs of γ rays emitted from the subject receiving a drug containing an RI radiation source (positron emitting nuclide) in directions opposite to each other by 180 °, thereby distributing the body distribution of the RI radiation source. This is an apparatus for generating a PET image obtained by imaging a living body and observing a biological function, and is applied to a clinical examination of a living body or a diseased patient.

  The PET apparatus includes a detector ring composed of a plurality of PET detector modules for detecting γ rays. The PET detector module is generally composed of a scintillator block having a scintillator array composed of a two-dimensional array of scintillator crystals, and a photomultiplier tube optically coupled to the scintillator block.

  The radiation emitted from within the subject is photoelectrically absorbed by the scintillator crystal of the PET detector module, thereby converting the energy of γ rays into light, and this light is further pulsed by the photomultiplier tube of the PET detector module. The incident position, energy and detection time of the γ-ray are specified based on this signal.

In the PET apparatus, a time window is set in consideration of a delay in the detection time of γ rays by the PET detector, a delay in processing by the detection circuit, and the like. If they are within the time window, the two detected γ-rays are determined to be γ-rays generated simultaneously by the annihilation of one electron-positron pair, and the annihilation point of this electron-positron pair is determined. Are identified on a straight line connecting the PET detectors that detected the γ rays.
Then, the PET apparatus simultaneously counts a pair of γ rays emitted from the inside of the subject in directions opposite to each other by 180 °, and based on the coincidence data collected over a certain time, the distribution image of the RI radiation source in the subject. (PET image) is reconstructed and displayed on the display.

In this case, in order to obtain a more accurate PET image, the coincidence timing is set so that the difference in detection time generated between the pair of PET detectors that detect the γ-ray pair is uniform throughout the detector ring. It is corrected.
In the correction of the timing of the coincidence counting, when a reference radiation source is arranged in the central opening of the detector ring and radiation is emitted from the reference radiation source, when the radiation is detected by the PET detector, the detection time is Recorded together with the identification information of the PET detector that has detected the radiation, and based on the recorded information, for all the PET detector pairs located within the field of view of the PET apparatus, the detection time generated between the paired PET detectors After obtaining the timing histogram by calculating the difference and counting the difference in detection time, the timing correction value is obtained by obtaining the peak position of the timing histogram.

  However, since the peak position of the timing histogram is determined based on the number of bins in the histogram, the accuracy of the timing correction value is limited by the measurement accuracy of the radiation detection time (time resolution of the PET apparatus), and therefore the timing In order to increase the accuracy of the correction value, the time resolution of the PET apparatus must be improved, but this is not easy to achieve.

Further, in order to increase the accuracy of the PET image, it is necessary to increase the spatial resolution of the PET detector, that is, the detection accuracy of the incident position of the γ-ray with respect to the PET detector.
By the way, the incident position of the γ-ray is detected by identifying the scintillator crystal that has emitted light by the interaction with the γ-ray, so that the γ-ray is incident on the detection surface of the PET detector obliquely and in the deep part of the scintillator crystal. When light emission occurs, a deviation occurs between the actual γ-ray incident position and the γ-ray detection position, thereby reducing the spatial resolution of the PET detector.

  Accordingly, a DOI detector has been proposed in which a plurality of micro scintillator crystals are stacked two-dimensionally to form a scintillator block and a photomultiplier tube is optically coupled thereto. According to the DOI detector, radiation detection information in the depth direction of the scintillator crystal can be obtained, so that a reduction in spatial resolution can be suppressed even when radiation is incident obliquely on the detection surface of the DOI detector (for example, Patent Documents). 1 and 2).

However, in the scintillator crystal, γ rays propagate at the speed of light, whereas fluorescence (emission by the scintillator crystal) propagates only at a speed of about 1/3 of the speed of light. For this reason, in the case of a PET detector having a laminated structure of scintillator crystals such as a DOI detector, the detection time by the scintillator crystal layer close to the photomultiplier tube is due to the scintillator crystal layer far from the photomultiplier tube. It becomes shorter than the detection time. Furthermore, even in one DOI detector, the quality varies for each layer of scintillator crystals, and the measurement error by the detection circuit differs for each layer of scintillator crystals.
For this reason, there has been a problem that the timing correction of the coincidence counting cannot be performed by the same method as the conventional method.

JP 2005-43104 A JP 2005-90979 A

Therefore, an object of the present invention is to provide a method capable of increasing the accuracy of the timing correction value for coincidence counting without increasing the time resolution of the PET apparatus.
It is another object of the present invention to provide a method capable of obtaining the timing correction value of the coincidence count of the PET apparatus with high accuracy even when the detector ring is composed of a DOI detector. .

In order to solve the above-mentioned problems, the first invention has a detector ring composed of a plurality of PET detectors, and received a drug containing an RI radiation source in the central opening of the detector ring. When the subject is placed, a pair of radiations emitted from the inside of the subject in an opposite direction of 180 ° are simultaneously counted by the PET detectors facing each other, and the subject is used by using the accumulated coincidence information. A timing correction value for coincidence counting of a PET apparatus for reconstructing a distribution image of an RI source inside
(1) disposing a reference radiation source in the central opening of the detector ring and irradiating radiation from the reference radiation source to a detection surface of the detector ring;
(2) When the radiation is detected by the PET detector, the detection time is recorded together with identification information of the PET detector that has detected the radiation;
(3) Based on the information recorded in the step (2), for all the pairs of PET detectors located in the field of view of the PET apparatus, a difference in detection time generated between the paired PET detectors Obtaining a timing histogram for each pair of PET detectors by counting the difference in detection time for each pair of PET detectors;
(4) For each pair of the PET detectors, the timing correction value for each pair of the PET detectors is obtained by obtaining the peak position of the timing histogram in a unit smaller than the unit of the time axis scale of the timing histogram. And obtaining the method.

  In the first invention, when the PET detector is a DOI detector having a laminated structure of scintillator crystals, the different PET detectors are formed for each layer of the scintillator crystals in the DOI detector. It is preferable to execute the steps (1) to (4).

In order to solve the above-mentioned problem, the second invention has a detector ring composed of a plurality of PET detectors, and administration of a medicine containing an RI radiation source is provided in the central opening of the detector ring. When the received subject is placed, a pair of radiations emitted from the inside of the subject in opposite directions by 180 ° are simultaneously counted by the PET detectors facing each other, and the accumulated count information is used to A method for obtaining a timing correction value for coincidence of a PET apparatus for reconstructing a distribution image of an RI radiation source in a subject,
(1) disposing a reference radiation source in the central opening of the detector ring and irradiating radiation from the reference radiation source to a detection surface of the detector ring;
(2) When the radiation is detected by the PET detector, the detection time is recorded together with identification information of the PET detector that has detected the radiation;
(3) Based on the information recorded in the step (2), for all the pairs of PET detectors located in the field of view of the PET apparatus, a difference in detection time generated between the paired PET detectors Calculating a timing histogram by counting the difference between the detection times; and
(4) obtaining a timing correction value of the detector ring by obtaining a peak position of the timing histogram in a unit finer than a unit of a scale of a time axis of the timing histogram. The method to do is comprised.

  In the second invention, when the PET detector is a DOI detector having a scintillator crystal laminated structure, the PET detector different for each layer of the scintillator crystal is formed in the DOI detector. It is preferable to execute the steps (1) to (4).

  According to the present invention, the timing histogram obtained for each pair of PET detectors constituting the detector ring or the peak positions of the timing histograms obtained for all pairs of PET detectors constituting the detector ring Since the timing correction value for the coincidence counting of the PET apparatus is obtained in a unit smaller than the unit of the time axis scale of the timing histogram, the measurement accuracy of the radiation detection time (time resolution of the PET apparatus) is conventional. The timing correction value can be obtained with the same accuracy as that obtained when the time resolution of the PET apparatus is increased.

  Further, according to the present invention, when the PET detector is a DOI detector having a scintillator crystal laminated structure, it is assumed that a different PET detector is formed for each scintillator crystal layer in the DOI detector. As in the case of a conventional PET detector, the timing correction value for coincidence counting can be easily obtained with high accuracy.

It is a flowchart of the method of calculating | requiring the timing correction value of the coincidence counting of PET apparatus by one Example of this invention. FIG. 6 is a flow diagram of a method for obtaining timing correction values for coincidence of a PET apparatus according to another embodiment of the present invention. It is a figure which shows schematic structure of the principal part of PET apparatus with which the method of this invention is applied. It is a perspective view which shows the structure of the PET detector module of the PET apparatus to which the method of this invention is applied. It is a graph which shows a timing histogram when the time resolution of a PET apparatus provided with a DOI detector having a two-layer structure of scintillator crystals is 0.5 ns. It is a graph which shows a timing histogram when time resolution of a PET apparatus provided with a DOI detector having a two-layer structure of scintillator crystals is 2 ns.

Hereinafter, preferred embodiments of the present invention will be described with reference to the accompanying drawings. FIG. 1 is a flowchart of a method for obtaining a timing correction value for coincidence counting of a PET apparatus according to an embodiment of the present invention, and FIG. 3 is a schematic configuration of a main part of the PET apparatus to which the method of the present invention is applied. FIG. FIG. 4 is a perspective view showing the configuration of the PET detector module of the PET apparatus to which the method of the present invention is applied.
Referring to FIG. 3, the PET apparatus includes a detector ring 1 composed of a plurality of PET detector modules 2. Then, the top plate 5 that supports the subject 4 receiving the medicine containing the RI radiation source in the supine position is detected so that the body axis of the subject 4 and the axis of the detector ring 1 are parallel to each other. It is arranged in the central opening 3 of the detector ring 1 and is reciprocated along the axial direction of the detector ring 1.

  In this embodiment, the PET detector module 2 includes a scintillator block 12 having a scintillator array composed of a two-dimensional array of scintillator crystals 11 (detection element matrix is 9 × 10), as shown in FIG. The scintillator block 12 is composed of four photomultiplier tubes 13 optically coupled.

  In the PET detector module 2, one PET detector is formed from each scintillator crystal 11 and a photomultiplier tube 13 optically coupled thereto, and the radiation emitted from the subject 4 is The scintillator crystal 11 of the PET detector is photoelectrically absorbed, whereby the energy of the radiation is converted into light, which is further converted into an electrical signal by the photomultiplier tube 13.

  A pre-circuit 6 is connected to the subsequent stage of each PET detector module 2, and each pre-circuit 6 outputs a detected wave height value, detection time data, and data including the PET detector identification information. The The coincidence counting circuit 7 performs coincidence counting based on the detection time data and the identification information of the PET detector, specifies the radiation detection position, and records it in the memory unit 8. The image reconstruction unit 9 performs image reconstruction processing based on the coincidence count data recorded in the memory unit 8 to generate a PET image. The generated PET image is displayed on the display device 10.

  With reference to FIG. 1, according to the method of the present invention, a reference radiation source is first disposed in the central opening 3 of the detector ring 1, and radiation is emitted from the reference radiation source to the detection surface of the detector ring 1. Irradiation (step 1). In this embodiment, the reference radiation source is made of a nuclide that emits 511 keV gamma rays.

Next, when radiation (γ rays) is detected by the PET detector, the detection time is recorded together with the identification information of the PET detector that detected the radiation (γ rays) (step 2).
In this case, the PET detector (scintillator crystal 11) that detects the radiation (γ rays) in the PET detector module 2 is identified in the pre-circuit 6 as follows. That is, the center of gravity is calculated using the collected output signals of the photomultiplier tubes 13, and the obtained center of gravity is projected onto a two-dimensional position map divided into 9 × 10 regions. Then, a position lookup table corresponding to each scintillator crystal 11 on the two-dimensional position map is created, and thereby, the scintillator crystal 11 that interacts with radiation (γ rays) is specified.

  Further, at this time, in the pre-circuit 6, the detection time of radiation (γ rays) by the PET detector (scintillator crystal 11) is determined by, for example, the LET (Leading Edge Trigger) method or the CFD (Constant Fraction Discriminator) method. Is done. The LET method is a method in which when the energy (intensity) of γ rays rises, the time when the energy reaches a preset threshold is used as the detection time. The CFD method shifts a curve obtained by increasing a measurement value by a predetermined value and a curve with a measurement value as it is in the time axis direction with respect to the time-energy (intensity) curve of radiation (γ-rays) and also in the energy axis direction. This is a method in which the intersection of both curves is determined and the intersection is set as the detection time of radiation (γ rays). Further, a TDC (Time to Digital Converter) element may be used.

  Then, based on the information recorded in step 2, for all the PET detector pairs located within the field of view of the PET apparatus, a difference in detection time generated between the PET detectors forming the pair is calculated, and PET detection is performed. A timing histogram for each pair of PET detectors is obtained by counting the difference in detection time for each pair of devices (step 3).

  Further, for each pair of PET detectors, the peak position of the timing histogram is obtained in a unit finer than the unit of the time axis scale of the timing histogram, and a timing correction value for each pair of PET detectors is obtained (step). 4).

Step 4 will be specifically described.
In the timing histogram, count values are plotted for each difference in detection time. Therefore, the unit of the scale on the time axis of the histogram is determined by the measurement accuracy of the detection time. For example, if the measurement accuracy of the radiation detection time (= time resolution of the PET apparatus) is 2 ns, the unit of the time axis of the timing histogram is 2 ns. In this case, according to the conventional method, since the peak position of the timing histogram is obtained in units of 2 ns, the timing correction accuracy is also 2 ns. Therefore, in the conventional method, in order to increase the timing correction accuracy, the time resolution of the PET apparatus has to be increased.
In the present invention, the peak position of the timing histogram is obtained in units smaller than the unit of the time axis scale of the timing histogram, and the timing correction value is calculated from the obtained peak position, so that the time resolution of the PET apparatus remains the same as before. So, the timing correction accuracy has been improved.

  As a method of calculating the peak position of the timing histogram, for example, in view of the fact that the timing histogram is basically symmetrical with respect to the peak, the method of obtaining the peak position from the centroid calculation, or the timing histogram by a Gaussian function It is possible to use a method of approximating and obtaining the peak position in a unit smaller than the unit of the time axis scale of the timing histogram based on the approximate curve.

  Thus, according to the method of the present invention, the peak position of the timing histogram obtained for each pair of PET detectors constituting the detector ring is obtained in a unit finer than the unit of the time axis scale of the timing histogram. By obtaining the timing correction value for the coincidence counting of the apparatus, the timing correction value can be obtained with the same accuracy as that obtained when the time resolution of the PET apparatus is increased while the time resolution of the PET apparatus remains the same. .

Although one preferred embodiment of the present invention has been described above, the configuration of the present invention is not limited to the above embodiment. For example, in the above embodiment, the present invention is applied to a PET apparatus in which the PET detector module constituting the detector ring has a single layer structure of a scintillator array composed of a two-dimensional array of scintillator crystals. Can also be applied to a PET apparatus provided with a detector ring composed of a DOI detector module as shown in FIG. 4B, a DOI detector module 2 ′ includes a scintillator block 12 ′ having a two-layer structure of a scintillator array composed of a two-dimensional array of scintillator crystals 11 ′ (detection element matrix is 9 × 10), and The four photomultiplier tubes 13 are optically coupled to the scintillator block 12 '.
In the DOI detector module 2 ', one DOI detector is formed from a pair of scintillator crystals 11' stacked one above the other and a photomultiplier tube 13 optically coupled thereto.

In the case of the DOI detector module, steps 1 to 4 in FIG. 1 are executed on the assumption that a different PET detector is formed for each layer of scintillator crystals in the DOI detector. Thereby, the timing correction value can be obtained with high accuracy in the same manner as in the case of the normal PET detector described above.
That is, in the case of a two-layer DOI detector as shown in FIG. 4B, there are 2 ² = 4 PET detector pairs for one pair of DOI detectors. Therefore, for each of these four PET detector pairs, the difference in detection time of radiation (γ rays) is counted to obtain a timing histogram, the peak position is obtained, and the timing correction value is calculated.

  Next, in order to confirm whether the accuracy of the timing correction value is increased by the method of the present invention, the same effect as that obtained by increasing the time resolution of the PET apparatus can be obtained. The relationship with accuracy was examined experimentally. In this experiment, a PET apparatus provided with a detector ring composed of the DOI detector module shown in FIG. 4B was used. For each DOI detector, a separate PET detector is formed for each DOI layer, with the scintillator crystal layer on the side far from the photomultiplier tube being the DOI0 layer and the scintillator crystal on the near side being the DOI1 layer. Each time radiation (γ rays) was detected, the detection time and the identification information of the detected PET detector were recorded. Based on this recorded information, the difference in detection time generated between the PET detectors is calculated for all PET detector pairs located within the field of view of the PET apparatus, and the difference in detection time for each PET detector pair is calculated. To obtain a timing histogram.

  The acquired timing histograms are shown in the graphs of FIGS. In FIGS. 5 and 6, for the sake of clarity, only timing histograms relating to a pair of a PET detector formed from the DOI1 layer and a PET detector formed from the DOI0 layer are shown. Further, the timing histogram in FIG. 5 is obtained when the time resolution of the PET apparatus is 0.5 ns, and the timing histogram in FIG. 6 is obtained when the time resolution of the PET apparatus is 2 ns.

Then, the respective peak positions of the timing histograms of FIGS. 5 and 6 were obtained, and the timing correction value was obtained. From the timing histogram of FIG. 5, timing correction values of -1.5 ns (DOI0 layer-DOI1 layer pair) and 2.5 ns (DOI1 layer-DOI0 layer pair) were obtained. Further, from the timing histogram of FIG. 6, timing correction values of −2 ns (DOI0 layer−DOI1 layer pair) and 2 ns (DOI1 layer−DOI0 layer pair) were obtained.
That is, it can be seen that the higher the time resolution, the higher the accuracy of the timing correction value obtained. From this, it was found that, when the accuracy of the timing correction value is increased by the method of the present invention, the same effect as that obtained by increasing the time resolution of the PET apparatus can be obtained.

In the embodiment of FIG. 1, the timing correction value is obtained for each pair of PET detectors constituting the detector ring. However, according to the present invention, the timing correction value of the detector ring can also be obtained. FIG. 2 shows a flowchart of an embodiment for obtaining the timing correction value of the detector ring.
As shown in FIG. 2, in this embodiment, steps 1 to 2 have the same configuration as that of the embodiment of FIG.

  Then, based on the information recorded in step 2, for all the PET detector pairs located in the field of view of the PET apparatus, a difference in detection time occurring between the PET detectors forming the pair is calculated, A timing histogram is acquired by counting the difference in detection time (step 3).

  Further, the peak position of the timing histogram is obtained in a unit finer than the unit of the time axis scale of the timing histogram, and the timing correction value of the detector ring is acquired (step 4). In this case, the method for calculating the peak position of the timing histogram is the same as in the embodiment of FIG.

  Also in this embodiment, as in the embodiment of FIG. 1, not only a PET detector having a single layer structure of scintillator crystals but also a DOI detector having a laminated structure of scintillator crystals is included in the DOI detector. Assuming that different PET detectors are formed for each layer of scintillator crystals, a highly accurate timing histogram can be easily acquired by executing steps 1 to 4.

  In the present invention, the accuracy of the timing correction value is increased by obtaining the peak position of the timing histogram in units smaller than the time resolution of the PET apparatus. However, as another method for increasing the accuracy of the timing correction value, detection is performed. By adding a low-order 1 bit to the radiation detection time data output from each PET detector of the detector ring using a random number, the number of data bits is increased, thereby simulating the time resolution of the PET apparatus. It is also conceivable to increase the accuracy of the timing correction value by expanding it.

DESCRIPTION OF SYMBOLS 1 Detector ring 2 PET detector module 2 'DOI detector module 3 Central opening part 4 Subject 5 Top plate 6 Prefix circuit 7 Simultaneous counting circuit 8 Memory part 9 Image reconstruction part 10 Display apparatus 11, 11' Scintillator crystal 12, 12 'scintillator block 13 photomultiplier tube

Claims (4)

A detector ring composed of a plurality of PET detectors, and when a subject receiving a drug containing an RI radiation source is disposed in a central opening of the detector ring, A pair of radiations emitted in the opposite directions from the inside by 180 ° are simultaneously counted by the PET detectors facing each other, and a distribution image of the RI radiation source in the subject is reconstructed using the accumulated coincidence information. A method of obtaining a timing correction value for coincidence counting of a PET apparatus,
(1) disposing a reference radiation source in the central opening of the detector ring and irradiating radiation from the reference radiation source to a detection surface of the detector ring;
(2) When the radiation is detected by the PET detector, the detection time is recorded together with identification information of the PET detector that has detected the radiation;
(3) Based on the information recorded in the step (2), for all the pairs of PET detectors located in the field of view of the PET apparatus, a difference in detection time generated between the paired PET detectors Obtaining a timing histogram for each pair of PET detectors by counting the difference in detection time for each pair of PET detectors;
(4) For each pair of the PET detectors, the timing correction value for each pair of the PET detectors is obtained by obtaining the peak position of the timing histogram in a unit smaller than the unit of the time axis scale of the timing histogram. Obtaining the method.   When the PET detector is a DOI detector having a laminated structure of scintillator crystals, it is assumed that the different PET detectors are formed for each layer of the scintillator crystals in the DOI detector. The method of claim 1, wherein: (4) is performed. A detector ring composed of a plurality of PET detectors, and when a subject receiving a drug containing an RI radiation source is disposed in a central opening of the detector ring, A pair of radiations emitted in the opposite directions from the inside by 180 ° are simultaneously counted by the PET detectors facing each other, and a distribution image of the RI radiation source in the subject is reconstructed using the accumulated coincidence information. A method of obtaining a timing correction value for coincidence counting of a PET apparatus,
(1) disposing a reference radiation source in the central opening of the detector ring and irradiating radiation from the reference radiation source to a detection surface of the detector ring;
(2) When the radiation is detected by the PET detector, the detection time is recorded together with identification information of the PET detector that has detected the radiation;
(3) Based on the information recorded in the step (2), for all the pairs of PET detectors located in the field of view of the PET apparatus, a difference in detection time generated between the paired PET detectors Calculating a timing histogram by counting the difference between the detection times; and
(4) obtaining a timing correction value of the detector ring by obtaining a peak position of the timing histogram in a unit finer than a unit of a scale of a time axis of the timing histogram. how to.   When the PET detector is a DOI detector having a laminated structure of scintillator crystals, it is assumed that the different PET detectors are formed for each layer of the scintillator crystals in the DOI detector. The method according to claim 3, wherein: (4) is performed.

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