JPH07113873A - Scattering simultaneous counting measuring method by gamma-ray absorber of positron tomograph and the same tomograph - Google Patents

Abstract

PURPOSE:To simply, accurately measure scattering simultaneous count by placing a gamma-ray absorber between each pair of detectors of many pairs of detector rings disposed at symmetrical positions and rotating it along a concentric circle with the rings. CONSTITUTION:A rodlike gamma-ray absorber 26 is placed inside gamma-rays detector rings arranged in a ring state, and rotated on a concentric circumference with the rings. When the absorber 26 is moved to opposed lines of a pair of gamma-ray detectors 12-N1 and 12-N2, positrons at a point 29 of an object 30 to be exposed collide with electrons of the object 30. Even if 7-rays are emitted toward the detectors N1, N2, the gamma-rays directed toward the detector N1 are absorbed the absorber 26, and only the t-rays directed toward the detector N2 are detected. Accordingly, scattering simultaneous counted value indicated at the other positrons 31 is measured by the detector N1, N2. If there is no absorber 26, an ordinary simultaneous counting is measured. A true simultaneous counting is obtained by removing scattering simultaneous counting from the measured simultaneous counting, and a positron image having high quantitativity is reformed.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention proposes an improvement of a measuring method for a positron emission tomography apparatus, in which a large number of pairs of detectors are arranged opposite to each other in a ring shape on a line passing through a subject, While rotating the γ-ray absorber on a concentric circle that surrounds the subject, the positrons emitted from the drug labeled with the positron-emitting nuclide injected into the subject and the electrons in the subject are combined in opposite directions. A method for detecting emitted γ-rays, separating only the scattered same clock value from the measured value containing the scattered same clock value and the true same clock value, and obtaining the true same clock value from both measured values .

[0002]

2. Description of the Related Art A positron emission tomography apparatus detects γ-rays by utilizing the characteristic of emitting two γ-rays that are 180 degrees opposite to each other when the positron and the electron in the living body are combined to annihilate the positron. This is a device for obtaining a distribution image of positrons.
The purpose of this study is to quantitatively investigate the functions such as metabolism of cells by labeling drugs with positron-emitting radionuclides and injecting them into the body, and measuring temporal changes in positron distribution images with a positron emission tomography device. To do. Therefore, the positron distribution image is required to be quantitative.

FIG. 1 is an explanatory view showing the outline of the configuration of a conventional positron emission tomography apparatus 1. In FIG. 1, 2 is a cyclotron, 3 is a labeling compound automatic synthesizer, and the compound circulating in the labeling drug automatic synthesizer 3 is irradiated with radiation of higher energy than the cyclotron 2 to make a positron labeled drug 4, and this positron labeled drug 4 is produced. Is injected into the subject 5, and a large number of γ-rays emitted when the positrons emitted from this drug are combined with the electrons in the living body and disappeared are provided on the circumference of mutually opposite positions on the same line. It is the principle of the positron emission tomography apparatus that obtains a positron distribution image by measuring the measured values, collecting and processing the measured values by the data collecting system 7 and the data processing system 8.

The injection of the positron-labeled drug 4 into the subject 5 can be carried out, for example, by supporting it on glucose. Positrons emitted from the drug injected into the subject have the property of combining with the electrons in the living body and disappearing together to generate a pair of γ rays. Since these two γ rays are emitted in a direction of 180 degrees with respect to each other, two γ rays are simultaneously counted when measured with two γ ray detectors as shown in FIG. In FIG.
9 is a positron emitted in the living body, 10 is an electron in the living body,
Reference numeral 11 is a γ ray, 12 is a γ ray detector, 13 is a time pick-off circuit, and 14 is a clock circuit. Here, when a positron collides with an electron in the living body, a pair of γ rays is generated. The degree of simultaneous emission of gamma rays depends on the time resolution of the detector system and is typically within a few nanoseconds. On the contrary, if the two detectors 12 and 12 arranged as shown in FIG. 2 perform coincidence counting, the positron was on the opposite line of these two detectors (this is called true coincidence counting). , This opposite line is called coincidence line).

Therefore, if the γ-ray detectors are arranged in a ring and the γ-rays from the positrons in the ring are measured, information on the distribution function of the positrons can be obtained. The conventional positron emission tomography apparatus includes a γ-ray detector 15 called a gantry in which a large number of γ-ray detectors 12 are arranged on the circumference as shown in FIG. 3, and a data collection system for collecting and analyzing data from the gantry 15. 7. A data processing system 8 that creates an image from the data and analyzes the image.

In FIG. 3, the data collection system 7 forms a detector ring 6 by a large number of pairs of γ-ray detectors 12 symmetrically arranged on the circumference, and the portion provided with the detector ring 6 is a gantry part. The gamma ray is detected by positioning the subject 5, for example, the head, chest, abdomen or the like in the detector ring 6.

The conventional positron emission tomography apparatus synthesizes a positron-emitting nuclide generated by high-energy radiation from a cyclotron and a drug or compound (eg glucose) to be injected by positron labeling and injection into a subject. , The positrons released from this positron-labeled drug are required electrons in the living body (for example, brain tumor or cancer cells)
When it collides with and disappears, the generated γ rays are detected, so
In order to generate a positron-emitting nuclide with a cyclotron and attach a positron label to a drug, the cyclotron cannot be used unless it is near a laboratory or the like that owns the cyclotron.
Although there are only a dozen or so positron emission tomography devices in Japan, the measured value by this device has a drawback that the scattered same clock value is mixed with the true same clock value to cause a false image. Therefore, if the conventional device can separately measure the scattered same clock value and the true same clock value, its value becomes an astronomical leap.

As shown in FIG. 4, the conventional positron emission tomography apparatus has a distribution S ₀ (x, θ) of coincidence counts in a certain direction.
(This is called a sinogram) is obtained, and the positron distribution function ρ (X, Y) is obtained as an image by the following equation using a computer.

[Equation 1] Here, A (x, θ) is a factor that corrects the effect of absorption of γ rays in the subject, and f (x) is a mathematical factor for image reconstruction called a filter function. S ₀ (x, θ)
Is related to ρ (X, Y) by the following equation.

[Equation 2]

Further, the detector rings 6 are overlapped in multiple layers, and positron distribution images of a plane in each ring or a plane between two detection rings are simultaneously measured, and the spatial distribution of positrons ρ (X, Y, Z) is required.

The above is a brief description of the principle and apparatus for obtaining a distribution image of positrons by the conventional planar image reconstruction method of positron distribution. A positron tomographic layer in which detector rings are stacked in multiple layers is described. The imaging device simultaneously captures not only the plane direction of the cross section containing the positron but also the three-dimensionally emitted γ-rays. Therefore, it is designed to reconstruct a positron distribution image from three-dimensional coincidence counting in a conventional device rather than a plane to dramatically increase the detection efficiency, and thereby to determine the amount of radioisotope to be administered into the body. I am less. However, the measurement value obtained by the three-dimensional coincidence counting method has a drawback in that there is a large possibility that false coincidence (error) occurs due to a large amount of scattered coincidence counting mixed into the true same clock value.

[0011]

The positron concentration obtained by a conventional positron emission tomography apparatus shows the degree of metabolism of cells by a drug labeled with a positron emitting nuclide. Since the function of the organ is diagnosed by quantitatively measuring and analyzing the positron distribution concentration, the positron distribution image needs to be quantitative. Scatter coincidence is a major cause of impairing the quantitative property. That is, just because γ-rays are simultaneously counted by two opposing detectors, it cannot be said that only γ-rays from positrons on the opposing line are measured. There are mainly the following coincidences that are not true coincidences.

(A) Simultaneous Scattering Counting: -γ rays emitted in the subject may be scattered by electrons in the subject. This scattering is called Compton scattering. In the scattering shown in FIG. 5B, the detector 12A,
12C simultaneously counts, and the positrons are measured as if they were on the opposite lines of the pair of detectors 12A and 12C. This is a malfunction. While true coincidence only measures positrons in the plane of the detector ring, this scatter coincidence also measures gamma rays from positrons in a volume containing that plane, so The contribution, ie the value of the scatter coincidence, is not negligible compared to the true coincidence.

At present, the following method is considered as a correction method for the simultaneous scattering count. (1) Use of septum In order to prevent the detector of the detector ring from detecting γ-rays other than positrons in the plane formed by the detector ring, in a multilayer positron tomography device (hereinafter referred to as positron CT device) Attach collimators 23 to both sides of the detector ring as shown in FIG. In FIG. 6, reference numeral 24 shows an image of inter-layer simultaneous counting, and 25 shows an image of intra-layer simultaneous counting. This collimator 23 is called a septum, which is also called a kind of slide shield plate, and it is possible to block and remove coincidence scattering from other than the detector ring plane, but from within the ring plane. The scatter coincidence count of cannot be eliminated. Moreover, since the septum is provided for collimation, the sensitivity is lowered. In particular, a positron emission tomography apparatus for three-dimensional image reconstruction cannot be used because it eliminates a combination of detectors that can be simultaneously counted in a three-dimensional space.

(2) Simulation Method From images obtained from coincidence counting including scatter coincidence counting,
This is a method of obtaining the influence of the scattering coincidence count by calculation, but this method has low reliability because the measured value is not the actual measured value.

(3) Method of selecting energy of γ-rays The energy of γ-rays generated from positrons is 511 keV, but when scattered, it becomes smaller than 511 keV. Therefore,
The scatter coincidence count is reduced by taking the coincidence count of γ-rays having an energy above a certain threshold. When the gamma ray energy range is divided into two and measured, it is possible to separate scattered coincidence counting and true coincidence counting (unscattered coincidence counting). However, in this method, the measuring circuit becomes very complicated and its discrimination depends largely on the energy resolution of the detector. Therefore, it is not possible to make a good discrimination with the current accuracy of the detector.

(B) Random coincidence counting: -Similar to scatter coincidence counting, there is random coincidence counting as a coincidence counting that is not a true coincidence counting. This is because two positrons disappear at the same time (within the time resolution of the measurement system),
When two of the four γ-rays are detected and the other two are not detected, the coincidence count is measured as if there was one positron on the opposite line of the detectors 12A and 12B. . However, this coincidence coincidence can be measured by coincidentally counting two γ-rays at a time interval at which the measurement system can determine that they are not coincident. This method is currently commonly used to measure contingency coincidences.

(C) Effect of scattering coincidence counting on positron image; _-If the contribution of scattering coincidence counting is B _sc (x, θ), the distribution S (x, θ) of coincidence counting actually measured is as follows. Given by the formula.

[Equation 3] Here, since the coincidence coincidence count can be removed by hardware by the method described above, the formula (3) is a coincidence coincidence count with its contribution removed. Further, B _sc (x, θ) also includes the effect of absorption in the subject.

Therefore, the positron distribution image ρ ₁ (X, Y) obtained from this S (x, θ) is

[Equation 4] Becomes Where Δρ (X, Y) is

[Equation 5] It is a false image by the coincidence counting of scattering given by.

In research and diagnosis using a positron emission tomography apparatus, since the quantification of images is important, this false image is a great obstacle, and many efforts have been made to eliminate it, and the method described above is used. Has been figured out.

Here, the influence of the scattering coincidence counting has been described by taking the two-dimensional image reconstruction method as an example. However, in the three-dimensional image reconstruction method, the influence of the false image on the image becomes larger.

[0021]

According to the present invention, a γ-ray absorber is placed between each pair of detectors of a multiplicity of pairs of detector rings which are arranged symmetrically with respect to each other on the circumference of a circle, and detect this. It is a simple and accurate method of measuring the scattering coincidence count in the positron image of a positron tomography apparatus by rotating it along a concentric circle with the instrument ring. Simultaneous counting is required to give a highly quantitative positron image.

According to the present invention, the detector of the positron emission tomography apparatus and the γ-ray absorber are positioned on opposite lines of the detector and supported by an appropriate supporting device, and the γ-ray absorber is arranged on a concentric circle surrounding the subject. Of the γ rays emitted from the positrons in the subject while rotating, the true same-clock value is excluded and only the scattered same-clock value is measured, and at the same time, the sum of the true same-clock value and the scattered same-clock value is simultaneously counted. , Is a simultaneous coincidence counting measurement method using a γ-ray absorber in a positron emission tomography apparatus, which is characterized in that true coincidence counting is obtained from the difference between the two.

According to the present invention, in a γ-ray detector of a positron emission tomography apparatus in which γ-rays emitted from positrons in a subject are detected by a large number of pairs of γ-ray detector rings, a large number of pairs of detectors arranged opposite to each other are detected. Positron tomography, characterized in that a rotator equipped with a γ-ray absorber located inside the detector ring is provided, and the γ-ray absorber is rotated concentrically inside the detector ring. On the device.

[0024]

The structure and operation of the present invention will be described below with reference to the drawings. According to the present invention, as shown in FIG. 7, a γ-ray absorber 26 is placed inside a γ-ray detector ring 6 arranged in a ring shape, and the γ-ray absorber 26 is rotated on a circle concentric with the detector ring 6.
Two gamma rays from the subject are counted simultaneously. In the multiplex positron emission tomography apparatus in which the detector rings 6 are overlapped in multiple layers, for example, as shown in FIG.
Rotate 26 to simultaneously count two gamma rays from the subject. Two detectors 12A and 12B that simultaneously count two γ rays
If there is a γ-ray absorber 26 on the opposite line of (1), this is labeled as scattering coincidence counting and recorded and stored. If the γ-ray absorber 26 was not present, it is recorded and stored as a normal coincidence count. γ
By rotating the line absorber 26, scatter coincidence and normal coincidence are measured simultaneously.

As shown in FIG. 8, the γ-ray absorber 26 is placed on the line opposite to the two detectors 12A and 12B which simultaneously count the two γ-rays.
If this is the case, this is clearly not the γ-rays from the positron annihilation that were on the opposite lines of the detectors 12A and 12B, but the γ-rays from the positrons other than the opposite lines of the detector are Compton scattering ( It was detected because of C). The scatter coincidence thus obtained is used to reconstruct the image from the true coincidence count, which does not include the scatter coincidence.

A positron tomography apparatus using the γ-ray absorber of the present invention is a multilayer positron tomography apparatus as shown in FIG.
The collimator (shielding plate) 23 as shown in FIG. 4 is not used, and the γ-ray absorber is rotated on a concentric circle inside the detector ring 6 to detect each pair of detectors 12A, When the opposite line of 12B passes through the γ-ray absorber, the same clock value by the pair of detectors 12A and 12B is measured as the scattering coincidence count. The principle of image reconstruction of true coincidence counting not including scatter coincidence counting using scatter coincidence counting obtained by coincidence counting measurement by the device of the present invention is as follows.

(1) Two-dimensional image reconstruction The scatter coincidence count obtained by the method of the present invention is B
If (x, θ), this is

[Equation 6] It is represented by. Here, the second term on the right side is the true coincidence count S measured because the γ-ray absorber could not be completely absorbed or the detector on the opposite line could not be hidden by the γ-ray absorber.
It is a part of ₀ (x, θ). α and β are geometric factors obtained by calculation or measuring a phantom. B _ac
(X, θ) is the coincidence coincidence count. Using the measured value of B _ac (x, θ) and removing the scatter coincidence count again, B (x, θ)

[Equation 7] From Eqs. (3) and (6), the scatter coincidence count and the true coincidence count can be calculated by the following equation.

[Equation 8] This true coincidence count is used to obtain a correct image without spurious coincidence counts.

(2) Three-dimensional image reconstruction The above-mentioned scattering coincidence counting method also relates to two-dimensional image reconstruction.
However, for 3D image reconstruction, FIG.
become that way. In other words, the rod-shaped γ-ray absorber 26 is
The same clock of γ-ray from the subject by rotating concentrically with the ring 6.
Make a number. Since coincidence counting and scatter coincidence counting are three-dimensional
For example, S (x, θ, N₁, N₂) And B _SC(X, θ, N
₁, N₂), Which correspond to (8) and (9)
The formula can be expressed by the following formula.

[Equation 9] Here, N ₁ and N ₂ represent the ring numbers of the detector. θ is from 0 to 360 degrees. When N ₁ = N ₂ , θ is 0 to 180 degrees.

FIG. 7 shows the details of the gantry portion provided with the detector ring 6 of the positron tomography apparatus used in the present invention shown in FIG. 1. In the present invention, the detector ring 6 of the positron tomography apparatus is shown. The γ-ray absorber 26 is supported inward by a rotating body 27 like a cantilever, and the γ-ray absorber 26 is supported inside the detector ring 6 so as to rotate concentrically therewith. While rotating the γ-ray absorber 26, the position of the γ-ray absorber 26 is measured by the angle encoder 28, and a large number of detectors No. 1 to N arranged in the detector ring 6 in a multiplicity of circles at 360 degrees. Simultaneous counting circuit 16
The position on the circumference of the γ-ray absorber is sent to the coincidence counting circuit 16 together with the position information of the γ-ray absorber from the position encoder 28, where the coincidence counting and the scattering coincidence counting are performed. Since the simultaneous measurement data is collected by the simultaneous data collection processor system 7 and the obtained data is processed by the data processing system 8, the data is classified, memorized, and the image is reconstructed. is there.

In the present invention, the γ-ray absorber 26 is an alloy containing lead as a main component (eg 90% lead-10% tantalum alloy) or an alloy containing bismuth as a main component (eg 90% bismuth-10).
% Tantalum alloy) is particularly effective. Further, in order to maintain the mechanical strength of the γ-ray absorber 26, the γ-ray absorber 26 can be formed by filling a stainless pipe (thickness 0.1 to 0.2 mm) with lead.

When mounting the γ-ray absorber 26 on the rotating body 27, the mounting position of the γ-ray absorbing body 26 is adjustable in the radial direction of the rotating body 27. That is, the γ-ray absorber 26
Is made to be able to approach or separate according to the size of the subject (for example, the head, chest, abdomen of the living body, etc.).

FIG. 8 shows the rotation position of the γ-ray absorber 26 in the detector ring 6 provided with the γ-ray absorber according to the present invention. The γ-ray absorption connecting the pair of detectors 12-N ₁ and 12-N ₂ The coordinates of the position of the γ-ray absorber are shown in FIG.
As shown in (B), the position encoder 28 of the γ-ray absorber
Can be measured by When the γ-ray absorber 26 comes on the line connecting the pair of detectors 12-N ₁ and 12-N ₂ , if there is a positron at the point 29 of the subject 30 containing the positron, the positron included in the subject Is the electronic subject (for example, the tissue of the affected area)
Colliding with the detector 12-N ₁ which arranges γ rays at 180 degrees.
Even if it emits in the direction of 12-N ₂ , one of the detectors 12-N ₁
In direction it is absorbed by colliding with the gamma ray absorber 26, only the other of the detector 12-N gamma rays toward the direction of the _two detectors 12-N ₂
Detected in. Therefore, the true same clock value (A) cannot be measured, but the scattered same clock value (B) shown in 31 is
It can be measured with 12-N ₁ and 12-N ₂ . Not only in the plane of the ring, but the detector ring 6, as shown in FIG.
Like the 3rd and 4th layers, multiple layers are arranged in parallel and arranged in a circular ring shape, so that the three-dimensional scattering same clock values coming from the electrons of the adjacent ranks are also measured.

In the conventional positron emission tomography apparatus, as described above, a lead collimator made of lead or a slide shield plate called a septum is inserted between the respective detector rings so that the scattering same clock numerical value can be obtained. I tried to prevent the generation of false images (that is, measurement errors) due to simultaneous measurement, but as mentioned above, when this septum is provided and collimation is performed,
Although the scatter coincidences can be blocked and removed to some extent, the scatter coincidences from within the detector ring plane cannot be completely eliminated and the measurement sensitivity is reduced.

In the present invention, in order to increase the measurement sensitivity and produce clear image information, the lead collimator such as a sceptor is used to stop the gap between the detector rings, and the circular coordinates of the detector rings are arranged on the circular coordinates. The position can be measured according to the movement of the γ-ray absorber, and the true same-clock numerical value is obtained by subtracting the sum of the true same-clock numerical value R and the scattered same-clock numerical value S and the measured value of only the scattering coincidence count S. R is required.

That is, when the γ-ray absorber 26 passes on the line opposite to the detectors 12-N ₁ and 12-N ₂ , these two detectors 12-
N _1, 12-N coincidence by ₂ is a scatter coincidence.

For example, when performing positron CT of the head, the diameter of the γ-ray absorber is reduced so that it rotates around the subject, and for example, the chest or abdomen of the living body is positron C.
When measuring with T, it is necessary to increase the diameter and measure. Therefore, the fixed position that supports the γ-ray absorber can be moved in the radial direction of rotation of the γ-ray absorber, the fixed position of the γ-ray absorber can be changed, and the γ-ray absorber can be rotated as close to the subject as possible. It is important to configure so that the radius can be changed. This is to improve the sensitivity in measurement and to prevent the formation of false images as much as possible.

FIG. 9 shows a partial cross section of the gantry of the positron emission tomography apparatus of the present invention. In FIG. 9, reference numeral 5 denotes a subject of a living body to be inspected. The subject 5 is placed on a pedestal 53 so that the subject 5 can move in and out of the circular cavity of the detector ring 54 of the gantry unit 32, and the rotation with the angle encoder 28 attached. The rod-shaped γ-ray absorber 26 is projected and supported on the support portion 56 in a cantilever shape so as to rotate on a concentric circle inward of the detector ring 6. In the device of the present invention, the simultaneous coincidence of scattering is measured even between the respective layers without providing a gamma ray blocking plate such as a lead collimator (septum) between the pairs of the detectors constituting the detector ring 6 and arranged in multiple layers. In this way, the scattered same-clock value is subtracted from the sum of the true same-clock value and the scattered same-clock value to obtain the true same-clock value.

Conventionally, the threshold energy of the γ-ray detector is set to be high in order to remove the scattering coincidence count as much as possible. Therefore, all the energy of the γ-rays entering the detector is not necessarily converted into an electric signal, so that even a γ-ray having a threshold energy or more is below the threshold and no signal is output. Therefore, the true coincidence count may be removed, resulting in a decrease in detection sensitivity.

According to the method of the present invention, since the scattering coincidence count can be measured irrespective of the energy of γ-rays, the threshold value can be lowered and detection can be performed, so that the sensitivity of the apparatus is increased.

Whereas the conventional X-ray CT apparatus and MRI-CT apparatus give a morphological image, the positron emission tomography apparatus of the present invention is an epoch-making apparatus capable of giving a functional image of a living organ. is there. However, quantitative data must be collected to obtain functional images. What was left behind due to the problem of the quantitative nature of the image is how to eliminate the false image due to the scattering coincidence counting. The present invention proposes a measuring method and apparatus which solves this problem and which can remove false images and obtain only true same-clock numerical values by a method that can be efficiently and easily performed.

In order to reduce the exposure to the body and to obtain an image with high statistical accuracy, a three-dimensional image reconstruction that is several times to ten times more sensitive than a conventional positron emission tomography apparatus for two-dimensional image reconstruction is performed. A positron emission tomography device capable of being developed has been developed. However, since this device collects three-dimensional data and obtains a three-dimensional image, the scattering coincidence count appears as a false image (error) larger than that in the two-dimensional case, and a major problem that the quantitativeness of the image cannot be expected. have. By using the present invention, this problem is solved. Therefore, according to the present invention, the measurement accuracy of the positron emission tomography apparatus for three-dimensional image reconstruction can be dramatically improved astronomically, which greatly contributes to its development. Is expected to take place.

[0042]

EXAMPLE A positron emission tomography apparatus (hereinafter referred to as positron CT) of the present invention comprises a detector for detecting γ-rays, a scanning mechanism for moving the detector to obtain a sufficient number of data, and a simultaneous counting. Electronic circuit for transferring the data to the computer, processing the sent coincidence count data,
It is composed of a computer system for displaying images.

(1) Detector As the γ-ray detector used in positron CT, a scintillation detector using an inorganic crystal as a scintillator can be used. As shown in FIG. 10, the scintillation detector 22 includes a scintillator 22A that converts γ-rays into visible light and a photomultiplier tube 22B that converts visible light into an electric signal. In Figure 10
22G is a photocathode, 22C is a preamplifier, 22D is a pulse output terminal, and 22E is a high voltage input terminal.

In positron CT, the energy of γ-rays handled is higher than that of ordinary nuclear medicine equipment, and high speed is required for simultaneous counting. Therefore, in addition to NaI (Tl), which is often used as a γ-ray detector, Bi ₄ Ge ₃ O
₁₂ (BGO) and CsF are used. These characteristics are shown in Table 1. BGO is 511 keV compared to NaI (Tl)
It has a high absorption coefficient for photons (about 2.7 times), and is therefore highly sensitive.

[0045]

[Table 1]

Further, even if the size of the crystal is made small, the sensitivity does not decrease so much, so that a crystal having a small size can be used and the spatial resolution can be improved. However, since the specific fluorescence output is small and the light emission decay time is relatively long, the width of the time window for simultaneous counting cannot be made very small. The width of the time window can be around 2-20 ns.

CsF is characterized by a very fast fluorescence decay time of 5 ns. Therefore, the time resolution of the detector is
It can be about 0.5 ns. However, in such a high-speed detector, the flight time of annihilation photons must be taken into consideration, and the width of the time window cannot be smaller than about 2 ns. Even so, the width of the time window is
Since it is about 1/5 to 1/10 of that of BGO, it is possible to reduce the influence of coincidence coincidence counting and improve high counting characteristics. Moreover, since the time resolution of the detector itself is smaller than the flight time of γ-ray photons, it is possible to measure the time difference of flight between two γ-ray photons, and the positron positron emission nuclide (RI) distribution that does not depend on the CT principle can be measured. The possibility of measurement (time of flight method-TOF method) is considered. However, the absorption coefficient of CsF is about NaI (Tl), which is not suitable for making the crystal small and realizing high spatial resolution. Scintillators for positron CT are currently being divided into BGO for high-sensitivity and high-resolution devices and CsF, which has excellent high count rate characteristics and is capable of the TOF method.

A photomultiplier tube is connected to the scintillator and converts generated light into electric pulses. X-ray C
Other photoelectric conversion elements (photodiodes, etc.) used in T and the like cannot be used because they are inferior to the photomultiplier tube in terms of high speed and fidelity required for coincidence counting.

(2) Electronic Circuit In normal γ-ray measurement, the electric pulse generated by one γ-ray photon may be counted in a time of about 1 μs (10 ⁻⁶ s). However, in the positron CT apparatus, it is necessary to perform coincidence counting, and the electronic circuit is required to have extremely high accuracy and high speed. The conditions are particularly severe when using Bi ₄ G ₃ O ₁₂ (BGO), which has a low fluorescence efficiency, as a scintillator or when measuring the time difference of flight using CsF. The former needs to be accurate enough to be able to distinguish each photoelectron pulse (the pulse due to γ-ray photons is this cluster). In the latter case, 0.1 ns (1
ns requires an accuracy of the order of 10 ^-9 s).

FIG. 11 is a block diagram showing a conceptual diagram of an electronic circuit of the positron CT apparatus. The signals generated by the detectors 12A and 12B enter the preamplifier 33.
Normally, the output signal of the photomultiplier tube 29 is sufficiently large that it does not need to be amplified. However, since the output resistance is generally large, the preamplifiers 33A and 33B are used as impedance converters to reduce the output resistance.

The time pick-off circuits 34A and 34B, the delay circuits 35A and 35B, and the energy discriminating circuit 36A, which are arranged in parallel with the time pick-off circuits 34A and 34B, respectively, are used to identify the time when the pulse signals are output from the preamplifiers 33A and 33B. 36B and these gate circuits 37A and 37B gate the time signal and send it to the coincidence counting circuit 38. It is confirmed that the pulse is not due to noise but is due to 511 keV γ-ray photons. Judgment is made by the discrimination circuits 36A and 36B. The energy discriminating circuits 36A and 36B are performed by integrating the pulse signals and selecting only the signals whose integrated value is a certain value or more, and the fluorescence decay time (300 n in the case of BGO).
It takes about s). On the other hand, the time signal is generated by the time pick-off circuits 34A and 34B at the same time when the pulse rises. Since the energy discrimination has not been completed at the stage when the time signal is generated, the time signal is delayed.
Receives a certain delay due to 35A and 35B. If the energy reaches the threshold value after a certain period of time, the gate
37A and 37B are opened, and the signal enters the coincidence counting circuit 38. On the other hand, when the energy is below the threshold value, the gate is not opened and the subsequent processing is not performed. Time pickoff circuit 34
The decomposition time in A and 34B is an important factor that determines the time window width of the coincidence counting circuit 38, and it is necessary to make it as short as possible. In the case of BGO, since the specific fluorescence efficiency is small and the fluorescence decay time is long, only one photoelectron pulse is generated every few ns. Therefore, when the time signal is generated only by the first photoelectron pulse, the time pickoff circuits 34A, 34B
The decomposition time of is determined by the fluctuation of the arrival time of the pulse and is a minimum of 2 to 3 ns. In the case of CsF, since the photoelectron pulse is generated very early, the decomposition time of the time pickoff circuit can be 0.5 ns or less.

In the coincidence counting circuit 38, the two detectors 12A,
It is judged whether the time signals from 12B come at the same time. If they come at the same time, the address encoder 39
Send a signal to. However, the time signals from the γ-rays due to positron annihilation do not enter the coincidence counting circuit 38 at exactly the same time. Therefore, the coincidence counting circuit 38 determines that the signals that have entered within a certain period of time are "simultaneous". This time is called the time window width. In the case of BGO, the time window width needs to be 10 to 20 ns due to the decomposition time of the time pickoff circuit and the error of the delay circuit. On the other hand, in the case of CsF, mainly about 2 ns is required due to the flight time difference.

In principle, the coincidence counting circuit 38 is required for the number of pairs of the detectors 12, and the outputs thereof are all connected to the address encoder 39. Address encoder 39
Outputs the address of the pair of detectors 12 that have been simultaneously counted as a binary number and sends it to the computer 40 shown in FIG. However, if the coincidence counting circuit 38 is added to each detector pair, the number becomes enormous. Therefore, in actuality, in order to save the coincidence counting circuit 38, the detectors 12 may be divided into some groups, and the coincidence counting may be performed for each group.

(3) Computer system The computer system of the positron CT system is an X-ray C
It is very similar to the computer system attached to the T-apparatus and the scintillation camera, and has both features.

FIG. 12 is a conceptual diagram of a computer system of the positron CT system. Gantry department
The detector 41, an electronic circuit (such as that shown in FIG. 11) are built in 41, and a scanning mechanism is also added. The gantry unit 41 outputs the address of the detector pair that is simultaneously counted.

In the positron CT, it is necessary to collect a large amount of data at high speed, and as described below, it is necessary to convert detector addresses and the like when collecting data. Therefore, it may not be possible to cope with the instruction execution speed of an ordinary general-purpose computer, and the address information output from the gantry 41 is temporarily stored in the external buffer memory 42.

The measurement data temporarily stored in the buffer memory 42 and subjected to processing such as address conversion is sent to the computer 40 at the same time as the end of data collection. The measurement data is subjected to some pre-processing in the computer 40 and then converted into a positron CT image as shown in FIG. 15 by the above-mentioned calculation method. This method is called image restoration, but a large amount of calculation processing is required for image restoration. Therefore, the high-speed arithmetic unit 43 is added in order to perform the image restoration calculation in a short time, and if the matrix size is 128 × 128, the calculation can be performed in units of seconds. Even if a dedicated arithmetic unit is not provided, it is necessary to be able to calculate in minutes at a minimum.

The restored image is displayed on the image display device 44 and used for diagnosis. A CRT monitor (CRT) is usually used as the image display device 44, and in the case of black and white (gray scale), it has a gradation recognition capability of about 16 to 32 levels. Similar to the X-ray CT, the gradation of the display device is controlled by the window and the level.

The displayed image is a multi-format
It is recorded on the film with a camera. Measurement data and image data are temporarily stored in the magnetic disk 46 and are immediately called from the computer 40 when necessary. The magnetic tape device 47 records and stores these data. Therefore, the image data is duplicated and stored on the film in a form that can be converted so that it can be seen by human eyes and on the magnetic tape in a form understandable by the computer 40. The console 48 is for the operator to control the entire system, and includes a keyboard for interaction, a character display, dedicated function keys, and the like.

Data Processing: Data processing in the Positron CT apparatus is all performed by a computer system attached to the apparatus, and consists of data collection, preprocessing, image restoration, restoration image processing, file management and the like. File management is important to efficiently store and use image data and projection data. (1) Data Collection Data collection refers to conversion and storage of coincidentally counted detector addresses sent from an electronic circuit in the gantry 41 into a form convenient for computer processing. As in the case of a normal scintillation camera, there are two possible data acquisition methods: list mode acquisition and frame mode acquisition.

The list mode acquisition is a method of accumulating coincidence counting addresses coming out of the gantry as they are in the magnetic disk via a computer. The address data is converted into projection data by the computer after the data collection is completed. In this method, high-speed collection can be performed because no processing other than data transfer is performed during data collection, and data conversion is flexible because it is collectively performed after software collection. However, since all the address data is stored in the disk, it is impossible to handle a large number of counts.

On the other hand, in the frame mode acquisition, each time the address data comes out of the gantry, it is converted into the coordinates of the measuring system and 1 is added to the contents of the memory corresponding to the coordinates. Simultaneously with the end of data collection, projection data is obtained as a two-dimensional array in the memory. The memory size required for frame mode acquisition is the frame size (128 × 12
8)), and it does not depend on the count value, so a large amount of data can be collected. However, it is necessary to perform a relatively complicated data conversion process at the same time as the data collection, and it takes a considerable amount of time to perform it by software. For example, in the case of the conventional device, the time for collecting one simultaneous count is about 30 μs, and the upper limit of the count rate that can be collected is determined by this.
This drawback is overcome by using a large-capacity and high-speed buffer memory and performing data conversion processing by dedicated hardware. In this case, one count can be collected in 1 μs as the designed value.

Finally, the outline of the conversion process from address data to projection data will be described for the positron emission tomography apparatus of the present invention having a ring type detector. FIG. 13 shows an example of conversion from address data to measurement system coordinates.
In the figure, the detector 12k and the detector 12l simultaneously count. The address pair (k, l) is sent to the computer 40. In the computer 44, first the two detectors 12 from (k, l)
The coordinates (t _kl , θ _kl ) of the straight line connecting k and 12l are obtained. Next, the coordinates (t, θ) of the measurement system are obtained by adding position information (rotation angle, etc.) of the entire system of the detector 12, and 1 is added to the corresponding two-dimensional array. When this conversion processing is performed on all address data obtained during data collection, each element of the array is the integrated value (line sum) of the positron-emitting nuclide (RI) distribution observed on the straight line (t, θ). Becomes In the list mode collection, such processing is performed after the data collection, whereas in the frame mode collection, it is performed in parallel with the collection.

(2) Pre-processing Generally, projection data directly measured from a subject includes variations in detector sensitivity, absorption of γ-rays in the subject, coincidence coincidence counting, scatter coincidence counting, etc. The projection is very different from the ideal projection obtained by simply integrating the positron emission nuclide (RI) distribution. The preprocessing is to remove such distortion of the measurement data. The pre-processing is performed on the measured projection data and is distinguished from the post-processing (image processing in a narrow sense) performed on the image data.

There are many possible types of detector arrays and scanning methods for positron CT. In order to improve the image quality and improve the quantitativeness, it is necessary to perform pre-processing suitable for each device. Therefore, the pretreatment method and order differ slightly depending on the device.

The scatter coincidence has a great influence on the quantitativeness of the restored image, but its removal method has not been established at present. Conventional methods are: 1) Increase the threshold for energy discrimination and remove it by hardware. 2) The measurement data at the edge of the field of view (where the subject does not exist) is subtracted as the scattered ray component, which cannot be said to be perfect.

In the present invention, the γ-ray absorber is rotated concentrically with the inside of the detector ring so that the scattered and same clock values can be measured separately. It can be said that this is an intentional invention in that the above can be separated.

On the other hand, the coincidence coincidence count can be estimated by hardware. That is, if a coincidence counting circuit in which the positions of the time windows are shifted not by "simultaneous" but by a fixed time is created, all the counts here are measured by random coincidences. Therefore, in order to remove the coincidence coincidence count, this count may be obtained separately and subtracted from the output of the normal coincidence count circuit.

The γ-ray absorption in the object and the variation in the sensitivity of the detector are corrected by the transmission positron emission nuclide data measured by the external radiation source and the blank data for calibration. Now, the positron emitting nuclide (R
I) is measured and the measured data is emission data E (t, θ), and the transmission data is T
(T, θ), and the blank data is B (t, θ) (Fig.
14). In FIG. 14, 49 is an external radiation source, 50 is a subject,
Reference numeral 51 represents a subject to which a positron emitting nuclide (RI) was administered. Here, θ is the projection angle, and t is the position coordinate of the line sum. These measurement data are the true line sum P (t, θ), the absorption A (t, θ) of the subject, and the dispersion D of the sensitivity of the detector.
(T, θ) and projection R (t) of the external source

[Equation 10] It is expressed like. Here, the external radiation source 49 is usually circular, and its projection R does not depend on the projection angle θ.

The true line sum P (t, θ) is obtained by the equations (12) to (14)
Than

[Equation 11] As described above, the measurement data (t, θ), T (t, θ) and the radius of the external radiation source 49 are represented by R (t). As described above, one of the remarkable features of the positron CT is that the absorption correction can be performed only by the transmission data.

However, the statistics of T (t, θ) are generally very poor, and if P (t, θ) in Eq. (15) is restored as it is, the resulting image quality becomes very poor. So T
After calibrating (t, θ) with B (t, θ), smoothing processing is required. An image smoothed by this pre-processing is obtained, and the image quality is significantly improved.

(3) Image Restoration Image restoration is performed by the same method as the image restoration of X-ray CT. This calculation method can be divided into two processes: convolution integral and backprojection.

The back projection is performed by returning the projections to their original directions as shown in FIG. The distribution created by this operation is called a backprojection image. The backprojection image is not the original image itself, but a blurred version of it. In order to obtain a correct image without blurring, it is only necessary to subject the projection data to a certain high frequency enhancement process as shown in FIG. This is done mathematically by superimposing the filter function (see Eq. 1). The correct distribution can be obtained by backprojecting the projection that has been superposed and integrated. When the shape of the correction function is changed, the contour of the image is emphasized to increase its sharpness, and conversely smoothed to obtain a smooth image.

(4) Post-processing of images Performing various processes on a completed image by a computer through operations on the console is called post-processing or image processing. Post-processing is performed with other image equipment (X-ray C
T, scintillation camera, etc.), but some are unique to positron CT. There are an extremely large number of types of image processing, but only typical ones are given here.

1. Processing common to other images (processing for a single image) Image display (level and window adjustment) Image enlargement / reduction Image translation, rotation, left / right (vertical) inversion Image smoothing , Contour enhancement ROI (region of interest) setting Measurement of statistical amount (average value, variance, area, etc.) in ROI (processing for multiple images) Multi-frame display of multiple images Superimposition display of multiple images (images of different nuclides are displayed) Calculations between images (addition, subtraction, multiplication and division, and multiplication by a constant) Graphs of changes in statistics of ROIs between images (especially dynamic curve is obtained from images that change with time)

2. Processing peculiar to positron CT image (processing on a single image) A glucose metabolism rate distribution is obtained from an ¹⁸ F-fluorodeoxyglucose ( ¹⁸ FDG) image. (Processing for Multiple Images) An oxygen metabolism rate distribution is obtained from the local blood flow (C ¹⁵ O ₂ image) and the ¹⁵ O ₂ image. The metabolic rate constant of deoxyglucose is determined from ¹⁸ FDG images that change with time. In positron CT,
Due to the importance of biochemical research, this type of image processing is required more than other imaging equipment. In particular, image processing peculiar to positron CT images is expected to be further improved in the future as research progresses.

[0077]

According to the conventional method, the threshold energy of the γ-ray detector is increased to measure the scattering coincidence count in order to remove it as much as possible. However, since all the energy of the γ-rays entering the detector is not converted into an electric signal, even a γ-ray having a threshold energy or more is below the threshold energy and no signal is output. Therefore, the true coincidence count may be removed, resulting in a decrease in detection sensitivity.

According to the present invention, by rotating the γ-ray absorber inside the detector ring so as to form a concentric circle with the γ-ray absorber, it is possible to measure the scattering coincidence count regardless of the magnitude of the γ-ray energy. The sensitivity of the device is very high because it can be detected by lowering it. Further, while the conventional X-ray CT apparatus and positron tomography apparatus (MRI-CT) give a morphological image, the positron tomography apparatus of the present invention is an epoch capable of giving a functional image of a living organ. Device.
However, quantitative data must be collected to obtain functional images. Due to the problem of quantitativeness of this image, what was left behind was a false image due to the coincidence of scattering.
The present invention is a method capable of solving this problem by preventing the generation of the false image and efficiently and easily. It is possible to give quantitativeness to the image of a positron imaging apparatus for two-dimensional image reconstruction which is mainly used at present.

Further, in order to reduce the exposure in the living body and obtain an image with high statistical accuracy, it is possible to perform three-dimensional image reconstruction which is several times to ten times more sensitive than the conventional positron imaging apparatus for two-dimensional image reconstruction. A further improved version of the positron emission tomography system can be developed. However, this 3D image reconstruction device
Since three-dimensional data is collected and three-dimensional images are acquired, the simultaneous coincidence of scattering appears as a larger error than in two-dimensional, and there is a big problem that the quantitativeness of images cannot be expected. This problem was solved by rotating the γ-ray absorber of the present invention on a concentric circle inside the detector ring, and measuring the data of the position and the same clock numerical value, so that the positron for three-dimensional image reconstruction was solved. The improvement and development of the tomography apparatus have an industrially great effect, which will be carried out more and more from now on.

According to the method of the present invention, since the sensitivity is increased, the three-dimensional image reconstruction method is improved so as not to generate a false image (error), so that the sensitivity is improved several tens of times. Development becomes easier. By using the positron emission tomography apparatus of the present invention, in the liver function test, not only is it used for the detection of neoplastic lesions, but it is also possible to improve the ability to detect deep tumors,
It has a remarkable medical effect when applied to liver dynamics imaging, brain tumor, cerebral infarction, schizophrenia, meningioma and the like.

[Brief description of drawings]

FIG. 1 is a system explanatory view of a conventional positron emission tomography apparatus.

FIG. 2 is a diagram for explaining the principle of a conventional positron emission tomography apparatus.

FIG. 3 is a diagram for explaining the principle of a conventional positron emission tomography apparatus.

FIG. 4 is a diagram for explaining the principle of a conventional positron emission tomography apparatus.

FIG. 5 is a diagram for explaining the principle of a conventional positron emission tomography apparatus.

FIG. 6 is a sectional view of a detection ring of a conventional positron emission tomography apparatus.

7A is a system explanatory view of a positron emission tomography apparatus of the present invention, and FIG. 7B is a circuit diagram showing an outline of a detector ring and a simultaneous measurement circuit of the apparatus.

FIG. 8 is a principle explanatory view showing a measurement principle of scatter coincidence counting according to the present invention.

FIG. 9 is an explanatory diagram of a gantry unit of the positron emission tomography apparatus according to the present invention.

FIG. 10 is a cross-sectional view showing an example of a scintillation detector used in the device of the present invention.

FIG. 11 is a block diagram of an electronic circuit of the device of the present invention.

FIG. 12 is a conceptual diagram for schematically explaining the computer system of the device of the present invention.

FIG. 13 is an explanatory diagram showing a method of converting address data into measurement system coordinates in the device of the present invention.

FIG. 14 is an explanatory diagram showing (a) blank data, (b) transmission data, and (c) emission data of measurement data in the device of the present invention.

FIG. 15 is an explanatory diagram showing a method of creating a two-dimensional image by the back projection method according to the device of the present invention.

FIG. 16 is an explanatory diagram showing a correction characteristic of projection by weight integration in the device of the present invention.

[Explanation of symbols]

1 Conventional positron tomography apparatus 2 Cyclotron 3 Automatic compound synthesizing device 4 Positron labeling drug 5 Subject 6 Detector ring 7 Data acquisition system 8 Data processing system 9 Positron 10 Electron in living body 11 γ ray 12 γ ray detector 12N ₁ , 12N _{2 A} pair of detectors 12A, 12B A pair of detectors 12k, 12l A pair of detectors 13 Time pick-off circuit 14 Simultaneous counting circuit 15 γ-ray detection unit (gantry unit) 16 Simultaneous counting circuit 17 Coincidence line determination circuit 18A Sinogram collection Memory 18B Image reconstruction processor 18C Image display 18D Computer CPU 18E Other peripherals 19 Positron 20 Subject 21 Other positron 22 Scintillation detector 22A Scintillator 22B Photomultiplier tube 22C Preamplifier 22D Power Output terminal 22E High voltage power terminal 22F γ-ray 22G Photocathode 23 Collimator (septum, Pb shielding plate) 24 Inter-layer simultaneous counting 25 In-layer simultaneous counting 26 γ-ray absorber 27 Rotating body 28 γ-ray absorber position encoder 29 Positron 30 Subject 31 Other positron 32 Gantry part 33A, 33B Preamplifier 34A, 34B Time pick-off circuit 35A, 35B Delay circuit 36A, 36B Energy discrimination circuit 37A, 37B Gate circuit 38 Simultaneous counting circuit 39 Address encoder circuit 40 Computer 41 Gantry 42 buffer memory-43 high-speed arithmetic unit 44 image display unit 45 multi-format camera 46 magnetic disk 47 magnetic tape 48 consol 49 external source 50 subject 51 positron emitting nuclide 52 projection 53 mount 5 Detector ring 55 rotating body 56 rotating support 56M Motor

Claims (4)

[Claims] 1. A detector of a positron emission tomography apparatus and a γ-ray absorber positioned in a line opposite to the detector and supported by an appropriate supporting device, and the γ-ray absorber rotates on a concentric circle surrounding an object. Of the γ-rays emitted from the positrons in the subject while excluding the true same-clock numerical value and measuring only the scattered same-clock numerical value, simultaneously counting the true same-clock numerical value and the scattered same-clock numerical value, A simultaneous coincidence counting measurement method using a γ-ray absorber in a positron emission tomography apparatus, which is characterized by obtaining a true coincidence count from the difference between the two. 2. A positron emission tomography apparatus for a subject in which γ-rays emitted from positrons in a subject are detected by a large number of pairs of γ-ray detector rings. A positron emission tomography apparatus, characterized in that a rotator equipped with a γ-ray absorber located on one side is provided, and the γ-ray absorber is rotated inward of the detector ring on a concentric circle. 3. The positron emission tomography apparatus according to claim 2, wherein the γ-ray absorber is made of an alloy containing lead as a main component. 4. The positron emission tomography apparatus according to claim 2, wherein the γ-ray absorber is made of an alloy containing bismuth as a main component.