JPH04331396A - Radiation detector - Google Patents

Abstract

PURPOSE:To exactly measure radiation quantity from a radiation source by detecting efficiently the radiation from a small amount of measuring sample including radiation source. CONSTITUTION:A scintillator unit 1 is constituted of more than three scintillators having at least one plane side wall and the radiation emitted from a measuring sample put in its hollow part goes in the plane side wall constituting the hollow and is detected as scintillation light by each scintillator having the plane side wall. Each scintillator is connected to the neighbouring scintillator with no gap to enable it to detect all the radiation emitted by the measuring sample without leakage. Thus the radiation detection count by the scintillation unit 1 is increased without increasing the concentration of the radiation source. The detected radiation is converted to electric signal by a photo-electric multipier 4, concurrence of the detection is decided by a concurrence decision unit 6 and is processed as the data indicating the concentration of the radiation source.

Description

[Detailed description of the invention]

0001

[Industrial Application Field] The present invention relates to a radiation detection device.
In particular, the present invention relates to a radiation detection device capable of measuring time changes in blood RI (radionuclide) concentration distribution with high accuracy.

0002

[Prior Art] Positron CT is used as a device to detect physiological and biochemical changes and metabolic functions within a living body through external measurements.
The device is known. Positron CT equipment is a system in which a drug labeled with a nuclide that emits positrons (positron emitting nuclide) is injected into a patient's body using a syringe, or a gasified drug is injected into the patient's body by inhalation. By detecting radiation from the body, the distribution of radionuclides in the body can be detected as a cross-sectional image. Positrons emitted from radionuclides (positron-emitting nuclides) taken into the body lose kinetic energy in their vicinity, then combine with electrons that make up the material and annihilate, sending a pair of annihilation gamma rays in opposite directions. released into the By detecting this pair of annihilation gamma rays with a measuring device placed outside the body, it is possible to quantitatively measure the concentration distribution of the radionuclide that is the radiation source of the annihilation gamma rays.

As described above, a radionuclide administered into the body emits a pair of annihilation gamma rays, and as a result of this emission, the concentration of the radionuclide decreases over time. Therefore, in order to accurately measure the concentration distribution of radionuclides in the body using a positron CT device, it is necessary to separately detect accurate temporal changes in the blood RI (radionuclide) concentration. As an RI measurement device for measuring changes in blood RI concentration over time, it is arranged to sandwich a catheter for collecting internal blood containing radionuclides to the outside of the body and a part of the blood supply tube attached to the catheter. 2 scintillators that generate scintillation light by detecting annihilation gamma rays emitted from blood RI (radionuclide) flowing in the blood supply tube, and a scintillator that is optically coupled to each scintillator and generates scintillation light. Some use photomultiplier tubes that convert light into amplified electrical signals. Among these, a straight blood supply tube sandwiched between two scintillators was developed by Moses (IEEE TRANSACTIONS ON) as shown in Figure 7.
NUCLEAR SCINENCE, VOL 37,
No. 2, APRIL 1990 P580-
584), and the U-shaped structure of the blood supply tube sandwiched between two scintillators as shown in Figure 8 was developed by Ericsson (IEEE TRANSAC).
TIONS ON NUCLEAR SCIENCE,
VOL 35, NO. 1, FEBRUARY 1
988, pp. 703-707), respectively.

[0004] In the Moses and Ericsson RI measuring devices, a coincidence determination process is performed using a coincidence circuit on the detection output of annihilation gamma rays from a scintillator. Through this simultaneous determination process, for example, background noise is removed by removing the influence of environmental radiation, and the S/N ratio of the measurement is increased.

[0005]

SUMMARY OF THE INVENTION It is desirable that the amount of radionuclides used for positron CT measurements administered into the body be small in order not to compromise the safety of biological systems. In order to accurately measure time changes in RI concentration using a small amount of radionuclides, the amount of blood supplied between the scintillators (pool amount) for RI measurement must be increased, and the amount of annihilated γ-rays detected by the scintillator must be increased. It is conceivable to increase the absolute amount of radiation.

[0006] In the Moses RI measuring device, the blood supply tube BT placed between the two scintillators S1 and S2 has a linear structure, and if the pool amount is small and the dose of radionuclide is small, the measurement accuracy will be reduced. It gets lower. On the other hand, in Ericsson's RI measurement device, two scintillators S1 and S
The blood supply tube BT placed between the two has a U-shaped structure, and has a larger pool amount than the Moses. However, as with Ericsson, the residence time of the measured blood in the annihilation gamma ray detection area (the gap between the two scintillators where the blood supply tube is placed) becomes longer (the flow path of the blood supply tube in the annihilation gamma ray detection area (longer), the accuracy of measuring changes over time will decrease. In addition, since it is impossible to return the collected blood to the living body, there is a problem that there is a limit to the amount of blood pool that can be measured.

[0007] Furthermore, in the Moses and Erickson RI measurement apparatus shown in FIGS. 7 and 8, two scintillators are arranged so as to sandwich the blood supply tube between them, and include the scintillators S1 and S2 and the blood supply tube BT. The cross section has a structure as shown in FIG. 9A or FIG. 9B. Figure 9
In case A, among the annihilation γ-rays emitted from, for example, the center of the blood supply tube BT in the detection area, the annihilation γ-rays emitted within angles d and d' (undetectable area) are not detected by the scintillators S1 and S2. There is a problem in that the amount of annihilation gamma rays detected by the scintillator becomes low because they are emitted in vain. In addition, in the case of FIG. 9B, all annihilation γ-rays emitted from the blood supply tube BT are detected by the scintillator, but
For example, when a pair of annihilation γ-rays are emitted in the directions of arrows A and B, each annihilation γ-ray will be detected by the same scintillator, and will be removed as background noise by simultaneous determination processing by the simultaneous measurement circuit. Put it away. Therefore, the annihilation γ-rays emitted at the same angle of incidence to the same scintillator are not apparently detected by the scintillators S1 and S2, similar to the annihilation γ-rays emitted to the undetectable area in FIG. 9A, and as shown in FIG. 9B. Even if the amount of detection is low,
Both Moses and Ericsson detect low amounts of annihilation gamma rays, making it impossible to improve measurement accuracy.

On the other hand, in addition to improving the measurement accuracy by increasing the absolute amount of annihilation gamma rays emitted, it is also possible to improve the measurement accuracy by increasing the detected amount of annihilation gamma rays emitted from the blood supply tube. Conceivable.

The purpose of the present invention was to solve the above-mentioned problems, and it is possible to detect radiation from a small amount of measurement sample containing a radiation source without waste, thereby removing radiation from the radiation source. The aim is to accurately measure radiation doses. In addition, the measurement sample is radioactive nuclide (RI
), it is possible to detect all of the annihilation gamma rays emitted from the measurement sample without waste, increasing the actual amount of annihilation gamma rays emitted from the measurement sample, thereby making it possible to measure temporal changes in blood RI concentration with high precision. The purpose is to make it measurable.

0010

[Means for Solving the Problems] In order to achieve this object, the radiation detection device of the present invention detects radiation from a measurement sample containing a radiation source at at least two positions, and at each detection position, scintillation light is emitted. A scintillator unit that generates light, a photoelectric conversion element that is optically coupled to the scintillator unit and converts the scintillation light generated by the scintillator unit into an amplified electrical signal, and a simultaneous counting process for the electrical signals of the photoelectric conversion element. In a radiation detection device comprising a simultaneous determination device that determines that radiation from a measurement sample is simultaneously incident on the two positions by The planar side walls of each scintillator form a cavity for placing the measurement sample, and the scintillators are connected to each other without any gaps so that the measurement sample is surrounded by approximately 360 degrees by the planar side walls. Features.

[0011]

[Operation] Radiation emitted from the measurement sample placed in the cavity of the scintillator unit is incident on the flat side wall forming the cavity, and is detected as scintillation light by each scintillator having the flat side wall. Each scintillator is connected to an adjacent scintillator without any gaps so that all the radiation emitted from the measurement sample can be detected without leaking.This allows the scintillator unit to emit radiation without increasing the concentration of the radiation source. The amount of detection can be increased. The radiation detected by the scintillator unit is converted into an electrical signal by a photomultiplier tube, and the simultaneity of radiation detection is determined by a coincidence determination device. If simultaneity is determined, it is determined that the radiation comes from the same radiation source, and the data is processed as data representing the concentration of the radiation source.

0012

DESCRIPTION OF THE PREFERRED EMBODIMENTS An example embodying the present invention will be described below with reference to the drawings.

The scintillator unit used in the present invention is composed of at least three scintillators, and has a structure in which a cavity with no gap is formed by one side wall of each scintillator. By flowing the measurement sample through the cavity or placing the measurement sample in the cavity, annihilation gamma rays emitted from the measurement sample are incident on one side wall of each scintillator facing the cavity and detected. do. The measurement sample is surrounded by the side wall of the scintillator almost 360° with no gaps except for the entrance/exit for inserting the measurement sample, so that almost all of the radiation-annihilated gamma rays can be detected by the scintillator. Furthermore, it is desirable that one side wall of each scintillator forming the cavity is flat. this is,
This is because if the side wall is a curved surface, a pair of annihilation gamma rays emitted from a measurement sample located near the side wall may both enter the same scintillator, and will apparently not be detected by the scintillator.

FIGS. 4A to 4E show various aspects of scintillator units used in radiation detection devices, and show cross sections including each scintillator and cavity that constitute the scintillator unit, viewed from the direction in which the measurement sample is inserted. be.

The scintillator units shown in FIGS. 4A to 4C use three, four, and six scintillators, respectively, so that the cross-sectional shape of the cavity (polygonal shape) and the outer cross-sectional shape of the scintillator unit are the same. scintillators are designed and arranged in The bonding surface CS of each scintillator with the adjacent scintillator has a completely joined structure, and the cavity 2 is surrounded by one side wall of each scintillator without any gap. If a scintillator unit with such a structure is used, all annihilation γ-rays other than the annihilation γ-rays incident on the bonding surface CS will be detected among the annihilation γ-rays emitted from the measurement sample, and the annihilation γ-rays by the scintillator will be detected. The amount of detection increases. Note that the number of scintillators is not limited to the above embodiment, and may be any number as long as it is three or more.

The scintillator unit shown in FIG. 4D has a structure in which four scintillators are assembled in a tortoise or spiral shape, and the scintillator unit shown in FIG. 4E has a pair of scintillators facing each other connected to the other pair. It has a structure in which four scintillators are arranged so as to be sandwiched between the scintillators. In either case, at least one of the bonding surfaces of at least one scintillator with an adjacent scintillator of the scintillators constituting the scintillator unit is on the same plane as the plane side wall of the scintillator. In this case, since each scintillator and the adjacent scintillator are joined without any gaps, it is possible to surround the measurement sample by 360° by one side wall of each scintillator, and the scintillator absorbs all annihilation γ-rays emitted from the measurement sample. can be detected. Furthermore, FIGS. 4D and 4
In the scintillator unit E, the annihilation γ-rays emitted in the direction of the bonding surface do not become undetectable due to the bonding surface and immediately enter the scintillator, so the annihilation γ-rays are more effectively detected than in the scintillator units shown in FIGS. 4A to 4C. can be detected. Note that the number of scintillators arranged as described above is not limited to the above embodiment, and may be any number as long as it is three or more.

The reason why the number of scintillators is set to three or more is that in order for the plane side walls of two scintillators that are always different in the radiation direction of annihilation gamma rays to exist in the cavity, the cavity must be formed by at least three scintillators. This is because it is necessary.

FIG. 2 shows an embodiment of the radiation detection apparatus of the present invention, and FIG. 1 shows a scintillator unit used in the radiation detection apparatus. In this example, a rectangular parallelepiped crystal (30
A structure in which four scintillators 1a, 1b, 1c, and 1d (mm x 45 mm x 30 mm) are arranged without gaps in a tomoe shape or a spiral shape with a cavity 2 in the center (Figure 4D
) scintillator unit 1 is used, and the scintillator unit 1 is covered with a lead shield LS to prevent gamma rays from leaking from the outside of the measurement system. Further, a blood supply tube 3 is used to flow blood containing radionuclides as a measurement sample.

A part of the blood supply tube 3 with one end attached to a blood collection needle is placed inside the cavity 2 of the scintillator unit 1, and the annihilation gamma rays emitted from the blood supply tube 3 inside the cavity 2 are absorbed into the cavity. four scintillators 1a forming part 2;
It is detected by being incident on each side wall (detection side wall) of 1b, 1c, and 1d. Therefore, the detection sidewall of each scintillator facing the cavity 2 forms an annihilation gamma ray detection region. Each scintillator 1a, 1b, 1c, 1d includes a photomultiplier tube (PMT) 4 supplied with a high voltage from a high voltage source.
a, 4b, 4c, and 4d, and a weak optical signal (scintillation light) from each scintillator is converted into a pulsed electrical signal amplified by a light multiplier. Pulse signals from each photomultiplier tube 4a, 4b, 4c, 4d are processed by preprocessing units 5a, 5b, 5c, which will be described later.
, 5d to the coincidence determination device 6, and the coincidence determination result is output to the counter 7. Further, the addition result by the counter 7 is output to a processing unit 9 consisting of a computer or the like, and a display or the like displays the change in RI concentration over time. Blood containing radionuclides from the blood collection unit flows through the blood supply pipe 3 by a blood collection pump (not shown) and is collected.

FIG. 3 is a block diagram of the circuits constituting the preprocessing units 5a, 5b, 5c, and 5d of the radiation detection apparatus shown in FIG.

[0021] Each pre-processing unit 5a, 5b, 5c, 5d
are a preamplifier 10, a time pickoff circuit 11, a delay circuit 12, and a gate circuit 13 connected to each photomultiplier tube.
, and an energy discrimination circuit 14 connected to a gate circuit 13 in parallel to a time pickoff circuit 11 and a delay circuit 12.

Each photomultiplier tube 4a (4b, 4c, 4d)
The pulse signal from is input to the preamplifier 10. Normally, the output signal of a photomultiplier tube is sufficiently large that there is no need for amplification, but since the output resistance is large, a preamplifier is used as an impedance converter. Noise signals are removed from the output pulse signal of the preamplifier, and annihilation gamma rays (511 keV
) It is necessary to perform energy discrimination processing in order to extract only the pulse signals caused by The energy discrimination process is accomplished by the energy discrimination circuit 14, which integrates the pulse signal and selects only those signals whose integral value is greater than or equal to a certain threshold. The output signal of the preamplifier is also input to a time pickoff circuit connected in parallel with the energy discrimination circuit 14 to form a time signal representative of the rise time of the pulse signal. At the stage when the time signal is generated, energy discrimination is not completed, so the time signal is subjected to a certain delay process by the delay circuit 12. After a certain period of time has passed, when the energy discrimination circuit 14 determines that the energy of the pulse signal has reached the threshold,
The gate of the gate circuit 13 is opened, and the time signal is input to the coincidence determiner 6 via the gate circuit, where it is subjected to coincidence counting processing. On the other hand, when the energy of the pulse signal is below the threshold value, the gate of the gate circuit 13 is not opened and subsequent processing is not performed.

The surface of each scintillator 1a, 1b, 1c, and 1d that connects with the adjacent scintillator is mirror-finished, and a reflector made of barium sulfate is further provided. Therefore, each scintillator is optically completely independent, and leakage of scintillation light to adjacent scintillators (optical crosstalk) is prevented.

Next, a method for measuring time changes in blood RI concentration using the above-mentioned radiation detection device will be described.

Blood containing radionuclides collected from a body artery using a blood collection unit is pumped into the blood supply tube 3 placed in the cavity 2 of the scintillator unit 1 at a rate of approximately 0.1 cc/min ( Flow at a flow rate of up to 1 cc/min). A pair of annihilation gamma rays emitted from the blood supply pipe 3 in the cavity enters two of the four scintillators (for example, 1a and 1c) and is converted into scintillation light. Each scintillation light generated by the scintillators 1a and 1c is transmitted to each scintillator 1a and 1c.
After being converted into electrical pulse signals by photomultiplier tubes 4a and 4c optically coupled to c, the pulse signals are sent to a simultaneous determination unit 6 via pre-processing units 5a and 5c that allow only pulse signals of a predetermined value or higher to pass through. is input. The simultaneity determiner 6 determines the simultaneity of the time signals from the preprocessing units 5a and 5c. That is, it is determined whether the annihilation gamma rays incident on the two scintillators 1a and 1c originate from the same source (radioactive nuclide). Specifically, signals input to the coincidence determiner 6 within a certain period of time (time window width) are determined to be simultaneous. The scintillator used is BG
For O, the time window width is preferably 10-20 nanoseconds. The signals determined to be simultaneous by the coincidence determiner 6 are immediately added up by the counter 7. This addition process is performed at regular intervals, and the processing unit 9 statistically calculates the average value at each cycle. The time change in the RI (radioactive nuclide) concentration is measured from the relationship between the count value obtained by the addition process of the counter 7 and time. The simultaneous measurement process by the simultaneous determination unit 6 removes the influence of environmental radiation, thereby removing background noise and increasing the S/N ratio of the measurement.

Furthermore, the photomultiplier tubes 4a, 4b, 4c, 4d
The optical coupling position with the scintillators 1a, 1b, 1c, and 1d can be any surface other than the side surface (detection side wall) facing the cavity 2 forming the annihilation gamma ray detection region. For example, the photomultiplier tube may be placed at a position perpendicular to the direction in which the blood supply tube is placed in the annihilation gamma ray detection area (
It is possible to arrange the photomultiplier tubes in a parallel position (FIG. 5A) or in a parallel position (FIG. 5B).

In this embodiment, the four scintillators are assembled in a tortoise shape or a spiral shape, and each scintillator is connected to the adjacent scintillator without any gaps, so that the blood supply pipe 3 in the annihilation gamma ray detection area The blood supply tube 3 can be surrounded by 360° by the detection side wall.
Almost all the annihilation gamma rays emitted from the scintillator can be detected by the scintillator. Furthermore, as is clear from FIG. 6, the blood supply pipe 3' (dotted line) has the inscribed diameter of the cavity 2.
By using the blood supply tube 3 (solid line) with a diameter smaller than the thickness of the binding surface CS, the annihilation gamma rays (arrow a), which had become undetectable due to incidence along the binding surface of the scintillator, can be removed. can also be detected (arrow b). In this case, the annihilation gamma rays (arrow c) emitted in the direction of the bonding surface will pass through the bonding surface and will always be incident on one of the two scintillators bonded via the bonding surface. It becomes possible to detect (completely detect) radiation annihilation gamma rays.

Furthermore, the scintillator unit 1 of the present invention can have a larger cross-sectional area or volume of the cavity, that is, the annihilation gamma ray detection region, than conventional scintillator units. Furthermore, it is possible to measure measurement samples with a large pool such as vial samples. Therefore, if it is possible to increase the pool amount, even if the amount of radionuclides is small, by using the radiation detection device of the present invention, it is possible to simultaneously increase the absolute radiation amount of annihilation gamma rays and the amount detected by the scintillator. Therefore, measurement accuracy can be considerably improved.

[0029]

According to the radiation detection device of the present invention, the scintillator unit is composed of three or more scintillators each having at least one flat side wall, and the flat side wall of each scintillator forms a cavity for placing a measurement sample. Since the scintillators are connected without any gaps so that the measurement sample is surrounded by approximately 360 degrees by each plane side wall, almost all of the radiation emitted from the measurement sample can be detected without waste. It is possible to accurately measure the radiation dose in a small amount of measurement sample. In addition, when the measurement sample is blood containing radionuclides (RI), the scintillator is assembled to surround the measurement sample almost 360 degrees without any gaps, so almost all the annihilation gamma rays emitted from the blood are absorbed by the scintillator. It becomes possible to detect. Therefore, even when using a small amount of radionuclides, it is possible to measure time changes in the concentration of radionuclides in the measurement sample (blood) with high precision, so measurement of RI concentration distribution in the body using a positron CT device using radionuclides is possible. accuracy can be increased.

[Brief explanation of drawings]

FIG. 1 shows a Tomoe-type scintillator unit used in the radiation detection device of the present invention.

FIG. 2 shows an example of a radiation detection device of the present invention using the scintillator unit shown in FIG. 1.

3 shows a configuration circuit of a preprocessing unit used in the radiation detection apparatus shown in FIG. 2. FIG.

FIG. 4A shows a scintillator unit having a triangular cross-sectional structure used in the present invention.

FIG. 4B shows a scintillator unit having a rectangular cross-sectional structure used in the present invention.

FIG. 4C shows a scintillator unit having a hexagonal cross-sectional structure used in the present invention.

FIG. 4D shows a scintillator unit having a tomoe-shaped cross-sectional structure used in the present invention.

FIG. 4E shows a scintillator unit with a sandwiched cross-sectional structure used in the present invention.

5A shows an example of a coupling position between a photomultiplier tube and a scintillator used in the radiation detection device shown in FIG. 2. FIG.

5B shows another example of the coupling position of the photomultiplier tube and the scintillator used in the radiation detection device shown in FIG. 2. FIG.

FIG. 6 is for explaining the complete detection function of the Tomoe-type scintillator unit used in the present invention.

FIG. 7 shows the structure of a conventional scintillator unit consisting of two scintillators using a straight blood supply tube.

FIG. 8 shows the structure of a conventional scintillator unit consisting of two scintillators using a U-shaped blood supply tube.

9A shows an example of a cross-sectional structure of the scintillator unit of FIG. 7. FIG.

9B shows another example of the cross-sectional structure of the scintillator unit shown in FIG. 7. FIG.

[Explanation of symbols]

1.Scintillator unit 3 Blood supply tube 4 Photomultiplier tube 5 Pre-treatment unit 6 Simultaneous judger 7 Counter

Claims (2)

[Claims] 1. A scintillator unit that detects radiation from a measurement sample containing a radiation source at at least two positions and generates scintillation light at each detection position; A photoelectric conversion element converts the scintillation light generated by the unit into an amplified electrical signal, and the electrical signal of the photoelectric conversion element is processed simultaneously to detect whether radiation from the measurement sample is incident on the two positions at the same time. In a radiation detection device comprising a simultaneous determination means for determining, the scintillator unit includes three or more scintillators each having at least one flat side wall, and the flat side wall of each scintillator provides a cavity for placing the measurement sample. The planar side wall allows the measurement sample to be approximately 360 mm wide.
A radiation detection device characterized in that the scintillators are connected and arranged so as to be surrounded without any gaps. 2. The radiation detection device according to claim 1, wherein the scintillator is arranged in a tomb shape with the cavity at the center.