WO2010035586A1 - Nuclear medicine diagnostic equipment and method for adjusting same - Google Patents
Nuclear medicine diagnostic equipment and method for adjusting same Download PDFInfo
- Publication number
- WO2010035586A1 WO2010035586A1 PCT/JP2009/064251 JP2009064251W WO2010035586A1 WO 2010035586 A1 WO2010035586 A1 WO 2010035586A1 JP 2009064251 W JP2009064251 W JP 2009064251W WO 2010035586 A1 WO2010035586 A1 WO 2010035586A1
- Authority
- WO
- WIPO (PCT)
- Prior art keywords
- radiation
- collimator
- nuclear medicine
- asymmetry
- degree
- Prior art date
Links
Images
Classifications
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/24—Measuring radiation intensity with semiconductor detectors
- G01T1/249—Measuring radiation intensity with semiconductor detectors specially adapted for use in SPECT or PET
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/24—Measuring radiation intensity with semiconductor detectors
Definitions
- a nuclear medicine diagnostic apparatus employing a pixel type as a radiation detector has been studied (for example, see Patent Document 1).
- a pixel type radiation detector a semiconductor detector using a scintillator such as thallium activated cesium iodide (CsI (Tl)) or a semiconductor radiation detection element such as cadmium telluride (CdTe) is used.
- CsI (Tl) thallium activated cesium iodide
- CdTe cadmium telluride
- the collimator's septa thickness thickness of the side wall of the through-hole
- the collimator's septa thickness thickness of the side wall of the through-hole
- the collimator's septa it is necessary to reduce the size of the holes.
- the pixel-type radiation detector 3 used in the pixel-type gamma camera includes thallium-activated sodium iodide (NaI (Tl)), thallium-activated cesium iodide (CsI (Tl)), and cerium-activated.
- Crystal scintillators such as GSO (GSO (Ce)), cerium activated LSO (LSO (Ce)), bismuth germanium oxide (BGO), etc., a plurality of optical devices for detecting the fluorescence, such as photomultiplier tubes, photo
- a radiation detector in which a plurality of detection units 3a are formed by combining diodes, avalanche photodiodes, or the like can be used.
- the diagnostic image is two-dimensional.
- the nuclear medicine diagnosis apparatus 1 is not limited to a gamma camera, and may be a SPECT apparatus configured using a gamma camera.
- the radiation detection unit 3 and the collimator 2 and the like rotate around the subject 12, and a plurality of two-dimensional diagnostic images can be generated at different angles.
- the SPECT apparatus can generate a three-dimensional diagnostic image and its tomographic image by performing image reconstruction by software based on these diagnostic images.
- the radiation source 17 can be an ampule with the radiopharmaceutical 13 placed in an ampule having a pitch smaller than the pitch.
- the arrangement position of the radiation source 17 is not limited to the bed 11 and may be disposed so as to face the collimator 2.
- the distance between the collimator 2 and the radiation source 17 is about several times the thickness of the collimator 2, or It is preferable that the size of an acquired image having a spread described later is a distance that is about 10 pixels or more in both the x-axis direction and the y-axis direction.
- the gap 3b and the gap center plane 3d become the detection unit boundary plane between the detection units 3a, and this detection unit boundary plane is the center of the side wall. It coincides with the surface 2d. Further, the pitch P2 of the detection units 3a is equal to the pitch P1 of the through holes 2a. Moire does not occur at such match points.
- FIG. 4 shows the positional relationship of the collimator 2 with respect to the radiation detector 3.
- the collimator 2 is not aligned, and the positional relationship of the collimator 2 with respect to the radiation detection unit 3 is not matched.
- the detection unit center axis 3c and the through-hole center axis 2c do not match.
- the gap center plane 3d and the side wall center plane 2d do not match. Moire occurs in such a state.
- the radiation 14 is incident with a negative traveling component in the x-axis direction. For this reason, the irradiation region 22 is generated on the negative side in the x-axis direction on the surface of the detection unit 3a. If the irradiation region 22 is shifted in the negative x-axis direction due to the shift of the collimator 2, the radiation 14 protrudes from the surface of the detection unit 3a, and the area of the irradiation region 22 is reduced.
- the areas of the irradiation regions 21 and 22 are larger on the larger x coordinate (pixel coordinate) side than on the smaller side. I understand that.
- the relative position of the collimator 2 with respect to the radiation detection unit 3 must be adjusted so that the asymmetry of the spread is not observed.
- An index that quantitatively expresses the degree of asymmetry (the degree of asymmetry) is required.
- the degree of asymmetry for example, it is conceivable to use an absolute value S of skewness known as a statistical index.
- the absolute value S of the skewness is given by the following equation when N radiations 14 are incident and the position (pixel coordinate (x coordinate)) where the i-th radiation 14 is incident is represented by xi .
- an index that quantitatively represents the size of the spread of the image 18 of the radiation source 17 is necessary.
- the size of the spread for example, a standard deviation or a full width at half maximum known as a statistical index can be considered.
- a range 23 to be reduced is determined based on the size of the spread, and a match point may be determined based on the degree of asymmetry of the spread within the range 23 that is reduced.
- FIG. 8A shows that the match point can be determined only by detecting the minimum value of the spread size.
- the extent of spread is one cycle increase / decrease at one pitch (P1, P2)
- the degree of spread asymmetry is two cycles increase / decrease at one pitch (P1, P2). Therefore, by using the degree of asymmetry of the spread to determine the match point, it is possible to determine the match point with higher positional accuracy than using the size of the spread.
- the radiation 14 emitted from the radiation source 17 passes through the collimator 2 and enters the detection unit 3 a of the radiation detection unit 3.
- the detection unit 3a converts the radiation 14 into a detection signal.
- the positional relationship between the plurality of detection units 3a and the plurality of pixels constituting the image is associated with each other, and the corresponding pixel is determined based on the count number (frequency) of the detection signal in the detection unit 3a.
- an image 18 of the radiation source 17 is created.
- the color information for example, a gray scale such as 256 gradations in which black corresponds to the count number 0 (zero) and white corresponds to the count number 255 can be used.
- the image processing unit 5 calculates the degree of asymmetry of the spread and the size of the spread using the image 18 of the radiation source 17.
- the image processing unit 5 causes the display device 8 to display the degree of asymmetry of the spread and the size of the spread.
- the manufacturer or the diagnostician moves the collimator 2 using the moving mechanism 6 so that the degree of asymmetry is reduced to a minimum value (zero) while observing the degree of asymmetry displayed on the display device 8.
- the moving mechanism 6 is preferably capable of measuring the relative position of the collimator 2, and the image processing unit 5 is preferably capable of receiving the relative position from the moving mechanism 6.
- the image processing unit 5 can store the relative position, the degree of asymmetry of the spread, and the size of the spread before and after the movement. For example, the image processing unit 5 compares the degree of asymmetry of the spread before and after the movement, so that the movement direction in the next movement is changed from the relative position after the movement to the relative position before the movement. Or the opposite direction.
- the image processing unit 5 causes the display device 8 to display the moving direction of the determination result.
- the manufacturer or the diagnostician looks at the moving direction displayed on the display device 8 and moves the collimator 2 in the moving direction using the moving mechanism 6.
- the size of the spread can be evaluated without storing a predetermined size threshold.
- the collimator 2 is arranged at one of the relative positions where the degree of asymmetry is minimum (minimum).
- the image processing unit 5 calculates and stores the size of the spread at this relative position.
- the moving mechanism 6 is used to move the relative position of the collimator 2 by a half pitch.
- An image of the radiation source 17 is generated at this relative position shifted by a half pitch, and the size of the spread is calculated.
- the size of the spread before and after the half-pitch movement is compared, and the relative position with the smaller spread is set as the match point.
- the image processing unit 5 uses the moving mechanism 6 to move the collimator 2 to a relative position where the size of the spread is smaller.
- the image processing unit 5 When the size of the spread is smaller than the size threshold, the image processing unit 5 performs a large movement of about a half pitch because the current relative position is included in the range 23 that is small in FIG. On the other hand, if it is larger, it is included in the range 24 where the current relative position becomes larger, so a large movement of about a half pitch is performed, and the relative position is moved to a range 23 where it becomes smaller.
- the image processing unit 5 determines a relative position at which the degree of asymmetry is minimized (minimum) based on the degree of asymmetry calculated for each different relative position.
- the image processing unit 5 may display the relative position at which the degree of asymmetry is minimized (minimum) on the display device 8.
- the image processing unit 5 causes the moving mechanism 6 to move the collimator 2 to a relative position where the degree of asymmetry is minimized (minimum).
- the image processing unit 5 determines the moving direction of the collimator 2 by a method similar to the manual method so that the degree of asymmetry is minimized (minimum) based on the degree of asymmetry calculated for each different relative position. You may decide.
- the linear radiation sources 25x and 25y it is possible to obtain the same effect as calculating the degree of asymmetry of spread a plurality of times by changing the position of the dotted radiation source 17. That is, when the relative positions of the collimator 2 and the radiation detector 3 and conditions such as statistical fluctuations overlap depending on the position of the point-like radiation source 17, the collimator 2 is not aligned, The degree of asymmetry may be reduced by chance. In such a case, the coincidence can be eliminated by changing the position of the point-like radiation source 17 and calculating the degree of asymmetry of the spread a plurality of times.
- a one-dimensional image in the x-axis direction is acquired in a row of detection units 3a having a pixel coordinate (y coordinate) y0 + 1 that is one larger than the pixel coordinate (y coordinate) y0, and the degree of asymmetry is calculated.
- This repetitive operation is repeated until the one-dimensional image indicated by the arrow 27 is repeatedly acquired, specifically, until the number of repetitions reaches a predetermined number N and the pixel coordinates (y coordinates) reach y0 + N.
- the average value of N + 1 asymmetry degrees obtained in this way is calculated.
- the average value of the degree of asymmetry is defined as the degree of asymmetry with respect to the linear source 25y.
- the extent of the spread with respect to the linear source 25y can be defined, and the degree of asymmetry and the extent of the extent with respect to the linear source 25x can be defined. And since the two linear radiation sources 25x and 25y are arrange
- a parallel hole collimator has been described as an example of the collimator 2.
- the collimator 2 is not limited thereto, and there is a correspondence between the detection unit 3a of the radiation detection unit 3 and the through hole 2a of the collimator 2.
- the present invention can be applied.
- the present invention is effective for a diverging collimator, a converging collimator, a slant collimator, and the like.
- the pitch of the through holes 2a of the collimator 2 is designed to be an integral multiple of the pitch of the detection units 3a of the radiation detection unit 3, the alignment can be performed by the same method, and the present invention is effective. .
- one pixel is not necessarily constituted by one scintillator or a semiconductor element.
- the pixel-type radiation detection unit 3 may be formed by attaching a pixel-like electrode to one semiconductor crystal.
- the pixel-type radiation detection unit 3 that has strip-shaped electrodes in a direction perpendicular to the front and back surfaces and performs a so-called cross-strip read operation may be used in the present invention.
- the gamma camera has been described.
- the effect of the present invention can be expected even in a SPECT apparatus.
- the point-like line source 17 and the line sources 25x and 25y intersecting each other are used.
- the present invention is not limited to this, and a plurality of point-like lines arranged on the line are used.
- the radiation source may be used.
Landscapes
- Physics & Mathematics (AREA)
- Health & Medical Sciences (AREA)
- Life Sciences & Earth Sciences (AREA)
- General Physics & Mathematics (AREA)
- High Energy & Nuclear Physics (AREA)
- Molecular Biology (AREA)
- Spectroscopy & Molecular Physics (AREA)
- Measurement Of Radiation (AREA)
- Nuclear Medicine (AREA)
Abstract
A nuclear medicine diagnostic equipment (1) comprising a radiation detecting section (3) where a plurality of detection units (3a) for detecting radiation (14) from a radiation source (17) are arranged in lattice structures, a collimator (2) for limiting the incident direction of radiation (14) which impinges on the detection units (3a), and an image forming section (4) for forming the diagnostic image of the radiation source (17) based on the count in a detection signal outputted from the detection units (3a) when the radiation (14) is detected, with the positional relations between the plurality of detection units (3a) and a plurality of pixels kept associated with one another,
is further provided with an image processing section (5) for calculating a degree of asymmetry of the expanse of a diagnostic image with respect to the radiation source (17) from the symmetry of the count, and a moving mechanism (6) for moving the position of the collimator (2) relative to the radiation detecting section (3). A nuclear medicine diagnostic equipment which can ensure high spatial resolution by controlling the occurrence of moiré is thereby provided.
Description
本発明は、線源から入射する放射線を検出する放射線検出部と、この放射線検出部に入射する放射線の入射方向を制限するコリメータとを有する核医学診断装置およびその調整方法に関する。
The present invention relates to a nuclear medicine diagnostic apparatus having a radiation detection unit that detects radiation incident from a radiation source, and a collimator that limits the incident direction of radiation incident on the radiation detection unit, and an adjustment method thereof.
核医学診断装置を用いた核医学診断では、特定の臓器や腫瘍に集積する性質があり放射性同位体で標識された薬剤を、被検体に投与する。その薬剤は集積された特定の臓器や腫瘍から放射線(ガンマ線)を放射するので、放射線検出器により放射線を検出することで、被検体内で集積された薬剤の分布に応じた放射線の検出頻度(カウント)分布を診断画像として得ることができる。核医学診断装置では、この診断画像に基づいて、その薬剤が集積される特定の臓器や腫瘍の被検体における位置を特定することができる。
In nuclear medicine diagnosis using a nuclear medicine diagnostic apparatus, a drug that has the property of accumulating in a specific organ or tumor and is labeled with a radioisotope is administered to a subject. Since the drug emits radiation (gamma rays) from a specific organ or tumor that has accumulated, detecting the radiation with a radiation detector allows the detection frequency of the radiation according to the distribution of the drug accumulated in the subject ( Count) distribution can be obtained as a diagnostic image. In the nuclear medicine diagnostic apparatus, based on this diagnostic image, it is possible to specify the position of the specific organ or tumor in which the drug is accumulated in the subject.
核医学診断装置、特に、ガンマカメラや、ガンマカメラを用いて構成される単光子放出型断層撮影(Single Photon Emission Computed Tomography; SPECT)装置では、放射線検出器の放射線の入射側にコリメータを設けることで、放射線の入射方向を制限し、前記診断画像の広がり(ぼけ)を抑制している。
In a nuclear medicine diagnostic device, especially a single-photon emission tomography (SPECT) device constructed using a gamma camera or gamma camera, a collimator is provided on the radiation incident side of the radiation detector. Thus, the incident direction of the radiation is limited, and the spread (blurring) of the diagnostic image is suppressed.
そして、近年、放射線検出器としてピクセル型を採用した核医学診断装置が検討されている(例えば、特許文献1等参照)。ピクセル型の放射線検出器として、タリウム賦活ヨウ化セシウム(CsI(Tl))などのシンチレータや、カドミウムテルル(CdTe)などの半導体放射線検出素子を利用した半導体検出器が用いられている。後者の半導体検出器は、光電子増倍管を介さないで検出信号を読み出すことができるために変換効率が良く、また小型化が容易であることから注目されている。
In recent years, a nuclear medicine diagnostic apparatus employing a pixel type as a radiation detector has been studied (for example, see Patent Document 1). As a pixel type radiation detector, a semiconductor detector using a scintillator such as thallium activated cesium iodide (CsI (Tl)) or a semiconductor radiation detection element such as cadmium telluride (CdTe) is used. The latter semiconductor detector is attracting attention because it can read the detection signal without using a photomultiplier tube, and therefore has high conversion efficiency and is easy to downsize.
しかし、シンチレータでも半導体検出器でもピクセル型を採用した核医学診断装置においては、診断画像にモアレが観察される場合があった。モアレは、診断画像の画質を低下させるので、正確な診断に影響する恐れがあると考えられる。そのため、モアレの発生を抑制することが望まれていた。
However, in the nuclear medicine diagnostic apparatus adopting the pixel type for both the scintillator and the semiconductor detector, moire may be observed in the diagnostic image. Moire degrades the image quality of diagnostic images and is considered to have the potential to affect accurate diagnosis. Therefore, it has been desired to suppress the occurrence of moire.
前記診断画像を構成する複数のピクセルは格子状に配置され、これら複数のピクセルに対応するように、放射線検出器、例えば、半導体検出器の有する複数の半導体放射線検出素子等も格子状に配置されている。そして、複数の半導体放射線検出素子等を配置するピッチと等しいピッチになるように、コリメータでは、複数の貫通孔が格子状に配置されている。モアレは、コリメータの放射線検出器に対する配置位置によって発生したり、発生しなかったりし、より詳細には、放射線検出器の複数の半導体放射線検出素子等に対するコリメータの複数の貫通孔の配置位置(位置合わせ)によって、モアレが発生したり発生しなかったりする。
A plurality of pixels constituting the diagnostic image are arranged in a grid, and a radiation detector, for example, a plurality of semiconductor radiation detection elements included in the semiconductor detector is also arranged in a grid so as to correspond to the plurality of pixels. ing. In the collimator, the plurality of through holes are arranged in a lattice shape so that the pitch is equal to the pitch at which the plurality of semiconductor radiation detection elements and the like are arranged. Moire may or may not be generated depending on the arrangement position of the collimator with respect to the radiation detector, and more specifically, the arrangement positions (positions) of the plurality of through holes of the collimator with respect to the plurality of semiconductor radiation detection elements and the like of the radiation detector. Moiré may or may not occur depending on the combination.
ただ、複数の半導体放射線検出素子等に対するコリメータの複数の貫通孔の配置位置の位置合わせは、放射線検出器の筐体がシールドされているため、中に配置される複数の半導体放射線検出素子等が見えず、光学的に位置合わせを行うことが困難である。
However, the alignment of the arrangement positions of the plurality of through holes of the collimator with respect to the plurality of semiconductor radiation detection elements etc. is because the housing of the radiation detector is shielded, so that the plurality of semiconductor radiation detection elements etc. arranged therein are It is not visible and it is difficult to perform optical alignment.
そこで、複数の半導体放射線検出素子等を配置するピッチよりも充分に細かいピッチで、コリメータの複数の貫通孔を配置したり、ピッチが等しい場合にコリメータのセプタ厚(貫通孔の側壁の厚さ)を大きくしたりして、モアレの発生を抑制することも考えられる。ところが、核医学診断装置においては常に診断画像の一層の高分解能化が求められており、診断画像のピクセルのサイズを小さくするために、放射線検出器の半導体放射線検出素子等のサイズとコリメータの貫通孔のサイズを小さくする必要がある。放射線をコリメートするために、コリメータのセプタ厚をある程度の厚さ(例えば0.2mm)に確保する必要があるので、前記細かいピッチを用いたり、前記セプタ厚を大きくしたりすると、貫通孔のサイズが極端に狭くなり、放射線検出器の感度の低下を招いてしまう。
Therefore, the collimator's septa thickness (thickness of the side wall of the through-hole) when the plurality of through-holes of the collimator are arranged at a pitch sufficiently finer than the pitch at which the plurality of semiconductor radiation detection elements are arranged. It is also conceivable to suppress the generation of moire by increasing the size. However, in nuclear medicine diagnostic devices, there is always a demand for higher resolution of diagnostic images, and in order to reduce the size of pixels in diagnostic images, the size of semiconductor radiation detection elements, etc. of radiation detectors and the penetration of collimators It is necessary to reduce the size of the holes. In order to collimate the radiation, it is necessary to ensure the thickness of the collimator's septa to some extent (for example, 0.2 mm). Therefore, if the fine pitch is used or the thickness of the septa is increased, the size of the through hole Becomes extremely narrow, leading to a decrease in sensitivity of the radiation detector.
そこで、本発明の目的は、モアレの発生を抑制しつつ高空間分解能化を可能にする核医学診断装置とその調整方法を提供することにある。
Therefore, an object of the present invention is to provide a nuclear medicine diagnostic apparatus that enables high spatial resolution while suppressing the occurrence of moire and an adjustment method thereof.
前記目的を達成するために、本発明の特徴は、線源から入射する放射線を検出する複数の検出単位を格子状に配置する放射線検出部と、前記検出単位に入射する前記放射線の入射方向を制限するコリメータと、複数の前記検出単位と複数のピクセルの位置関係を対応させ、前記放射線を検出した際に前記検出単位から出力される検出信号のカウント値に基づいて、前記検出単位に対応する前記ピクセルに色情報を設定して、前記線源の診断画像を作成する画像作成部とを有する核医学診断装置において、
前記線源に対する前記診断画像の広がりの非対称化度を前記カウント値の対称性から計算し、前記放射線検出部に対する前記コリメータの相対位置を移動させる核医学診断装置およびその調整方法にある。 In order to achieve the above object, a feature of the present invention is that a radiation detection unit that arranges a plurality of detection units for detecting radiation incident from a radiation source in a grid pattern, and an incident direction of the radiation incident on the detection unit The collimator to be limited is associated with the positional relationship between the plurality of detection units and the plurality of pixels, and corresponds to the detection unit based on the count value of the detection signal output from the detection unit when the radiation is detected. In a nuclear medicine diagnostic apparatus having an image creation unit that sets color information in the pixels and creates a diagnostic image of the radiation source,
In the nuclear medicine diagnostic apparatus and the adjustment method thereof, the degree of asymmetry of the spread of the diagnostic image with respect to the radiation source is calculated from the symmetry of the count value, and the relative position of the collimator with respect to the radiation detection unit is moved.
前記線源に対する前記診断画像の広がりの非対称化度を前記カウント値の対称性から計算し、前記放射線検出部に対する前記コリメータの相対位置を移動させる核医学診断装置およびその調整方法にある。 In order to achieve the above object, a feature of the present invention is that a radiation detection unit that arranges a plurality of detection units for detecting radiation incident from a radiation source in a grid pattern, and an incident direction of the radiation incident on the detection unit The collimator to be limited is associated with the positional relationship between the plurality of detection units and the plurality of pixels, and corresponds to the detection unit based on the count value of the detection signal output from the detection unit when the radiation is detected. In a nuclear medicine diagnostic apparatus having an image creation unit that sets color information in the pixels and creates a diagnostic image of the radiation source,
In the nuclear medicine diagnostic apparatus and the adjustment method thereof, the degree of asymmetry of the spread of the diagnostic image with respect to the radiation source is calculated from the symmetry of the count value, and the relative position of the collimator with respect to the radiation detection unit is moved.
本発明によれば、モアレの発生を抑制しつつ高空間分解能化を可能にする核医学診断装置とその調整方法を提供することができる。
According to the present invention, it is possible to provide a nuclear medicine diagnostic apparatus that enables high spatial resolution while suppressing the occurrence of moire and an adjustment method thereof.
次に、本発明の実施形態について、適宜図面を参照しながら詳細に説明する。なお、各図において、共通する部分には同一の符号を付し重複した説明を省略する。
Next, embodiments of the present invention will be described in detail with reference to the drawings as appropriate. In each figure, common portions are denoted by the same reference numerals, and redundant description is omitted.
図1に、本発明の実施形態に係る核医学診断装置1のブロック図を示す。図1では、核医学診断装置1を用いた被検体12に対する診断時の様子を示している。核医学診断装置1として、図1では、ピクセル型のガンマカメラを示している。
FIG. 1 shows a block diagram of a nuclear medicine diagnostic apparatus 1 according to an embodiment of the present invention. In FIG. 1, a state at the time of diagnosis for the subject 12 using the nuclear medicine diagnostic apparatus 1 is shown. As the nuclear medicine diagnosis apparatus 1, FIG. 1 shows a pixel-type gamma camera.
放射性同位体で標識された放射性薬剤13は、被検体12に投与されると、特定の臓器や腫瘍に集積する。放射性薬剤13は、その特定の臓器や腫瘍から放射線14を放射する。被検体12がベッド11上に横たえられコリメータ2に対向させられると、放射性薬剤13から放射された放射線14は、コリメータ2の貫通孔2aを通り、放射線検出部3の検出単位3aで検出される。なお、放射線14のうち、コリメータ2および放射線検出部3の表面に対して垂直方向から大きく外れて入射する放射線14は、コリメータ2によって遮られて、放射線検出部3で検出されない。逆に、コリメータ2および放射線検出部3の表面に対して略垂直方向から入射する放射線14は、コリメータ2(貫通孔2a)を通過して、放射線検出部3によって検出される。
The radiopharmaceutical 13 labeled with a radioisotope accumulates in a specific organ or tumor when administered to the subject 12. The radiopharmaceutical 13 emits radiation 14 from the specific organ or tumor. When the subject 12 lies on the bed 11 and faces the collimator 2, the radiation 14 emitted from the radiopharmaceutical 13 passes through the through-hole 2 a of the collimator 2 and is detected by the detection unit 3 a of the radiation detection unit 3. . Of the radiation 14, the radiation 14 that is significantly deviated from the vertical direction with respect to the surfaces of the collimator 2 and the radiation detection unit 3 is blocked by the collimator 2 and is not detected by the radiation detection unit 3. On the contrary, the radiation 14 incident from the substantially vertical direction on the surfaces of the collimator 2 and the radiation detection unit 3 passes through the collimator 2 (through hole 2 a) and is detected by the radiation detection unit 3.
ピクセル型のガンマカメラに用いられるピクセル型の放射線検出部3としては、放射線14により蛍光を発するタリウム賦活ヨウ化ナトリウム(NaI(Tl))、タリウム賦活ヨウ化セシウム(CsI(Tl))、セリウム賦活GSO(GSO(Ce))、セリウム賦活LSO(LSO(Ce))、ビスマスゲルマニウムオキサイド(BGO)などの結晶シンチレータに、前記蛍光を検出するための複数の光デバイス、例えば、光電子増倍管、フォトダイオード、アバランシェフォトダイオードなどを結合させて複数の検出単位3aとした放射線検出器を用いることができる。また、ピクセル型の放射線検出部3として、シリコン(Si)、ゲルマニウム(Ge)、カドミウムテルル(CdTe)、カドミウムヂィンクテルライド(CZT)、臭化タリウム(TlBr)、ヨウ化水銀(HgI2)、ガリウム砒素(GaAs)など半導体を用いた複数の半導体放射線検出素子を複数の検出単位3aとした半導体検出器を用いることができる。また、結晶シンチレータを縦横それぞれ10cm以上の大きさを有する平板として、この平板に複数の光デバイスを格子状に配置することで、ピクセル型の放射線検出部3を構成してもよい。
The pixel-type radiation detector 3 used in the pixel-type gamma camera includes thallium-activated sodium iodide (NaI (Tl)), thallium-activated cesium iodide (CsI (Tl)), and cerium-activated. Crystal scintillators such as GSO (GSO (Ce)), cerium activated LSO (LSO (Ce)), bismuth germanium oxide (BGO), etc., a plurality of optical devices for detecting the fluorescence, such as photomultiplier tubes, photo A radiation detector in which a plurality of detection units 3a are formed by combining diodes, avalanche photodiodes, or the like can be used. In addition, as the pixel type radiation detection unit 3, silicon (Si), germanium (Ge), cadmium tellurium (CdTe), cadmium diinctelluride (CZT), thallium bromide (TlBr), mercury iodide (HgI 2 ), A semiconductor detector in which a plurality of semiconductor radiation detection elements using a semiconductor such as gallium arsenide (GaAs) are used as a plurality of detection units 3a can be used. Alternatively, the pixel-type radiation detection unit 3 may be configured by arranging the crystal scintillator as a flat plate having a size of 10 cm or more in the vertical and horizontal directions and arranging a plurality of optical devices in a lattice pattern on the flat plate.
放射線検出部3で検出された放射線14は検出信号に変換され、後段の画像作成部4のアナログ処理部4aへ送信される。アナログ処理部4aにおいて、検出信号は、前置増幅部、波形整形部、ピークホールド部、AD変換部を通して、デジタルの波高値情報へ変換され、放射線検出部3の検出単位3aの位置情報とともに、画像作成部4のデジタル処理部4bへ送信される。波高値情報は、検出した放射線のエネルギの大きさを表している。位置情報については、検出単位3aさらにはピクセルの一つ一つに対応するように複数のアナログ処理部4aを設けている場合は、アナログ処理部4aから波高値情報が送信されてきた際のその送信経路によって、対応する位置情報を生成することができる。
The radiation 14 detected by the radiation detection unit 3 is converted into a detection signal and transmitted to the analog processing unit 4a of the subsequent image creation unit 4. In the analog processing unit 4a, the detection signal is converted into digital peak value information through the preamplification unit, the waveform shaping unit, the peak hold unit, and the AD conversion unit, and together with the position information of the detection unit 3a of the radiation detection unit 3, It is transmitted to the digital processing unit 4b of the image creation unit 4. The peak value information represents the magnitude of the detected radiation energy. As for the position information, when a plurality of analog processing units 4a are provided so as to correspond to each detection unit 3a and each pixel, the peak value information when the analog processing unit 4a transmits the peak value information. Corresponding position information can be generated according to the transmission path.
デジタル処理部4bは、波高値情報と位置情報に対して必要な補正を行った後、波高値情報が有効なイベントであるか否かの判定を行う。波高値情報のイベントが有効か否かの判定として、放射線14のエネルギを表す波高値情報が、例えば、入力装置9から入力されたエネルギウィンドウの範囲内にあるか否かの判定が挙げられる。波高値情報がエネルギウィンドウの範囲内であり、有効なイベントであった場合は、対応する位置情報のカウント数(頻度)を1増やす。前記のことを放射線検出部3で検出される複数の放射線14について行うことにより、被検体12内の放射性薬剤13の分布に応じた放射線14の検出頻度(カウント)分布を診断画像として得ることができる。
The digital processing unit 4b performs necessary correction on the peak value information and the position information, and then determines whether the peak value information is a valid event. The determination as to whether the event of the peak value information is valid includes determination of whether the peak value information representing the energy of the radiation 14 is within the range of the energy window input from the input device 9, for example. When the peak value information is within the energy window range and is a valid event, the count number (frequency) of the corresponding position information is increased by one. By performing the above for a plurality of radiations 14 detected by the radiation detection unit 3, the detection frequency (count) distribution of the radiation 14 according to the distribution of the radiopharmaceutical 13 in the subject 12 can be obtained as a diagnostic image. it can.
診断画像は、被検体12の診断時には、画像作成部4から画像処理部5へ送信されるが、画像処理部5では処理されること無く、そのままコンソール7の表示装置8に送信される。また、被検体12の診断時に移動機構6は動かない。移動機構6は、コリメータ2を、放射線検出部3に対して相対的に移動させたり、固定させたりすることができるが、被検体12の診断時には、コリメータ2を放射線検出部3に対して変位しないように固定する。
The diagnostic image is transmitted from the image creation unit 4 to the image processing unit 5 when the subject 12 is diagnosed, but is not processed by the image processing unit 5 and is transmitted to the display device 8 of the console 7 as it is. Further, the moving mechanism 6 does not move when the subject 12 is diagnosed. The moving mechanism 6 can move or fix the collimator 2 relative to the radiation detection unit 3, but displace the collimator 2 relative to the radiation detection unit 3 when diagnosing the subject 12. Do not fix.
表示装置8は、診断画像15を表示する。診断画像15は、放射性薬剤13のそのままの形状に対応する広がりのない理想の診断画像16があったとすると、その理想の診断画像16に対して、ぼけて広がる。この広がりは、コリメータ2によって抑制されており図1では誇張して記載しているが僅かに存在している。後記するが、本発明では、この広がりを使用している。そして、診断者は、この診断画像15に基づいて、放射性薬剤13が集積した特定の臓器や腫瘍の被検体12における位置を特定することができる。
The display device 8 displays the diagnostic image 15. If there is an ideal diagnostic image 16 having no spread corresponding to the shape of the radiopharmaceutical 13 as it is, the diagnostic image 15 is blurred and spreads with respect to the ideal diagnostic image 16. This spread is suppressed by the collimator 2 and is exaggerated in FIG. As will be described later, this spread is used in the present invention. Then, based on the diagnostic image 15, the diagnostician can specify the position of the specific organ or tumor in which the radiopharmaceutical 13 is accumulated in the subject 12.
前記のように、核医学診断装置1がガンマカメラであれば、診断画像は2次元的になる。核医学診断装置1は、ガンマカメラに限らず、ガンマカメラを用いて構成されるSPECT装置であってもよい。SPECT装置においては、被検体12を中心として、放射線検出部3およびコリメータ2等が回転し、角度を変えて複数の2次元的な診断画像を生成させることができる。そして、SPECT装置では、これらの診断画像を元にソフトウェアによる画像再構成を行うことで、3次元的な診断画像やその断層画像を生成させることができる。
As described above, if the nuclear medicine diagnostic apparatus 1 is a gamma camera, the diagnostic image is two-dimensional. The nuclear medicine diagnosis apparatus 1 is not limited to a gamma camera, and may be a SPECT apparatus configured using a gamma camera. In the SPECT apparatus, the radiation detection unit 3 and the collimator 2 and the like rotate around the subject 12, and a plurality of two-dimensional diagnostic images can be generated at different angles. The SPECT apparatus can generate a three-dimensional diagnostic image and its tomographic image by performing image reconstruction by software based on these diagnostic images.
図2に、本発明の実施形態に係る核医学診断装置1のブロック図を示す。図2では、被検体12の診断時ではなく、核医学診断装置1の製造時のような、放射線検出部3に対するコリメータ2の位置合わせ時の様子を示している。被検体12の診断時と異なっている点は、被検体12に替えて、ベッド11上に点状の線源17が配置されている点である。点状とはいっても、線源17の大きさは、コリメータ2の複数の貫通孔2aのピッチ、および、放射線検出部3の複数の検出単位3aのピッチと同程度の大きさかそれ以下であればよい。具体的には、放射性薬剤13は通常液体であるので、線源17としては、前記ピッチの大きさ以下のアンプルに放射性薬剤13を入れて封入したものを用いることができる。また、線源17の配置位置は、ベッド11上に限られず、コリメータ2に対向して配置されればよく、コリメータ2と線源17の距離は、コリメータ2の厚みの数倍程度以上、あるいは、後記する取得される広がりを有する画像の大きさがx軸方向、y軸方向ともに10ピクセル程度以上となる距離であることが好ましい。
FIG. 2 shows a block diagram of the nuclear medicine diagnostic apparatus 1 according to the embodiment of the present invention. FIG. 2 shows a state when the collimator 2 is aligned with the radiation detection unit 3, such as when the nuclear medicine diagnostic apparatus 1 is manufactured, not when the subject 12 is diagnosed. The difference from the time of diagnosis of the subject 12 is that a point-like radiation source 17 is arranged on the bed 11 instead of the subject 12. Even if it is point-like, the size of the radiation source 17 should be equal to or smaller than the pitch of the plurality of through holes 2 a of the collimator 2 and the pitch of the plurality of detection units 3 a of the radiation detector 3. That's fine. Specifically, since the radiopharmaceutical 13 is usually a liquid, the radiation source 17 can be an ampule with the radiopharmaceutical 13 placed in an ampule having a pitch smaller than the pitch. Further, the arrangement position of the radiation source 17 is not limited to the bed 11 and may be disposed so as to face the collimator 2. The distance between the collimator 2 and the radiation source 17 is about several times the thickness of the collimator 2, or It is preferable that the size of an acquired image having a spread described later is a distance that is about 10 pixels or more in both the x-axis direction and the y-axis direction.
線源17から放射線14が放射され、コリメータ2の貫通孔2aを通った放射線14が放射線検出部3の検出単位3aで検出される。コリメータ2と、放射線検出部3と、画像作成部4とでは、図1で説明したのと同様な処理が行われ、画像作成部4で作成された線源17の画像が、画像処理部5へ送信され、また、画像処理部5を介して表示装置8へ送信される。画像処理部5では、詳細は後記するが、線源17の画像に基づいて、位置合わせをするために、移動機構6がどのように移動させられるのかが決定される。この決定に基づいて、移動機構6を用いて、自動(オート)または手動(マニュアル)で、放射線検出部3に対するコリメータ2の相対位置が移動(変更)される。
Radiation 14 is emitted from the radiation source 17, and the radiation 14 that has passed through the through hole 2 a of the collimator 2 is detected by the detection unit 3 a of the radiation detection unit 3. The collimator 2, the radiation detection unit 3, and the image creation unit 4 perform the same processing as described in FIG. 1, and the image of the radiation source 17 created by the image creation unit 4 is converted into the image processing unit 5. And to the display device 8 via the image processing unit 5. Although details will be described later, the image processing unit 5 determines how the moving mechanism 6 is moved for alignment based on the image of the radiation source 17. Based on this determination, the relative position of the collimator 2 with respect to the radiation detector 3 is moved (changed) automatically (automatically) or manually (manually) using the moving mechanism 6.
表示装置8は、線源17の画像18を表示する。画像18は、線源17のそのままの形状に対応する広がりのない理想の画像19があったとすると、その理想の画像19に対して、ぼけて広がる。放射線検出部3に対するコリメータ2の相対位置の位置合わせが行われていないと、この広がりは、非対称になる。本実施の形態では、この広がりの非対称性に着目し、画像処理部5において、広がりの非対称性を数値化した非対称化度を計算し、広がりの非対称化度が小さくなるように、移動機構6を用いて、コリメータ2の位置合わせを行っている。なお、非対称化度の計算の詳細は後記する。
The display device 8 displays an image 18 of the radiation source 17. If there is an ideal image 19 having no spread corresponding to the shape of the radiation source 17 as it is, the image 18 blurs and spreads with respect to the ideal image 19. If the relative position of the collimator 2 with respect to the radiation detection unit 3 is not aligned, this spread becomes asymmetric. In the present embodiment, paying attention to the asymmetry of the spread, the image processing unit 5 calculates the asymmetry degree obtained by quantifying the asymmetry of the spread, and the moving mechanism 6 so that the degree of asymmetry of the spread becomes small. Is used to align the collimator 2. Details of the calculation of the degree of asymmetry will be described later.
図3に、コリメータ2の放射線検出部3に対する位置関係を示す。そして、例えば、核医学診断装置1の製造段階で、コリメータ2の位置合わせが行われ、コリメータ2の放射線検出部3に対する位置関係がマッチ(整合)している状態(マッチポイント)を示している。マッチ(整合)している状態では、検出単位3aの検出単位中心軸3cと、貫通孔2aの貫通孔中心軸2cとが一致している。また、マッチ(整合)している状態では、隣り合う検出単位3aの間の間隙3bの間隙中央面3dと、隣り合う貫通孔2aを仕切る側壁2bの側壁中央面2dとが一致している。なお、間隙3bが存在せず、検出単位3a同士が接している場合は、間隙3bおよび間隙中央面3dは、検出単位3a同士の検出単位境界面になり、この検出単位境界面は、側壁中央面2dに一致する。また、検出単位3aのピッチP2は、貫通孔2aのピッチP1に等しくなっている。このようなマッチポイントでは、モアレは発生しない。
FIG. 3 shows the positional relationship of the collimator 2 with respect to the radiation detector 3. For example, in the manufacturing stage of the nuclear medicine diagnostic apparatus 1, the collimator 2 is aligned, and the positional relationship of the collimator 2 with respect to the radiation detection unit 3 is matched (match point). . In the matched state, the detection unit central axis 3c of the detection unit 3a and the through-hole central axis 2c of the through-hole 2a coincide. In the matched state, the gap center surface 3d of the gap 3b between the adjacent detection units 3a and the side wall center surface 2d of the side wall 2b partitioning the adjacent through-holes 2a coincide with each other. When the gap 3b does not exist and the detection units 3a are in contact with each other, the gap 3b and the gap center plane 3d become the detection unit boundary plane between the detection units 3a, and this detection unit boundary plane is the center of the side wall. It coincides with the surface 2d. Further, the pitch P2 of the detection units 3a is equal to the pitch P1 of the through holes 2a. Moire does not occur at such match points.
図4に、コリメータ2の放射線検出部3に対する位置関係を示す。そして、コリメータ2の位置合わせが行われておらず、コリメータ2の放射線検出部3に対する位置関係がマッチ(整合)していない状態を示している。マッチ(整合)していない状態では、検出単位中心軸3cと、貫通孔中心軸2cとが一致していない。また、マッチ(整合)していない状態では、間隙中央面3dと、側壁中央面2dとが一致していない。このような状態でモアレは発生する。
FIG. 4 shows the positional relationship of the collimator 2 with respect to the radiation detector 3. In addition, the collimator 2 is not aligned, and the positional relationship of the collimator 2 with respect to the radiation detection unit 3 is not matched. In a state where there is no match (matching), the detection unit center axis 3c and the through-hole center axis 2c do not match. Further, in a state where they are not matched, the gap center plane 3d and the side wall center plane 2d do not match. Moire occurs in such a state.
核医学診断装置1が製造後、完成し、診断者に引き渡され診断者に使用されている際にも、コリメータ2は診断の用途に合わせて使い分けるため、複数のコリメータ2が取り付けられたり外されたりする。この取り付けの際に、マッチ(整合)していない状態でコリメータ2が取り付けられる可能性がある。
Even when the nuclear medicine diagnostic apparatus 1 is manufactured and completed, delivered to the diagnostician, and used by the diagnostician, the collimator 2 is used depending on the purpose of diagnosis. Or At the time of this attachment, there is a possibility that the collimator 2 is attached without matching.
図5に、コリメータ2の位置合わせ時、例えば、製造時におけるコリメータ2と放射線検出部3と線源17の位置関係を模式的に示す。コリメータ2には、四角孔の貫通孔2aを有するパラレルホールコリメータを用いている。図4に示したマッチ(整合)していない状態は、図5のx軸方向に対しても、y軸方向に対しても発生すると考えられる。そこで、後記では、x軸方向に対してコリメータ2の位置が放射線検出部3の位置とマッチ(整合)していない場合を例に、線源17の画像の広がりが非対称性を持つことを説明する。
FIG. 5 schematically shows a positional relationship among the collimator 2, the radiation detection unit 3, and the radiation source 17 at the time of alignment of the collimator 2, for example, at the time of manufacture. For the collimator 2, a parallel hole collimator having a square hole 2a is used. It is considered that the unmatched state shown in FIG. 4 occurs both in the x-axis direction and in the y-axis direction in FIG. Therefore, in the following description, it is explained that the spread of the image of the radiation source 17 has asymmetry, taking as an example the case where the position of the collimator 2 does not match (match) the position of the radiation detection unit 3 with respect to the x-axis direction. To do.
図6は、図5のA-A方向の矢視断面図であり、放射線14が放射線検出部3に照射される様子を示している。この例では、コリメータ2の放射線検出部3に対する位置関係は、図4と同様にマッチ(整合)していない。線源17とコリメータ2との距離に対して、貫通孔2aの幅は充分に小さいので、個々の貫通孔2aに入射する複数の放射線14は互いにほぼ平行になっている。この放射線14が貫通孔中心軸2cと平行に入射している検出単位3aにおいて、広がりのない線源17の理想の画像19(図2参照)の検出信号が生成される。放射線14が貫通孔中心軸2cに対して斜めに入射している検出単位3aにおいて、理想の画像19に対して広がりのある線源17の実際の画像18(図2参照)の検出信号が生成される。
FIG. 6 is a cross-sectional view taken along the line AA in FIG. 5 and shows a state in which the radiation 14 is irradiated to the radiation detection unit 3. In this example, the positional relationship of the collimator 2 with respect to the radiation detector 3 is not matched (matched) as in FIG. Since the width of the through hole 2a is sufficiently small with respect to the distance between the radiation source 17 and the collimator 2, the plurality of radiations 14 incident on the individual through holes 2a are substantially parallel to each other. In the detection unit 3a in which the radiation 14 is incident in parallel to the through-hole central axis 2c, a detection signal of an ideal image 19 (see FIG. 2) of the radiation source 17 having no spread is generated. In the detection unit 3a in which the radiation 14 is incident obliquely with respect to the through-hole central axis 2c, a detection signal of the actual image 18 (see FIG. 2) of the radiation source 17 having a spread with respect to the ideal image 19 is generated. Is done.
図6では、x軸方向にコリメータ2の位置が放射線検出部3の位置とマッチ(整合)しておらず、コリメータ2は放射線検出部3に対してx軸方向の負方向にずれている場合を考える。このずれにより、コリメータ2の放射線検出部3上の影は、放射線検出部3の検出単位3aの表面において、x軸方向の負方向(図6の右方向)にずれることになる。同様に、放射線検出部3の検出単位3aの表面上の放射線14の照射領域21、22は、放射線検出部3の検出単位3aの表面において、x軸方向の負方向にずれることになる。
In FIG. 6, the position of the collimator 2 in the x-axis direction is not matched with the position of the radiation detection unit 3, and the collimator 2 is shifted in the negative x-axis direction with respect to the radiation detection unit 3. think of. Due to this shift, the shadow on the radiation detection unit 3 of the collimator 2 shifts in the negative direction in the x-axis direction (the right direction in FIG. 6) on the surface of the detection unit 3a of the radiation detection unit 3. Similarly, the irradiation regions 21 and 22 of the radiation 14 on the surface of the detection unit 3a of the radiation detection unit 3 are shifted in the negative direction in the x-axis direction on the surface of the detection unit 3a of the radiation detection unit 3.
x座標(ピクセル座標)の小さい側では、例えば、ピクセル座標が2や3では、放射線14はx軸方向の負方向の進行成分を持って入射する。このため、照射領域22は、検出単位3aの表面上のx軸方向の負方向の側に生じる。この照射領域22が、コリメータ2のずれにより、x軸方向の負方向にずれると、放射線14が検出単位3aの表面上からはみ出し、照射領域22の面積が減少してしまう。
On the small side of the x coordinate (pixel coordinate), for example, when the pixel coordinate is 2 or 3, the radiation 14 is incident with a negative traveling component in the x-axis direction. For this reason, the irradiation region 22 is generated on the negative side in the x-axis direction on the surface of the detection unit 3a. If the irradiation region 22 is shifted in the negative x-axis direction due to the shift of the collimator 2, the radiation 14 protrudes from the surface of the detection unit 3a, and the area of the irradiation region 22 is reduced.
x座標(ピクセル座標)の大きい側では、例えば、ピクセル座標が11や12では、放射線14はx軸方向の正方向の進行成分を持って入射する。このため、照射領域21は、検出単位3aの表面上のx軸方向の正方向の側に生じる。この照射領域22が、コリメータ2のずれにより、x軸方向の負方向にずれても、放射線14が検出単位3aの表面上からはみ出すことは無く、照射領域21の面積は減少しない。
On the larger side of the x coordinate (pixel coordinate), for example, when the pixel coordinate is 11 or 12, the radiation 14 is incident with a traveling component in the positive direction of the x axis. For this reason, the irradiation region 21 is generated on the positive side in the x-axis direction on the surface of the detection unit 3a. Even if the irradiation region 22 is displaced in the negative direction in the x-axis direction due to the displacement of the collimator 2, the radiation 14 does not protrude from the surface of the detection unit 3a, and the area of the irradiation region 21 does not decrease.
前記より、コリメータ2が放射線検出部3に対してx軸方向の負方向にずれている場合、x座標(ピクセル座標)の大きい側では、小さい側より、照射領域21、22の面積が大きくなることがわかる。
As described above, when the collimator 2 is deviated in the negative x-axis direction with respect to the radiation detection unit 3, the areas of the irradiation regions 21 and 22 are larger on the larger x coordinate (pixel coordinate) side than on the smaller side. I understand that.
図7に、線源17のx軸方向の一次元画像(コリメータ2が放射線検出部3に対してx軸方向の負方向にずれている場合)を示す。放射線14のカウント数は、照射領域21、22の面積にほぼ比例するので、図7に示すように、x座標(ピクセル座標)の大きい側では、小さい側より、カウント数(頻度)が大きくなり、線源17の画像の広がりにおいて非対称性が生じることがわかる。
7 shows a one-dimensional image of the radiation source 17 in the x-axis direction (when the collimator 2 is displaced in the negative direction of the x-axis direction with respect to the radiation detection unit 3). Since the count number of the radiation 14 is substantially proportional to the areas of the irradiation regions 21 and 22, as shown in FIG. 7, the count number (frequency) is larger on the side where the x coordinate (pixel coordinate) is larger than on the smaller side. It can be seen that asymmetry occurs in the spread of the image of the radiation source 17.
コリメータ2の位置を放射線検出部3にマッチさせるには、前記の広がりの非対称性が見られないようにコリメータ2の放射線検出部3に対する相対位置を調整しなくてはならず、そのためには広がりの非対称性の大きさ(非対称化度)を定量的に表す指標が必要である。非対称化度としては、例えば、統計指標として知られる歪度の絶対値Sを用いることが考えられる。歪度の絶対値Sは、N個の放射線14が入射し、i番目の放射線14の入射した位置(ピクセル座標(x座標))がxiで表される場合に次式で与えられる。
S=|Σ(xi-m)3/N/σ3|
ここで、Σはiが1からNまでに渡っての(xi-m)3の総和、mはxiの平均値、σはxiの標準偏差である。 In order to match the position of thecollimator 2 with the radiation detection unit 3, the relative position of the collimator 2 with respect to the radiation detection unit 3 must be adjusted so that the asymmetry of the spread is not observed. An index that quantitatively expresses the degree of asymmetry (the degree of asymmetry) is required. As the degree of asymmetry, for example, it is conceivable to use an absolute value S of skewness known as a statistical index. The absolute value S of the skewness is given by the following equation when N radiations 14 are incident and the position (pixel coordinate (x coordinate)) where the i-th radiation 14 is incident is represented by xi .
S = | Σ (x i −m) 3 / N / σ 3 |
Here, sigma is the summation i is from 1 to over to N of (x i -m) 3, m is an average value of x i, sigma is the standard deviation of x i.
S=|Σ(xi-m)3/N/σ3|
ここで、Σはiが1からNまでに渡っての(xi-m)3の総和、mはxiの平均値、σはxiの標準偏差である。 In order to match the position of the
S = | Σ (x i −m) 3 / N / σ 3 |
Here, sigma is the summation i is from 1 to over to N of (x i -m) 3, m is an average value of x i, sigma is the standard deviation of x i.
図8(b)に、コリメータ2の放射線検出部3に対する相対位置と、画像の広がりの非対称化度(歪度の絶対値S)の関係を示す。歪度の絶対値Sは、カウント数(頻度)分布が左右対称の場合に、最小値0(ゼロ)となるので、コリメータ2の相対位置を放射線検出部3にマッチするマッチポイントに設定するには、非対称化度(歪度の絶対値S)が小さくなる方向へコリメータ2の位置を移動させていけばよいことがわかる。
FIG. 8B shows the relationship between the relative position of the collimator 2 with respect to the radiation detector 3 and the degree of asymmetry of the spread of the image (absolute value S of the skewness). Since the absolute value S of the skewness has a minimum value of 0 (zero) when the count number (frequency) distribution is symmetrical, the relative position of the collimator 2 is set as a match point that matches the radiation detector 3. Indicates that the position of the collimator 2 may be moved in the direction in which the degree of asymmetry (the absolute value S of the skewness) decreases.
なお、ずれが小さい場合は問題ないが、ずれが大きくなり、コリメータ2の相対位置が、マッチポイントからちょうど半ピッチずれている場合も、非対称化度(歪度の絶対値S)は最小値0(ゼロ)となる。半ピッチのように、ずれが大きくなった場合、単に、非対称化度(歪度の絶対値S)が最小値0(ゼロ)となるように、コリメータ2の位置調整を行ったのでは、半ピッチずれた状態に調整されてしまう可能性がある。ただ、マッチポイントと、マッチポイントから半ピッチずれた状態とでは、線源17の画像18の広がりの大きさ(放射線14の放射線検出部3上での広がりの大きさに相当)が異なる。半ピッチずれた状態では、マッチポイントより、空間分解能が悪化し、線源17の画像18がぼけている。
Note that there is no problem if the deviation is small, but the degree of asymmetry (absolute value S of skewness) is the minimum value 0 even when the deviation is large and the relative position of the collimator 2 is exactly half a pitch away from the match point. (Zero). When the deviation becomes large as in the half pitch, the position of the collimator 2 is simply adjusted so that the degree of asymmetry (the absolute value S of the skewness) becomes the minimum value 0 (zero). There is a possibility that the pitch is adjusted to be shifted. However, the size of the spread of the image 18 of the radiation source 17 (corresponding to the size of the spread of the radiation 14 on the radiation detection unit 3) differs between the match point and the state shifted by a half pitch from the match point. In the state shifted by a half pitch, the spatial resolution is worse than the match point, and the image 18 of the radiation source 17 is blurred.
マッチポイントと、マッチポイントから半ピッチずれた状態とを判別するためには、線源17の画像18の広がりの大きさを定量的に表す指標が必要である。広がりの大きさとしては、例えば、統計指標として知られる標準偏差や半値全幅などが考えられる。
In order to discriminate between a match point and a state shifted by a half pitch from the match point, an index that quantitatively represents the size of the spread of the image 18 of the radiation source 17 is necessary. As the size of the spread, for example, a standard deviation or a full width at half maximum known as a statistical index can be considered.
図8(a)に、コリメータ2の放射線検出部3に対する相対位置と、線源17の画像18の広がりの大きさの関係を示す。半ピッチずれた状態では、マッチポイントよりも、広がりの大きさが大きくなって、線源17の画像18がぼけていることがわかる。そして、広がりの大きさが大きくなる範囲24の相対位置で、マッチポイントから半ピッチずれた状態が生じており、広がりの大きさが大きくなる範囲24より小さくなる範囲23の相対位置で、マッチポイントが生じていることがわかる。これより、広がりの大きさに基づいて、小さくなる範囲23を決定し、この小さくなる範囲23の中で、広がりの非対称化度に基づいて、マッチポイントを決定すればよい。なお、図8(a)より、広がりの大きさの最小値を検出することだけでも、マッチポイントを決定できることがわかる。しかし、広がりの大きさは、1ピッチ(P1、P2)で1周期の増減であるところ、広がりの非対称化度は、1ピッチ(P1、P2)で2周期の増減である。このため、マッチポイントを決定するのに、広がりの非対称化度を用いることで、広がりの大きさを用いるより、高い位置精度でマッチポイントを決定することができる。
FIG. 8A shows the relationship between the relative position of the collimator 2 with respect to the radiation detector 3 and the size of the spread of the image 18 of the radiation source 17. It can be seen that in the state of being shifted by a half pitch, the size of the spread is larger than the match point, and the image 18 of the radiation source 17 is blurred. Then, a state that is shifted by a half pitch from the match point occurs in the relative position of the range 24 in which the size of the spread becomes large, and the match point in the relative position of the range 23 that is smaller than the range 24 in which the size of the spread becomes large. It turns out that has occurred. Thus, a range 23 to be reduced is determined based on the size of the spread, and a match point may be determined based on the degree of asymmetry of the spread within the range 23 that is reduced. Note that FIG. 8A shows that the match point can be determined only by detecting the minimum value of the spread size. However, the extent of spread is one cycle increase / decrease at one pitch (P1, P2), while the degree of spread asymmetry is two cycles increase / decrease at one pitch (P1, P2). Therefore, by using the degree of asymmetry of the spread to determine the match point, it is possible to determine the match point with higher positional accuracy than using the size of the spread.
次に、コリメータ2の位置合わせにおける、核医学診断装置1の働きにつき、主に、画像処理部5と移動機構6の働きについて説明する。
Next, the operation of the image processing unit 5 and the moving mechanism 6 will be mainly described with respect to the operation of the nuclear medicine diagnosis apparatus 1 in the alignment of the collimator 2.
図2に示すように、線源17から放射された放射線14は、コリメータ2を通って、放射線検出部3の検出単位3aに入射する。検出単位3aは、放射線14を検出信号に変換する。
As shown in FIG. 2, the radiation 14 emitted from the radiation source 17 passes through the collimator 2 and enters the detection unit 3 a of the radiation detection unit 3. The detection unit 3a converts the radiation 14 into a detection signal.
画像作成部4において、複数の前記検出単位3aと、画像を構成する複数のピクセルの位置関係は対応付けられており、検出単位3aにおける検出信号のカウント数(頻度)に基づいて、対応するピクセルに色情報を設定する。こうして、線源17の画像18が作成される。なお、色情報としては、例えば、黒をカウント数0(ゼロ)に対応させ、白をカウント数255に対応させる256階調といったグレースケールを用いることができる。
In the image creating unit 4, the positional relationship between the plurality of detection units 3a and the plurality of pixels constituting the image is associated with each other, and the corresponding pixel is determined based on the count number (frequency) of the detection signal in the detection unit 3a. Set color information to. Thus, an image 18 of the radiation source 17 is created. As the color information, for example, a gray scale such as 256 gradations in which black corresponds to the count number 0 (zero) and white corresponds to the count number 255 can be used.
画像処理部5は、線源17の画像18を用いて、前記広がりの非対称化度と、前記広がりの大きさを計算する。画像処理部5は、表示装置8に、広がりの非対称化度と、広がりの大きさを表示させる。製造者又は診断者は、表示装置8に表示された非対称化度を見ながら、非対称化度が、小さくなり、最小値(ゼロ)になるように、移動機構6を用いて、コリメータ2を移動させる。
The image processing unit 5 calculates the degree of asymmetry of the spread and the size of the spread using the image 18 of the radiation source 17. The image processing unit 5 causes the display device 8 to display the degree of asymmetry of the spread and the size of the spread. The manufacturer or the diagnostician moves the collimator 2 using the moving mechanism 6 so that the degree of asymmetry is reduced to a minimum value (zero) while observing the degree of asymmetry displayed on the display device 8. Let
また、移動機構6は、コリメータ2の相対位置を計測可能になっていることが好ましく、画像処理部5は、移動機構6から相対位置を受信可能になっていることが好ましい。このことによれば、画像処理部5は、移動の前後での、相対位置と、広がりの非対称化度と、広がりの大きさを記憶することができる。画像処理部5は、例えば、移動の前後での広がりの非対称化度の大小を比較することにより、次の移動の際の移動方向が、移動の後の相対位置から移動の前の相対位置への方向か、その逆方向かを判定することができる。画像処理部5は、判定結果の移動方向を表示装置8に表示させる。製造者又は診断者は、表示装置8に表示された移動方向を見て、移動機構6を用いて、その移動方向にコリメータ2を移動させる。
The moving mechanism 6 is preferably capable of measuring the relative position of the collimator 2, and the image processing unit 5 is preferably capable of receiving the relative position from the moving mechanism 6. According to this, the image processing unit 5 can store the relative position, the degree of asymmetry of the spread, and the size of the spread before and after the movement. For example, the image processing unit 5 compares the degree of asymmetry of the spread before and after the movement, so that the movement direction in the next movement is changed from the relative position after the movement to the relative position before the movement. Or the opposite direction. The image processing unit 5 causes the display device 8 to display the moving direction of the determination result. The manufacturer or the diagnostician looks at the moving direction displayed on the display device 8 and moves the collimator 2 in the moving direction using the moving mechanism 6.
また、画像処理部5が、所定の大きさ閾値を記憶しておいてもよい。画像処理部5は、計算された広がりの大きさが、大きさ閾値よりも小さいか否かを判定する。画像処理部5は、この判定結果を表示装置8に表示する。この表示に基づいて、製造者又は診断者は、広がりの大きさが大きさ閾値よりも小さい場合は、現在の相対位置が図8(a)の小さくなる範囲23に含まれていると考えられるので、半ピッチ程度の大きな移動を控えることができる。一方、大きい場合は、現在の相対位置が大きくなる範囲24に含まれていると考えられるので、半ピッチ程度の大きな移動を行うことになる。
Further, the image processing unit 5 may store a predetermined size threshold value. The image processing unit 5 determines whether or not the calculated spread size is smaller than the size threshold value. The image processing unit 5 displays the determination result on the display device 8. On the basis of this display, the manufacturer or the diagnostician, when the size of the spread is smaller than the size threshold, is considered that the current relative position is included in the smaller range 23 in FIG. Therefore, a large movement of about half a pitch can be avoided. On the other hand, when it is large, it is considered that the current relative position is included in the range 24 in which the relative position becomes large, and therefore, a large movement of about a half pitch is performed.
なお、所定の大きさ閾値を記憶しておかなくても、広がりの大きさを評価することができる。これにはまず、非対称化度が最小(極小)となる相対位置の1つにコリメータ2を配置する。そして、画像処理部5は、この相対位置で広がりの大きさを計算し記憶する。次に、移動機構6を用いて、コリメータ2の相対位置を半ピッチだけ移動させる。半ピッチ移動したこの相対位置で、線源17の画像を生成し、広がりの大きさを計算する。半ピッチの移動の前後での広がりの大きさを比較し、広がりの大きさが小さい方の相対位置を、マッチポイントに設定する。画像処理部5は、移動機構6で、コリメータ2を広がりの大きさが小さい方の相対位置に移動させる。
Note that the size of the spread can be evaluated without storing a predetermined size threshold. For this, first, the collimator 2 is arranged at one of the relative positions where the degree of asymmetry is minimum (minimum). The image processing unit 5 calculates and stores the size of the spread at this relative position. Next, the moving mechanism 6 is used to move the relative position of the collimator 2 by a half pitch. An image of the radiation source 17 is generated at this relative position shifted by a half pitch, and the size of the spread is calculated. The size of the spread before and after the half-pitch movement is compared, and the relative position with the smaller spread is set as the match point. The image processing unit 5 uses the moving mechanism 6 to move the collimator 2 to a relative position where the size of the spread is smaller.
前記のように、コリメータ2の相対位置を調整(移動)するにあたっては、移動機構6を用いて、製造者又は診断者が手動で調整してもよいが、後記のように、移動機構6を自動で動かしコリメータ2の相対位置を少しずつずらしながら、広がりの非対称化度と広がりの大きさを計算し、広がりの非対称化度と広がりの大きさが最小(極小)となる相対位置を自動的にマッチポイントに設定してもよい。また、広がりの大きさの最小(極小)が決定し難い場合は、前記のように、画像処理部5が、所定の大きさ閾値を記憶しておき、画像処理部5は、計算された広がりの大きさが、大きさ閾値よりも小さいか否かを判定すればよい。画像処理部5は、広がりの大きさが大きさ閾値よりも小さい場合は、現在の相対位置が図8(a)の小さくなる範囲23に含まれているので、半ピッチ程度の大きな移動を行わず、一方、大きい場合は、現在の相対位置が大きくなる範囲24に含まれているので、半ピッチ程度の大きな移動を行い、相対位置を小さくなる範囲23へ移動させる。
As described above, when the relative position of the collimator 2 is adjusted (moved), the manufacturer or the diagnostician may manually adjust the moving mechanism 6 using the moving mechanism 6, but the moving mechanism 6 may be adjusted as described later. While automatically moving the relative position of the collimator 2 little by little, the degree of asymmetry of spread and the size of spread are calculated, and the relative position at which the degree of spread asymmetry and spread is minimized (minimum) automatically. You may set it as a match point. When it is difficult to determine the minimum (minimum) size of the spread, the image processing unit 5 stores a predetermined size threshold as described above, and the image processing unit 5 calculates the calculated spread. What is necessary is just to determine whether the magnitude | size of is smaller than a magnitude | size threshold value. When the size of the spread is smaller than the size threshold, the image processing unit 5 performs a large movement of about a half pitch because the current relative position is included in the range 23 that is small in FIG. On the other hand, if it is larger, it is included in the range 24 where the current relative position becomes larger, so a large movement of about a half pitch is performed, and the relative position is moved to a range 23 where it becomes smaller.
自動の場合、画像処理部5は、異なる相対位置毎に計算された非対称化度に基づいて、非対称化度が最小(極小)になる相対位置を決定する。画像処理部5は、表示装置8に、非対称化度が最小(極小)になる相対位置を表示してもよい。画像処理部5は、移動機構6に、コリメータ2を非対称化度が最小(極小)になる相対位置に移動させる。
In the case of automatic, the image processing unit 5 determines a relative position at which the degree of asymmetry is minimized (minimum) based on the degree of asymmetry calculated for each different relative position. The image processing unit 5 may display the relative position at which the degree of asymmetry is minimized (minimum) on the display device 8. The image processing unit 5 causes the moving mechanism 6 to move the collimator 2 to a relative position where the degree of asymmetry is minimized (minimum).
なお、画像処理部5は、異なる相対位置毎に計算された非対称化度に基づいて、非対称化度が最小(極小)になるように、手動における方法と同様の方法でコリメータ2の移動方向を決定してもよい。
The image processing unit 5 determines the moving direction of the collimator 2 by a method similar to the manual method so that the degree of asymmetry is minimized (minimum) based on the degree of asymmetry calculated for each different relative position. You may decide.
図9に、実施の形態の変形例を示す。実施の形態の変形例では、図5の実施の形態と比較して、点状の線源17が、線状の線源25x、25yになっている点が異なっている。線状の線源25xは、直線状の線源であり、xy平面(コリメータ2の入射面)と平行な平面上に沿って配置され、x軸方向(コリメータ2の貫通孔2aの一方の配列方向)から僅かに傾いている。線状の線源25yは、直線状の線源であり、xy平面と平行な平面上に沿って配置され、y軸方向(コリメータ2の貫通孔2aの他方の配列方向)から僅かに傾いている。傾きの角度は、5度以上10度以下程度の大きさである。線状の線源25xと、線状の線源25yとは、互いに交差している。また、線源25x、25yの線状の太さは、前記ピッチP1、P2以下であることが好適である。線状の線源25x、25yは、太さがピッチP1、P2以下の筒状のアンプルに、液体の放射性薬剤13を封入したものを用いることができる。
FIG. 9 shows a modification of the embodiment. The modification of the embodiment is different from the embodiment of FIG. 5 in that the point-like radiation source 17 is a linear radiation source 25x, 25y. The linear radiation source 25x is a linear radiation source, which is arranged along a plane parallel to the xy plane (incident surface of the collimator 2), and is arranged in the x-axis direction (one of the through holes 2a of the collimator 2). Direction). The linear radiation source 25y is a linear radiation source, is disposed along a plane parallel to the xy plane, and is slightly inclined from the y-axis direction (the other arrangement direction of the through holes 2a of the collimator 2). Yes. The inclination angle is about 5 degrees to 10 degrees. The linear source 25x and the linear source 25y intersect each other. Further, it is preferable that the line thickness of the radiation sources 25x and 25y is equal to or less than the pitches P1 and P2. As the linear radiation sources 25x and 25y, a cylindrical ampule having a thickness of pitch P1 or less than P2 and sealed with a liquid radioactive medicine 13 can be used.
線状の線源25x、25yを用いることにより、点状の線源17の位置を変えて複数回、広がりの非対称化度を計算したのと同じ効果を得ることができる。すなわち、点状の線源17の位置によって、コリメータ2と放射線検出部3との相対的な位置や、統計揺らぎなどの条件が重なると、コリメータ2の位置合わせができていないにも拘らず、偶然に非非対称化度が小さくなる可能性がある。このような場合、点状の線源17の位置を変えて複数回、広がりの非対称化度を計算することで、その偶然を排除することができる。そして、x、y軸方向から僅かに傾いた線状の線源25x、25yであれば、一次元画像を生成する放射線検出部3の検出単位3aの列を変えることにより、線源25x、25yのコリメータ2と放射線検出部3とに対する相対的な位置を変更することができる。そして、カウント数を2次元的に1回計測すればよいので、点状の線源17の位置を変えて複数回計測する場合に比べて、トータルの処理時間を短縮することができる。また、封入する放射性薬剤13の量が多くなるので、短時間のうちに多くの放射線をカウントすることができ、さらに、トータルの処理時間を短縮することができるだけでなく、非対称化度の計算精度、位置合わせ精度を高めることができる。
By using the linear radiation sources 25x and 25y, it is possible to obtain the same effect as calculating the degree of asymmetry of spread a plurality of times by changing the position of the dotted radiation source 17. That is, when the relative positions of the collimator 2 and the radiation detector 3 and conditions such as statistical fluctuations overlap depending on the position of the point-like radiation source 17, the collimator 2 is not aligned, The degree of asymmetry may be reduced by chance. In such a case, the coincidence can be eliminated by changing the position of the point-like radiation source 17 and calculating the degree of asymmetry of the spread a plurality of times. If the line sources 25x and 25y are slightly inclined from the x and y axis directions, the source units 25x and 25y are obtained by changing the columns of the detection units 3a of the radiation detection unit 3 that generates a one-dimensional image. The relative positions of the collimator 2 and the radiation detector 3 can be changed. Since the count number only needs to be measured once in a two-dimensional manner, the total processing time can be shortened as compared to the case where the position of the point-like radiation source 17 is changed and measured a plurality of times. Further, since the amount of the radiopharmaceutical 13 to be encapsulated is increased, it is possible to count a large amount of radiation within a short time, and not only can the total processing time be shortened, but also the calculation accuracy of the degree of asymmetry. The positioning accuracy can be increased.
図10に、2本の線状の線源25x、25yの平面図を示す。そして、線状の線源25x、25yを投影した放射線検出部3の入射表面において、x軸方向の一次元画像を取得する際の位置と一次元画像の幅とを示す矢印26xを記載している。また、y軸方向の一次元画像を取得する際の位置と一次元画像の幅とを示す矢印26yを記載している。そして、一次元画像を繰り返し取得する範囲を矢印27として記載している。
FIG. 10 shows a plan view of two linear sources 25x and 25y. An arrow 26x indicating the position and width of the one-dimensional image at the time of acquiring the one-dimensional image in the x-axis direction on the incident surface of the radiation detection unit 3 projected with the linear radiation sources 25x and 25y is described. Yes. In addition, an arrow 26y indicating the position when acquiring a one-dimensional image in the y-axis direction and the width of the one-dimensional image is described. A range in which a one-dimensional image is repeatedly acquired is indicated as an arrow 27.
広がりの非対称化度の計算を、線状の線源25yに関して行う場合について述べる。まず、あるピクセル座標(y座標)y0を持つ検出単位3aの列(図10中では紙面横方向に並ぶ列)において、x軸方向の一次元画像を取得し、非対称化度を計算する。次に、ピクセル座標(y座標)y0より一つ大きいピクセル座標(y座標)y0+1を持つ検出単位3aの列において、x軸方向の一次元画像を取得し、非対称化度を計算する。この繰り返し作業を、矢印27の示す一次元画像を繰り返し取得する範囲、具体的には、繰り返し回数が所定の回数Nに達し、ピクセル座標(y座標)がy0+Nに達するまで繰り返す。こうして得られたN+1個の非対称化度の平均値を計算する。この非対称化度の平均値を、線状の線源25yに対する非対称化度と定義する。同様に、線状の線源25yに対する広がりの大きさを定義することができ、線状の線源25xに対する広がりの非対称化度と大きさを定義することができる。そして、2本の線状の線源25x、25yが配置されているので、x軸、y軸両方向に対して一度に位置合わせを行うことができる。そして、図10に示すように、4つの矢印27で示す4つの一次元画像を繰り返し取得する範囲を設けることにより、コリメータ2のx軸方向とy軸方向のずれだけでなく、xy平面上での回転方向のずれも、4つの範囲に対応する4つの非対称化度として検出でき、これらの非対称化度をそれぞれ最小になるように、移動機構6を調整することでずれを修正することができる。
Described below is the case where the spread asymmetry degree is calculated for the linear source 25y. First, in a row of detection units 3a having a certain pixel coordinate (y coordinate) y0 (a row aligned in the horizontal direction in the drawing in FIG. 10), a one-dimensional image in the x-axis direction is acquired and the degree of asymmetry is calculated. Next, a one-dimensional image in the x-axis direction is acquired in a row of detection units 3a having a pixel coordinate (y coordinate) y0 + 1 that is one larger than the pixel coordinate (y coordinate) y0, and the degree of asymmetry is calculated. This repetitive operation is repeated until the one-dimensional image indicated by the arrow 27 is repeatedly acquired, specifically, until the number of repetitions reaches a predetermined number N and the pixel coordinates (y coordinates) reach y0 + N. The average value of N + 1 asymmetry degrees obtained in this way is calculated. The average value of the degree of asymmetry is defined as the degree of asymmetry with respect to the linear source 25y. Similarly, the extent of the spread with respect to the linear source 25y can be defined, and the degree of asymmetry and the extent of the extent with respect to the linear source 25x can be defined. And since the two linear radiation sources 25x and 25y are arrange | positioned, it can align at once with respect to both x-axis and y-axis directions. Then, as shown in FIG. 10, by providing a range in which four one-dimensional images indicated by four arrows 27 are repeatedly acquired, not only the deviation of the collimator 2 in the x-axis direction and the y-axis direction but also on the xy plane. In the rotational direction can be detected as four asymmetries corresponding to the four ranges, and the shift can be corrected by adjusting the moving mechanism 6 so that these asymmetries are minimized. .
なお、本実施の形態では、コリメータ2として、パラレルホールコリメータを例に説明したが、これに限らず、放射線検出部3の検出単位3aと、コリメータ2の貫通孔2aとに対応関係があれば、本発明を適用することが可能である。例えば、ダイバージングコリメータ、コンバージングコリメータ、スラントコリメータなどに対しても本発明は有効である。また、例えば、コリメータ2の貫通孔2aのピッチが、放射線検出部3の検出単位3aのピッチの整数倍とした設計の場合でも、同様の方法で位置合わせを実施でき、本発明は有効である。
In the present embodiment, a parallel hole collimator has been described as an example of the collimator 2. However, the collimator 2 is not limited thereto, and there is a correspondence between the detection unit 3a of the radiation detection unit 3 and the through hole 2a of the collimator 2. The present invention can be applied. For example, the present invention is effective for a diverging collimator, a converging collimator, a slant collimator, and the like. Further, for example, even when the pitch of the through holes 2a of the collimator 2 is designed to be an integral multiple of the pitch of the detection units 3a of the radiation detection unit 3, the alignment can be performed by the same method, and the present invention is effective. .
また、ピクセル型の放射線検出部3においては、必ずしも一つのピクセルが一つのシンチレータあるいは半導体素子によって構成されていなくても良い。例えば、半導体素子の場合、一つの半導体結晶にピクセル状の電極を付けることでピクセル型の放射線検出部3としてもよい。また、表面と裏面で互いに直行する方向に短冊状の電極をつけ、いわゆるクロスストリップ型の読み出し動作を行うピクセル型の放射線検出部3を、本発明で用いてもよい。
Further, in the pixel-type radiation detection unit 3, one pixel is not necessarily constituted by one scintillator or a semiconductor element. For example, in the case of a semiconductor element, the pixel-type radiation detection unit 3 may be formed by attaching a pixel-like electrode to one semiconductor crystal. In addition, the pixel-type radiation detection unit 3 that has strip-shaped electrodes in a direction perpendicular to the front and back surfaces and performs a so-called cross-strip read operation may be used in the present invention.
また、本実施の形態では、ガンマカメラについて述べたが、SPECT装置でも本発明の効果が期待できる。
In the present embodiment, the gamma camera has been described. However, the effect of the present invention can be expected even in a SPECT apparatus.
また、本実施の形態では、1点の点状の線源17と、互いに交差する線状の線源25x、25yとを用いたが、これらに限られず、線上に配置された複数の点状の線源を用いてもよい。
Further, in the present embodiment, the point-like line source 17 and the line sources 25x and 25y intersecting each other are used. However, the present invention is not limited to this, and a plurality of point-like lines arranged on the line are used. The radiation source may be used.
1 核医学診断装置
2 コリメータ
2a 貫通孔
2b 側壁
2c 貫通孔中心軸
2d 側壁中央面
3 放射線検出部
3a 検出単位(半導体放射線検出素子)
3b 間隙(検出単位境界面)
3c 検出単位中心軸
3d 間隙中央面(検出単位境界面)
4 画像作成部
4a アナログ処理部
4b デジタル処理部
5 画像処理部
6 移動機構
7 コンソール
8 表示装置
9 入力装置
11 ベッド
12 被検体
13 放射性薬剤
14 放射線(ガンマ線)
15 広がりのある実際の診断画像
16 広がりのない理想の診断画像
17 点状の線源
18 広がりのある点状の線源の実際の画像
19 広がりのない点状の線源の理想の画像
21、22 照射領域
23 小さくなる範囲
24 大きくなる範囲
25x x軸方向から僅かに傾いた線状の線源
25y y軸方向から僅かに傾いた線状の線源
26x x軸方向の一次元画像の画像幅
26y y軸方向の一次元画像の画像幅
27 一次元画像の取得範囲 DESCRIPTION OFSYMBOLS 1 Nuclear medicine diagnostic apparatus 2 Collimator 2a Through-hole 2b Side wall 2c Through-hole center axis | shaft 2d Side wall center surface 3 Radiation detection part 3a Detection unit (semiconductor radiation detection element)
3b Gap (detection unit boundary surface)
3c detectionunit center axis 3d gap center plane (detection unit boundary surface)
4Image creation unit 4a Analog processing unit 4b Digital processing unit 5 Image processing unit 6 Moving mechanism 7 Console 8 Display device 9 Input device 11 Bed 12 Subject 13 Radioactive agent 14 Radiation (gamma ray)
DESCRIPTION OFSYMBOLS 15 Actual diagnostic image with spread 16 Ideal diagnostic image without spread 17 Point source 18 Actual image of point source with spread 19 Ideal image 21 of point source without spread 21, 22 Irradiation area 23 Decrease range 24 Increase range 25x Line source slightly tilted from the x-axis direction 25y Line source slightly tilted from the y-axis direction 26x Image width of a one-dimensional image in the x-axis direction 26y Image width of one-dimensional image in y-axis direction 27 Acquisition range of one-dimensional image
2 コリメータ
2a 貫通孔
2b 側壁
2c 貫通孔中心軸
2d 側壁中央面
3 放射線検出部
3a 検出単位(半導体放射線検出素子)
3b 間隙(検出単位境界面)
3c 検出単位中心軸
3d 間隙中央面(検出単位境界面)
4 画像作成部
4a アナログ処理部
4b デジタル処理部
5 画像処理部
6 移動機構
7 コンソール
8 表示装置
9 入力装置
11 ベッド
12 被検体
13 放射性薬剤
14 放射線(ガンマ線)
15 広がりのある実際の診断画像
16 広がりのない理想の診断画像
17 点状の線源
18 広がりのある点状の線源の実際の画像
19 広がりのない点状の線源の理想の画像
21、22 照射領域
23 小さくなる範囲
24 大きくなる範囲
25x x軸方向から僅かに傾いた線状の線源
25y y軸方向から僅かに傾いた線状の線源
26x x軸方向の一次元画像の画像幅
26y y軸方向の一次元画像の画像幅
27 一次元画像の取得範囲 DESCRIPTION OF
3b Gap (detection unit boundary surface)
3c detection
4
DESCRIPTION OF
Claims (16)
- 線源から入射する放射線を検出する複数の検出単位を格子状に配置する放射線検出部と、
前記検出単位に入射する前記放射線の入射方向を制限するコリメータと、
複数の前記検出単位と複数のピクセルの位置関係を対応させ、前記放射線を検出した際に前記検出単位から出力される検出信号のカウント値に基づいて、 前記線源の診断画像を作成する画像作成部とを有する核医学診断装置において、
前記線源に対する前記診断画像の広がりの非対称化度を前記カウント値の対称性から計算する画像処理部と、
前記放射線検出部に対する前記コリメータの相対位置を移動させる移動機構とを有することを特徴とする核医学診断装置。 A radiation detector that arranges a plurality of detection units for detecting radiation incident from the radiation source in a grid pattern;
A collimator for limiting the incident direction of the radiation incident on the detection unit;
Image creation for creating a diagnostic image of the radiation source based on a count value of a detection signal output from the detection unit when the radiation is detected by associating a plurality of detection units with a plurality of pixels. In a nuclear medicine diagnostic apparatus having a
An image processor that calculates the degree of asymmetry of the spread of the diagnostic image with respect to the radiation source from the symmetry of the count value;
A nuclear medicine diagnosis apparatus comprising: a moving mechanism that moves a relative position of the collimator with respect to the radiation detection unit. - 計算された前記非対称化度を表示する表示装置を有することを特徴とする請求の範囲第1項に記載の核医学診断装置。 The nuclear medicine diagnostic apparatus according to claim 1, further comprising a display device that displays the calculated degree of asymmetry.
- 前記画像処理部は、異なる前記相対位置毎に計算された前記非対称化度に基づいて、前記非対称化度が極小になる前記相対位置を決定することを特徴とする請求の範囲第1項又は第2項に記載の核医学診断装置。 The image processing unit determines the relative position at which the degree of asymmetry is minimized based on the degree of asymmetry calculated for each of the different relative positions. The nuclear medicine diagnostic apparatus according to Item 2.
- 前記非対称化度が極小になる前記相対位置を表示する表示装置を有することを特徴とする請求の範囲第3項に記載の核医学診断装置。 The nuclear medicine diagnosis apparatus according to claim 3, further comprising a display device that displays the relative position at which the degree of asymmetry is minimized.
- 前記移動機構は、前記非対称化度が極小になる前記相対位置に、前記コリメータを移動させることを特徴とする請求の範囲第3項に記載の核医学診断装置。 The nuclear medicine diagnosis apparatus according to claim 3, wherein the moving mechanism moves the collimator to the relative position where the degree of asymmetry is minimized.
- 前記画像処理部は、異なる前記相対位置毎に計算された前記非対称化度に基づいて、前記非対称化度が極小になるように、前記コリメータを前記放射線検出部に対して移動させる移動方向を決定することを特徴とする請求の範囲第1項又は第2項に記載の核医学診断装置。 The image processing unit determines a moving direction in which the collimator is moved with respect to the radiation detection unit based on the degree of asymmetry calculated for each of the different relative positions so that the degree of asymmetry is minimized. The nuclear medicine diagnosis apparatus according to claim 1 or 2, wherein the apparatus is a nuclear medicine diagnosis apparatus.
- 前記移動方向を表示する表示装置を有することを特徴とする請求の範囲第6項に記載の核医学診断装置。 The nuclear medicine diagnosis apparatus according to claim 6, further comprising a display device that displays the moving direction.
- 前記移動機構は、前記移動方向に基づいて、前記コリメータを移動させることを特徴とする請求の範囲第6項に記載の核医学診断装置。 The nuclear medicine diagnosis apparatus according to claim 6, wherein the moving mechanism moves the collimator based on the moving direction.
- 前記画像処理部は、前記広がりの大きさを計算することを特徴とする請求の範囲第1項に記載の核医学診断装置。 The nuclear medicine diagnosis apparatus according to claim 1, wherein the image processing unit calculates the size of the spread.
- 計算された前記広がりの前記大きさを表示する表示装置を有することを特徴とする請求の範囲第9項に記載の核医学診断装置。 The nuclear medicine diagnosis apparatus according to claim 9, further comprising a display device that displays the calculated size of the spread.
- 前記画像処理部は、前記コリメータの1ピッチの範囲を前記広がりの前記大きさが小さくなる範囲と前記小さくなる範囲より大きくなる範囲の2つに分け、
前記小さくなる範囲において、前記非対称化度が極小になる前記相対位置を決定することを特徴とする請求の範囲第9項に記載の核医学診断装置。 The image processing unit divides the range of one pitch of the collimator into two ranges, a range in which the size of the spread is reduced and a range that is larger than the reduced range,
10. The nuclear medicine diagnosis apparatus according to claim 9, wherein the relative position at which the degree of asymmetry is minimized is determined in the smaller range. - 複数の前記検出単位それぞれは、検出素子を有することを特徴とする請求の範囲第1項に記載の核医学診断装置。 2. The nuclear medicine diagnosis apparatus according to claim 1, wherein each of the plurality of detection units includes a detection element.
- 複数の前記検出単位それぞれは、半導体放射線検出素子を有することを特徴とする請求の範囲第1項に記載の核医学診断装置。 2. The nuclear medicine diagnostic apparatus according to claim 1, wherein each of the plurality of detection units includes a semiconductor radiation detection element.
- 前記検出単位と、前記コリメータに設けられる貫通孔とは、平面視で四角形であることを特徴とする請求の範囲第1項に記載の核医学診断装置。 The nuclear medicine diagnosis apparatus according to claim 1, wherein the detection unit and the through hole provided in the collimator are square in a plan view.
- 線源から入射する放射線を検出する複数の検出単位を格子状に配置する放射線検出部と、
前記検出単位に入射する前記放射線の入射方向を制限するコリメータと、
複数の前記検出単位と複数のピクセルの位置関係を対応させ、前記放射線を検出した際に前記検出単位から出力される検出信号のカウント値に基づいて、前記線源の診断画像を作成する画像作成部とを有する核医学診断装置の調整方法において、
前記線源に対する前記診断画像の広がりの非対称化度を前記カウント値の対称性から計算し、
前記放射線検出部に対する前記コリメータの相対位置を移動させ、
異なる前記相対位置毎に計算された前記非対称化度に基づいて、前記非対称化度が極小になる前記相対位置を決定することを特徴とする核医学診断装置の調整方法。 A radiation detector that arranges a plurality of detection units for detecting radiation incident from the radiation source in a grid pattern;
A collimator for limiting the incident direction of the radiation incident on the detection unit;
Image creation for creating a diagnostic image of the radiation source based on a count value of a detection signal output from the detection unit when the radiation is detected by associating a plurality of detection units with a plurality of pixels. In a method for adjusting a nuclear medicine diagnostic apparatus having a part,
Calculating the degree of asymmetry of the spread of the diagnostic image relative to the source from the symmetry of the count value;
Moving the collimator relative to the radiation detector;
An adjustment method for a nuclear medicine diagnosis apparatus, wherein the relative position at which the degree of asymmetry is minimized is determined based on the degree of asymmetry calculated for each different relative position. - 前記線源に対する前記診断画像の広がりの非対称化度を計算する際に用いる前記線源として、点状又は線状の線源を用いることを特徴とする請求の範囲第15項に記載の核医学診断装置の調整方法。 16. The nuclear medicine according to claim 15, wherein a point-like or linear source is used as the source used in calculating the degree of asymmetry of the spread of the diagnostic image with respect to the source. Diagnostic device adjustment method.
Applications Claiming Priority (2)
Application Number | Priority Date | Filing Date | Title |
---|---|---|---|
JP2008-244115 | 2008-09-24 | ||
JP2008244115A JP5249694B2 (en) | 2008-09-24 | 2008-09-24 | Nuclear medicine diagnostic apparatus and adjustment method thereof |
Publications (1)
Publication Number | Publication Date |
---|---|
WO2010035586A1 true WO2010035586A1 (en) | 2010-04-01 |
Family
ID=42059593
Family Applications (1)
Application Number | Title | Priority Date | Filing Date |
---|---|---|---|
PCT/JP2009/064251 WO2010035586A1 (en) | 2008-09-24 | 2009-08-12 | Nuclear medicine diagnostic equipment and method for adjusting same |
Country Status (2)
Country | Link |
---|---|
JP (1) | JP5249694B2 (en) |
WO (1) | WO2010035586A1 (en) |
Cited By (1)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
CN113557448A (en) * | 2019-03-29 | 2021-10-26 | 深圳帧观德芯科技有限公司 | Imaging method |
Families Citing this family (3)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
JP4902753B2 (en) * | 2010-01-29 | 2012-03-21 | 株式会社日立製作所 | Radiation imaging apparatus and collimator position estimation method |
JP5852540B2 (en) * | 2012-10-04 | 2016-02-03 | 株式会社日立製作所 | Radiation imaging device |
KR102390955B1 (en) * | 2020-03-02 | 2022-04-25 | 한국수력원자력 주식회사 | Radiation source imaging apparatus to decommissioning area in nuclear power plant |
Citations (2)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
JPH11337646A (en) * | 1998-05-29 | 1999-12-10 | Toshiba Corp | Radiation semiconductor detector, radiation semiconductor detector array and collimator installating device |
JP2001346795A (en) * | 2000-06-07 | 2001-12-18 | Canon Inc | Method of evaluating alignment between x-ray source and scattered ray removing grid, and feedback method, feedback device, x-ray radiography and x-ray radiographic device |
-
2008
- 2008-09-24 JP JP2008244115A patent/JP5249694B2/en not_active Expired - Fee Related
-
2009
- 2009-08-12 WO PCT/JP2009/064251 patent/WO2010035586A1/en active Application Filing
Patent Citations (2)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
JPH11337646A (en) * | 1998-05-29 | 1999-12-10 | Toshiba Corp | Radiation semiconductor detector, radiation semiconductor detector array and collimator installating device |
JP2001346795A (en) * | 2000-06-07 | 2001-12-18 | Canon Inc | Method of evaluating alignment between x-ray source and scattered ray removing grid, and feedback method, feedback device, x-ray radiography and x-ray radiographic device |
Cited By (1)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
CN113557448A (en) * | 2019-03-29 | 2021-10-26 | 深圳帧观德芯科技有限公司 | Imaging method |
Also Published As
Publication number | Publication date |
---|---|
JP2010078350A (en) | 2010-04-08 |
JP5249694B2 (en) | 2013-07-31 |
Similar Documents
Publication | Publication Date | Title |
---|---|---|
US8294110B2 (en) | Method for improved correction of SiPM non-linearity in multiplexed radiation detectors | |
JP3928647B2 (en) | Radiation imaging apparatus and nuclear medicine diagnostic apparatus using the same | |
US8809793B2 (en) | System and method for pixelated detector calibration | |
JP6684787B2 (en) | PET detector timing calibration | |
Sanchez et al. | Design and tests of a portable mini gamma camera | |
JP2005533245A (en) | Gamma-ray detectors for positron emission tomography (PET) and single photon emission computed tomography (SPECT) | |
JP2018537658A (en) | Radiation detector for combined detection of low energy radiation quanta and high energy radiation quanta | |
JP4972474B2 (en) | Nuclear medicine diagnostic apparatus, image processing method, and gamma camera | |
WO2011058891A1 (en) | Radiation imaging device and nuclear medicine diagnostic device using same | |
KR101898794B1 (en) | Photon-counting detector | |
US9383457B2 (en) | Detector for detecting the traces of ionizing particles | |
JP5249694B2 (en) | Nuclear medicine diagnostic apparatus and adjustment method thereof | |
US10545249B2 (en) | Radiation detection device | |
JP4902753B2 (en) | Radiation imaging apparatus and collimator position estimation method | |
JP2019020334A (en) | Wave height frequency distribution acquisition device, wave height frequency distribution acquisition method, wave height frequency distribution acquisition program, and radiation imaging device | |
JP2016223997A (en) | Radiation camera | |
US9612344B2 (en) | Positron emission tomography and single photon emission computed tomography based on intensity attenuation shadowing methods and effects | |
JP5672061B2 (en) | Positron emission tomography system | |
JP2011107122A (en) | Gamma ray detector module with variable light guide thickness | |
JP2012177555A (en) | Radiation imaging apparatus | |
WO2014054755A1 (en) | Radiation imaging device | |
JP4479699B2 (en) | Gamma camera | |
Dokhale et al. | Intrinsic spatial resolution and parallax correction using depth-encoding PET detector modules based on position-sensitive APD readout | |
US20050029461A1 (en) | Gamma camera using rotating scintillation bar detector and method for tomographic imaging using the same | |
WO2024048515A1 (en) | Image acquisition device and image acquisition method |
Legal Events
Date | Code | Title | Description |
---|---|---|---|
121 | Ep: the epo has been informed by wipo that ep was designated in this application |
Ref document number: 09816003 Country of ref document: EP Kind code of ref document: A1 |
|
NENP | Non-entry into the national phase |
Ref country code: DE |
|
122 | Ep: pct application non-entry in european phase |
Ref document number: 09816003 Country of ref document: EP Kind code of ref document: A1 |