JP5497304B2 - Tomography equipment - Google Patents

Description

  The present invention relates to a tomography apparatus that performs tomography by detecting radiation emitted from a radioisotope in a subject or radiation transmitted through a subject.

  In recent years, tomographic apparatuses have been widely used to obtain information inside a living body (subject). Examples of the tomography apparatus include an X-ray computed tomography (X-ray CT) apparatus, a magnetic resonance imaging apparatus, a SPECT (single photon emission CT) apparatus, and a positron tomography (PET) apparatus. An X-ray CT apparatus irradiates a cross section of a living body with a narrow X-ray beam from multiple directions, detects transmitted X-rays, and calculates the spatial distribution of the degree of X-ray absorption in the cross section by a computer. Calculated and imaged. In this way, morphological abnormalities inside the living body, such as bleeding foci, can be grasped (see, for example, Patent Documents 1 and 2).

In addition, since the PET apparatus can obtain precise information on the function information in the subject, development has been proceeding in recent years. In a diagnostic method using a PET apparatus, first, a test drug labeled with a positron nuclide is introduced into a subject by injection or inhalation. The test drug introduced into the subject is accumulated in a specific part having a function corresponding to the test drug. For example, when a saccharide test drug is used, it is selectively accumulated at sites with high metabolism such as cancer cells. At this time, positrons are emitted from the positron nuclide of the test agent, and when the emitted positrons and surrounding electrons are combined and annihilated, two gamma rays (so-called annihilation gamma rays) are emitted in a direction of about 180 degrees from each other. The Therefore, coincidence detection is performed with a radiation detector arranged around the subject, and the image is regenerated by a computer or the like, thereby obtaining radioisotope distribution image data in the subject. As described above, since the function information in the body of the subject can be obtained with the PET apparatus, the pathology of various intractable diseases can be elucidated.
(For example, refer to Patent Document 3).

JP 2005-253572 A JP 2006-225 A JP 2006-119095 A

  By the way, if there is a positional shift in the arrangement of the detectors that detect radiation, an error due to this positional shift is included in an image regenerated by a computer or the like, and the accuracy and resolution of the regenerated image deteriorate.

  When the resolution is relatively low, the effect on the finally regenerated image is unlikely to be a problem even if the detector arrangement includes misalignment, but the resolution increases as the device performance improves. As a result, the influence of the displacement of the detector on the image increases.

  SUMMARY OF THE INVENTION An object of the present invention is to provide a tomography apparatus in which the accuracy and resolution of a regenerated image is improved by performing correction based on the positional deviation of the detector with respect to detection data of the detector. To do.

A tomography apparatus according to one aspect of the present invention is provided on a base around a subject installation region, and a plurality of detectors that detect radiation emitted from the subject or radiation that passes through the subject; A correction unit that corrects position information included in the detection result of the detector by using the installation position deviation of the detector; and a regeneration unit that regenerates an image based on the detection result corrected by the correction unit. The regenerating unit is configured to regenerate the image by regenerating data in which the detection results are arranged in a pixel shape, and the correcting unit is configured to regenerate the image according to the installation position deviation. The pixel is divided into sub-pixels, and the average value of the values in the sub-pixels is used as the value in the pixels, thereby correcting the position information included in the detection result of the detector .

  Further, the regeneration unit may use a sinogram as data arranged in a pixel shape.

The correction unit may correct the position information included in the detection result of the detector using a ratio between the displacement of the installation position of the detector and the width of the detection element included in the detector.

  According to the present invention, it is possible to provide a tomographic apparatus in which the accuracy and resolution of a regenerated image can be improved by correcting the detection data of the detector based on the positional deviation of the detector. Is obtained.

  Embodiments to which the tomography apparatus of the present invention is applied will be described below.

  The tomography apparatus of the present invention is an X-ray computed tomography (X-ray CT) apparatus, magnetic resonance imaging apparatus, SPECT (Single Photon Emission CT) apparatus, which is a general tomography apparatus, The present invention can be applied to all positron tomography (PET) apparatuses. Here, an embodiment applied to a PET apparatus using a semiconductor element having a particularly high resolution will be described.

  FIG. 1 is a block diagram illustrating a configuration of a PET apparatus according to an embodiment.

  The PET apparatus 10 of the present embodiment is arranged around the subject S so as to surround the subject S 360 degrees in plan view, and detects the gamma rays detected by the radiation detector 11 and the radiation detector 11. An information processing unit 12 that processes data and regenerates image data of the position of the positron nuclide RI in the body of the subject S, a display unit 13 that displays image data, and the like, a subject S and a radiation detector 11, a control unit 14 that controls movement of the control unit 11, an input / output unit 15 that includes a terminal that sends instructions to the information processing unit 12 and the control unit 14, a printer that outputs image data, and the position data of the radiation detector 11. And a position data storage unit 100 to be stored.

  The radiation detector 11 is a plurality of detectors that are arranged on an annular base 16 around the installation area of the subject S and detect γ rays emitted from the subject S.

  The radiation detector 11 includes a semiconductor detection unit 20 and a detection circuit 30, and the semiconductor detection unit 20 is arranged so that the incident surfaces of the gamma rays γa and γb face the subject S. Note that a test drug labeled with a positron nuclide RI is introduced into the subject S in advance.

The semiconductor detector 20 detects two gamma rays γ _a and γ _b that are generated simultaneously when the positron from the positron nuclide RI disappears. Since the two gamma rays γ _a and γ _b are emitted at an angle of approximately 180 degrees, they enter the semiconductor detector 20 of the radiation detector 11 facing each other with the subject S interposed therebetween. Each of the two semiconductor detector 20 the gamma ray gamma _a, is gamma _b incident sends gamma gamma _a, an electric signal (detection signal) generated by the incidence of the gamma _b to the detection circuit 30.

The detection circuit 30 includes, for example, an ASIC for amplifying the detection signal and an FPGA that sends the incident time and the incident position to the information processing unit 12 as digital data. The detection circuit 30 determines the time (incident time) and the incident position where the gamma rays γ _a and γ _b are incident on the detection element from the detection signal detected by the semiconductor detection unit 20, and detects the detection signal, the incident time, and the incident position. Is sent to the information processing unit 12. The data representing the incident position includes the identification number of the detection element.

  The information processing unit 12 performs coincidence detection and image data regeneration by an image regeneration algorithm based on the detection data input from the detection circuit 30. In the coincidence detection, when there are two detection data whose incident times are equal to or less than the set time width, the detection data is determined to be valid and becomes coincidence detection data, while detection data whose incident times do not match is invalidated. Determine and discard.

  The information processing unit 12 creates a sinogram using the coincidence detection data and position information (incidence position information) included in the coincidence detection data, and uses a sinogram to generate a predetermined image regeneration algorithm (for example, expectation value maximization). The image data is regenerated based on (Expectation Maximization) method. The display unit 13 displays the image data regenerated in response to a request from the input / output unit 15. The coincidence detection and image regeneration algorithm is not limited to the above-described items, and known items can be used.

  The position data storage unit 100 stores the identification number of the detection element included in the semiconductor detection unit 20 in the radiation detector 11 and the coordinate value of the incident surface of the detection element in association with each other. Although the coordinates and coordinate values will be described later, the position data storage unit 100 also includes a position shift value indicating a shift of each installation position of the radiation detector 11 and an identification number of the detection element and an incident surface of the detection element. Stored in association with coordinate values.

  With the above configuration and operation, the PET apparatus 10 detects gamma rays from the positron nuclide RI selectively located in the body of the subject S, and reproduces the image data of the distribution state of the positron nuclide RI through creation of a sinogram. To do. The PET apparatus 10 according to the present embodiment is mainly characterized by the radiation detector 11. Hereinafter, the radiation detector 11 will be described in detail.

  When regenerating the image data, the information processing unit 12 corrects the sinogram using the position shift value indicating the shift of each installation position of the radiation detector 11 stored in the position data storage unit 100, and performs correction. Image data is regenerated based on the later sinogram. Note that each misregistration value of the radiation detector 11 may be accurately measured, for example, when the PET apparatus 10 is assembled, and data representing the misregistration value is stored in the position data storage unit 100.

The radiation detectors 11 _{1 to} 11 ₈ of the PET apparatus 10 are arranged around the subject S over 360 degrees. Each of the radiation detectors 11 _{1 to} 11 ₈ is provided with a semiconductor detector 20 on the subject S side. Here, the body axis direction of the subject S (the direction penetrating the paper surface) is defined as the Z-axis direction (the direction from the back side to the front side is the + Z direction, and the direction from the front side to the back side is the −Z direction).

In addition, the x-axis with the width direction of the semiconductor detector 20 of the radiation detector 11 ₁ as the + x direction is taken, and the y-axis with the depth direction of the semiconductor detector 20 of the radiation detector 11 ₁ as the + y direction is taken.

Here, the radiation detector 11 may be movable relative to the subject S in the Z-axis direction. Although eight radiation detectors 11 _{1 to} 11 ₈ are shown in FIG. 1, these numbers are merely examples, and the number of radiation detectors is appropriately selected.

Figure 2 is a perspective view showing a structure of a semiconductor detector 20 included in the radiation detector 11 ₁ is a view of the semiconductor detector substantially from the incident side of the gamma ray. FIG. 3 is a schematic plan view of the semiconductor detection unit. FIG. 4 is a diagram for explaining the operation of the semiconductor detection element.

2 to 4, the semiconductor detection unit 20 includes a wiring board 21, two semiconductor detection element arrays 22 a and 22 b disposed on the wiring board 21, and an output of the semiconductor detection unit 20 as a detection circuit ( It includes a connector 29 for sending out to the detection circuit 30) shown in FIG. Each of the semiconductor detection element arrays 22a and 22b is provided with a first electrode 25 and a second electrode 26 on two surfaces of a substantially planar semiconductor crystal body 24 perpendicular to the Z-axis direction. Each of the semiconductor detection element arrays 22a and 22b includes a semiconductor detection element 23 _{1 in which the} semiconductor crystal bodies 24 are separated from each other in the X-axis direction by grooves 24b formed in the surface on the second electrode 26 side along the Y-axis direction. -23 ₆ and guard members 28a, 28b and the like provided on both sides in the arrangement direction of the semiconductor detection elements 23 ₁ -23 ₆ (outside in the X-axis direction). The semiconductor crystal body 24 and the guard members 28a and 28b are made of the same material and are integrated. 2 and 3, the separation position is indicated by a broken line or a solid line between the adjacent semiconductor detection elements 23 ₁ to 23 ₆ , but the separation position passes through the center of the groove portion in the X-axis direction and is in the Y-axis. It extends along the direction. The interval between the semiconductor detection elements 23 ₁ to 23 ₆ corresponds to the distance between adjacent separation positions.

In FIG. 2, the direction in which the semiconductor detection elements 23 ₁ to 23 ₆ in the semiconductor detection unit 20 included in the radiation detector 11 ₁ are arranged is defined as the arrangement direction (X-axis direction), and two semiconductor detection element arrays 22a. , 22b is referred to as the depth direction (Y-axis direction), and the direction in which the wiring board 21 and the semiconductor detection element arrays 22a, 22b are stacked is referred to as the stacking direction (Z-axis direction). Here, the number of the semiconductor detection elements 23 ₁ to 23 _{6 in} each of the semiconductor detection element arrays 22a and 22b is six, but the number is not particularly limited as long as it is two or more.

Further, FIG. 2 shows a state where the semiconductor detection unit 20 included in the radiation detector 11 ₁ is mounted on the wiring board 21, the semiconductor detection unit 20 included in the radiation detector 11 _{2-11 8} also identical Is mounted on the wiring board 21.

Each of the semiconductor detection elements 23 ₁ to 23 ₆ includes a semiconductor crystal body 24, a first electrode portion 25 formed on the upper surface of the semiconductor crystal body 24, and a second electrode portion 26 formed on the lower surface.

The material of the semiconductor crystal 24 is, for example, cadmium telluride (CdTe), Cd _1-x Zn _x Te (CZT), thallium bromide (TlBr), silicon or the like sensitive to gamma rays with energy of 511 keV. Can be mentioned. Moreover, the dopant for controlling electroconductivity etc. may be contained in these materials.

The semiconductor crystal 24, any one of the semiconductor detection device 23 ₁ to 23 _6, the width, depth, and width are equally set, respectively. For example, the semiconductor crystal body 24 has dimensions of a width (X-axis direction) of 1.2 mm, a depth (Y-axis direction) of 5 mm, and a thickness of about 1 mm. The semiconductor crystal body 24 forms a semiconductor crystal using a Bridgeman method, which is a semiconductor crystal growth method, or a moving heating method, and is cut out in a predetermined crystal orientation.

  The first electrode 25 is a conductive film that substantially covers the upper surface of the semiconductor crystal body 24. A negative bias voltage Vb is applied to the first electrode 25 to serve as a cathode. In the case where the semiconductor crystal body 24 is made of CdTe, for example, Pt is used for the first electrode 25. The bias voltage Vb is a DC voltage and is set to, for example, −60V to −1000V. The first electrode 25 is continuously formed over the entire upper surface of the six semiconductor crystal bodies 24. However, although the first electrode 25 covers the upper surfaces of the guard members 28a and 28b, this is not essential. The bias voltage is supplied from the outside of the wiring board 21 via the wiring pattern 36 and the wire wiring 35.

The second electrode 26 includes a conductive film that substantially covers the lower surface between the groove 24 b and the groove 24 b of the semiconductor crystal body 24. The second electrode 26 functions as an anode. Note that In (indium) is implanted into the semiconductor crystal body 24 on the second electrode 26 side. When the semiconductor crystal body 24 is made of CdTe, for example, Au is used for the second electrode 26. The second electrode 26 is provided in each of the semiconductor detection elements 23 ₁ to 23 ₆ , and the second electrodes 26 adjacent to each other are electrically insulated from each other. The second electrode 26 is electrically connected to the connector 29 via a wiring pattern (not shown) provided on the wiring board 21 via the conductive adhesive layer 27 and the pad electrode 32.

  The second electrode 25 is also formed on the lower surfaces of the guard members 28 a and 28 b and is grounded via the conductive adhesive layer 27 and the pad electrode 32. By doing so, it is possible to avoid mixing the detection signal with noise as the electron-hole pair generated by the gamma rays incident on the guard members 28a and 28b flows to the ground potential. Note that. The structure of the lower surface of the guard members 28a and 28b is not essential, and the guard member may not be provided when a material other than the semiconductor crystal is used.

  The conductive adhesive layer 27 is composed of a metal powder or carbon filler selected from Au, Ag, Cu, and alloys thereof, and a resin-made conductive adhesive, such as a conductive paste or an anisotropic adhesive. Can be used.

  As shown in FIG. 4, in the semiconductor detection element 23, when gamma rays γ are incident on the semiconductor crystal body 24, the number of electron-hole pairs corresponding to the energy of the gamma rays is generated. Since an electric field is applied to the semiconductor crystal body 24 in the direction from the second electrode 26 to the first electrode 25, holes are attracted to the first electrode 25 and electrons are attracted to the second electrode 26. As a result, a detection signal is generated and sent to the detection circuit. It should be noted that even when gamma rays γ are incident on the semiconductor crystal body 24, the generation of electron-hole pairs is stochastic, and the electron-hole pairs are not generated in the semiconductor crystal body 24 and may penetrate.

2 and 3, the guard members 28a and 28b are provided at both ends of the semiconductor detection element arrays 22a and 22b in the arrangement direction, and protect the outer surfaces of the semiconductor detection elements 23 ₁ and 23 ₆ . The material of the guard members 28a and 28b is not particularly limited as long as it is a material that does not electrically connect the first electrode 25 and the second electrode 26. However, adverse effects due to stress caused by a difference in thermal expansion coefficient due to heating in the manufacturing process can be avoided. In this respect, it is preferable that the same material as that of the semiconductor crystal body 24 is included. Furthermore, the guard members 28 a and 28 b are particularly preferably formed from the same crystal plate continuous with the semiconductor crystal body 24. Thereby, it is not necessary to bond the guard members 28a and 28b to the semiconductor crystal body 24, and the process of forming the guard members 28a and 28b can be simplified. However, it is preferable to form a groove 24c between the guard member 28a and the semiconductor detection element 23 ₁ and between the guard member 28b and the semiconductor detection element 23 ₆ . Thereby, it is possible to avoid the electron-hole pair generated by the gamma rays incident on the guard members 28a and 28b from flowing into the second electrodes 26 of the adjacent semiconductor detection elements 23 ₁ and 23 ₆ . The groove 24c may be formed in the same shape as the groove 24b. The width of the guard members 28a and 28b is set to a predetermined width based on the interval between the semiconductor detection elements 23 ₁ to 23 ₆ , which will be described in detail later.

  Although the two semiconductor detection element arrays 22a and 22b both have the above-described structure, the arrangement on the wiring board 21 is different as shown in FIG. Hereinafter, the semiconductor detection element array disposed on the side on which the gamma rays γ are incident is referred to as a first semiconductor detection element array 22a, and the semiconductor detection element array disposed on the back side is referred to as a second semiconductor detection element array 22b. However, when it is not necessary to distinguish between the two, they are simply referred to as semiconductor detection element arrays 22a and 22b.

The first semiconductor detection element array 22a and the second semiconductor detection element array 22b are characterized by the positional relationship between the semiconductor detection elements. Both the first semiconductor detection element array 22a and the second semiconductor detection element array 22b are arranged at predetermined intervals PT in the arrangement direction (X-axis direction). Further, the left side surface 23 _1a of the semiconductor detection element 23 ₁ of the first semiconductor detection element array 22a is arranged at a position spaced apart from the reference line Xa-Xa by the interval PT. On the other hand, the left side surface 23 _1b of the semiconductor detection element 23 ₁ of the second semiconductor detection element array 22a is arranged at a position spaced apart from the reference line Xa-Xa by the interval PT / 2. In this manner, with each of the semiconductor detection device 23 ₁ to 23 ₆ of the first semiconductor detection device array 22a, and each of the semiconductor detection device 23 ₁ to 23 ₆ of the second semiconductor detection device array 22b, to the array directions Displaced by PT / 2.

Further, the guard member 28a at the left end of the drawing of the first semiconductor detection element array 22a is set to a width equal to the interval PT, and the guard member 28b at the right end of the drawing is set to ½ of the interval PT. On the other hand, the guard member 28b at the left end of the drawing of the second semiconductor detection element array 22b is set to ½ of the interval PT, and the guard member 28a at the right end of the drawing is set to the interval PT. By providing the guard members 28a and 28b as described above, the widths (lengths in the arrangement direction) of the semiconductor detection element arrays 22a and 22b are made equal to each other. The semiconductor detection element arrays 22a and 22b are arranged so that the left end of the drawing coincides with the reference line Xa-Xa. By doing so, when the two semiconductor detection element arrays 22a and 22b are arranged in the assembly process of the semiconductor detection unit 20, the semiconductor detection elements 23 ₁ to 23 ₆ are in the above-described positional relationship. Positioning is greatly facilitated. Although it is preferable to provide the guard members 28a and 28b as described above, they are not essential to the present invention.

  In the above description, the positioning is performed at the left end of the semiconductor detection element arrays 22a and 22b. However, the positioning may be similarly performed with the reference line Xb-Xb at the right end of the paper.

Thus, the first semiconductor detection element array 22a and the second semiconductor detection element array 22b are shifted in the arrangement direction by 1/2 of the interval PT between the semiconductor detection elements 23 ₁ to 23 ₆ and arranged in the depth direction. As a result, the semiconductor detector 20 has a spatial resolution at the center of the field of view (that is, a gamma ray incident from substantially the front of the semiconductor detector 20) more than the spatial resolution at the center of the field of view when only one semiconductor detector array is provided. Will also improve. The degree of improvement is that the spatial resolution at the center of the visual field of the semiconductor detection unit 20 is substantially equal to the spatial resolution at the center of the visual field when the width of the semiconductor detection element is set to ½.

  Furthermore, the first semiconductor detection element array 22a and the second semiconductor detection element array 22b are arranged in the depth direction (Y-axis direction), and each semiconductor detection element array 22a, 22b is provided with only one semiconductor detection element array. The depth of the semiconductor detection element is made shorter than the case where the depth is set equal to the sum of the depths of the semiconductor detection element arrays 22a and 22b. As a result, as described above with reference to FIG. 2B, the spatial resolution in the peripheral portion of the visual field of the semiconductor detection unit 20 can be improved.

  Furthermore, the detection efficiency of the semiconductor detection unit 20 is the same as when only one semiconductor detection element array is provided and the depth of the semiconductor detection element is set equal to the sum of the depths of the semiconductor detection element arrays 22a and 22b. The detection efficiency can be maintained. Although these points are shown by the simulation mentioned later, the effect is demonstrated easily below.

  FIG. 5 is a schematic front perspective view of the semiconductor detection unit. In FIG. 5, the first electrode, the second electrode, and the conductive adhesive layer are not shown.

Referring to FIG. 5, when the semiconductor detector 20 is seen through from the front (gamma ray incident side), the semiconductor detectors 23 ₁ to 23 _{6 of} the first semiconductor detector array 22a and the semiconductors of the second semiconductor detector array 22b. The detection elements 23 ₁ to 23 ₆ overlap with each other, and it appears as if semiconductor detection elements having a half of the interval PT are arranged in a line. The gamma rays include those that generate electron-hole pairs in the first semiconductor detection element array 22a and those that pass through the first semiconductor detection element array 22a and generate electron-hole pairs in the second semiconductor detection element array 22b. Probably present. Therefore, for gamma rays, the semiconductor detection element arrays 22a and 22b are substantially the same as a half of the semiconductor detection elements formed in a line as shown in the perspective view of FIG. From this, it is fully inferred that the semiconductor detector 20 approaches the spatial resolution when the spatial resolution at the center of the visual field is set to ½ of the interval PT of the semiconductor detection elements. Further, this is presumed to contribute to the improvement of the spatial resolution around the visual field of the semiconductor detector 20.

Here, although described semiconductor detection device 23 ₁ to 23 ₆ in the semiconductor detection unit 20 included in the radiation detector 11 _1, the radiation detector ₁₁ 2-11 ₈ can also be considered as well.

Incidentally, if there is a positional deviation in the position of one of the semiconductor detector 20 of the radiation detector 11 ₁ to 11 _8, the accuracy of the image to be regenerated is lowered. This is because a position shift is included in the sinogram generated in the stage of regenerating the image.

  For this reason, in the PET apparatus according to the present embodiment, the position data storage unit 100 shown in FIG. 1 includes a position shift that indicates a shift in the installation position of each radiation detector 11 (a shift in the installation position of the semiconductor detection unit 20). The value is stored, and when the sinogram is created, the detection signal of the semiconductor detection unit 20 is corrected based on the positional deviation value.

The locus value of the incident surface of the semiconductor detector 20 included in each of the radiation detectors 11 _{1 to} 11 ₈ of the PET apparatus 10 is the origin (x, y, z) = the center of the area where the subject S is installed = As (0, 0, 0), it is accurately measured during assembly. The measured coordinate values are stored in the position data storage unit 100 as coordinate values (x, y, z) of the semiconductor detection unit 20 included in each of the radiation detectors 11 _{1 to} 11 ₈ .

Here, as described above, all the semiconductor detection units 20 included in the radiation detectors 11 _{1 to} 11 ₈ are mounted on the same wiring board 21. For this reason, the positional deviation about the z coordinate value among the coordinate values (x, y, z) can be ignored. Further, the positional deviation of the semiconductor detection unit 20 in the depth direction is only a slight difference with respect to the optical path length and can be ignored. However, the positional deviation in the width direction of the semiconductor detector 20 is a positional deviation in the x-coordinate value and the y-coordinate value of the incident surface, and the positional deviation of the semiconductor crystal body 24 whose incident surface width (X-axis direction) is 1.2 mm. It becomes. For example, this positional deviation has a large influence on the detection signal even if only a few millimeters of the comma occurs.

  For this reason, in the tomography apparatus of the present embodiment, the detection signal output from the semiconductor detection unit 20 is corrected using the position shift value of the incident surface (position shift amount in the width direction of the semiconductor detection unit 20). .

Specifically, the radiation detector 11 represents the data representing the position of the incident surface of each semiconductor detector 20 in (x, y) coordinates, the coordinate value, and the positional deviation value representing the deviation between the coordinate value and the design value. ₁ to 11 ₈ for all of the semiconductor detection unit 20 included in the stored in the position data storage unit 100 is used to correct the incident position of detection signals.

In other words, the information processing unit 12 uses each of the radiation detectors 11 _{1 to} 11 ₈ by using the position shift value (position shift amount in the width direction of the semiconductor detection unit 20) stored in the position data storage unit 100. The incident position of the detection data (detection signal, incident time, and incident position) detected in (1) is corrected.

Here, each of the radiation detectors 11 _{1 to} 11 ₈ includes a semiconductor detection unit 20, and each semiconductor detection unit 20 includes semiconductor detection element arrays 22 a and 22 b, and the radiation detectors 11 _{1 to} 11 _8. semiconductor detection device array 22a inside the respective, for 22b and semiconductor positional deviation of the detector elements 23 ₁ to 23 ₆ are unlikely to occur, as the positional deviation value, tomography apparatus the displacement of the radiation detector 11 ₁ to 11 ₈ It is measured as a deviation from the design value at the time of assembly.

The incident surface of the semiconductor detection unit 20 is a set of incident surfaces of the semiconductor detection elements 23 ₁ to 23 ₆ included in the semiconductor detection element arrays 22a and 22b of each semiconductor detection unit 20.

For this reason, the positions of the incident surfaces of all the semiconductor detection elements 23 ₁ to 23 ₆ are defined within the range of values that can be taken by the coordinate values (x, y), the x coordinate value, and the y coordinate value, and the semiconductor on which the γ rays are incident. The incident positions of the detection elements 23 ₁ to 23 ₆ are corrected.

  FIG. 6 is a diagram illustrating a procedure of sinogram correction processing and tomographic image regeneration processing in the tomography apparatus according to the present embodiment. This process is a process executed by the information processing unit 12.

The information processing unit 12 determines whether detection data is input from the radiation detectors 11 _{1 to} 11 ₈ (step S1). This is because it is necessary to correct the positional deviation of the positional information included in the detection data. In addition, the process of step S1 is repeatedly performed until it reaches the measurement time set by the user.

The information processing unit 12 corrects the positional deviation of the position information included in the detection data, and creates a sinogram (step S2). Specifically, using the identification number of the detection element included in the detection data input from the radiation detectors 11 _{1 to} 11 _{8, the} position shift value stored in the position data storage unit 100 is read, and the read position shift The incident position included in the detection data is corrected using the value. Then, a sinogram is created using the detection data with the incident position corrected.

Next, the information processing unit 12 regenerates a tomographic image based on the sinogram (step S3). Since the sinogram used here is created based on the detection data in which the incident position is corrected in the process of step S2, the resolution and accuracy caused by the displacement of the installation positions of the radiation detectors 11 _{1 to} 11 ₈ are reduced. The removed tomographic image can be regenerated.

  FIG. 7 is a diagram conceptually showing a method of sinogram correction processing in the tomography apparatus of the present embodiment.

Here, in case of detecting the radiation detector 11 ₁ and 11 ₅ γ rays shown in FIG. 1, and when there is no positional deviation in the radiation detector 11 ₅ (FIG. 7 (a)), if there is a positional deviation ( FIG. 7B will be described.

FIG. 7 shows the sinogram as 9-pixel data. Among the nine pixels, the vertical direction represents an angle formed by the direction of γ rays and the x axis (an angle given in a counterclockwise direction with respect to the x axis in FIG. 1), and this is referred to as an angular direction. Of the nine pixels, the horizontal direction is referred to as a radial direction and represents a direction perpendicular to the angular direction. For convenience of description, the radiation detector 11 ₁ comprises a semiconductor detecting element 1-3, the radiation detector 11 ₅ is intended to include semiconductor detecting element 4-6.

As shown in FIG. 7A, when the radiation detectors 11 ₁ and 11 ₅ are not misaligned, 5 counts are obtained by the semiconductor detection element 1-4 and 3 counts are obtained by the semiconductor detection element 2-5 due to the incidence of γ rays. Assume that the semiconductor detection element 3-6 detects a 1-count detection signal.

In this case, the information processing unit 12 calls the position data storage unit 100, the radiation detector 11 ₁ and 11 position data of ₅ (x, y) a. Here, the position of the semiconductor detection element 1 is (x1, y1), the position of the semiconductor detection element 2 is (x2, y1), the position of the semiconductor detection element 3 is (x3, y1), and the position of the semiconductor detection element 4 is ( x4, y5), the position of the semiconductor detection element 5 is (x5, y5), and the position of the semiconductor detection element 6 is (x6, y5). First, the information processing unit 12 calculates the angular direction coordinate θ using Equation 1. As shown in FIG. 1, the angle direction coordinate θ is shown as an angle formed by the direction in which the gamma ray γa is emitted with respect to the positive direction of the y axis (counterclockwise angle with respect to the positive direction of the y axis). This is the same as the angle formed by the direction in which the gamma ray γb is emitted with respect to the negative direction of the y-axis.

Next, the information processing unit 12 calculates the radial direction coordinate s using Equation 2.

The obtained counts are assigned to the pixels of the sinogram corresponding to the obtained angular direction coordinate θ and radial direction coordinate s. That is, 5 counts are given to the pixels (θ1, s1), 3 counts are given to the pixels (θ1, s2), and 1 count is given to the pixels (θ1, s3).

On the other hand, in FIG. 7 (b), the radiation detector 11 ₅ to the displacement E corresponding to half the width W of the semiconductor crystal 24 (= W / 2) is referred to as being one. In this case, in the tomography apparatus of the present embodiment, the sinogram pixels are divided into an integer number obtained by dividing the width W of the semiconductor crystal body 24 by E, and the angular direction coordinates and the moving radius are divided according to the above-described processing. Determine the direction coordinates.

  Here, since W / E = W / (W / 2) = 2, the pixel is divided into two subpixels.

Data call from the position data storage unit is a radiation detector 11 ₁ is the same, the radiation detector 11 _5, the position of the semiconductor detection element 4 (x4 + E, y5), the position of the semiconductor detection element 5 (x5 + E, y5), the position of the semiconductor detection element 6 is (x6 + E, y5).

  First, the information processing unit 12 calculates an angular direction coordinate θ ′ having a positional shift using Equation 3.

Next, the information processing unit 12 calculates the radial direction coordinate s ′ according to Equation 4.

The obtained counts are assigned to two subpixels of the sinogram corresponding to the angular direction coordinate θ ′ and the radial direction coordinate s ′ obtained by the calculation. That is, 5 counts for 2 subpixels in (θ1 ′, s1 ′), 3 counts for 2 subpixels in (θ1 ′, s2 ′), and 2 subpixels in (θ1 ′, s3 ′) 1 count is given. As a result, values of “5”, “5”, “3”, “3”, “1”, “1”, “0” are given to the seven subpixels in the radial direction. Note that the value of the leftmost subpixel in the figure is “0” because there is no count value.

  Here, when there is no positional deviation between the semiconductor detection element 1 and the semiconductor detection element 4, since x1 = x4, (x1-x4) / (y1-y5) = 0.

  In addition, since the width W of the semiconductor crystal body 24 is 1.2 mm and E = W / 2, when the semiconductor detection element 1 and the semiconductor detection element 4 are misaligned, (x1-x4-E) /(Y1-y5)≈0.005.

  Therefore, here, since it can be handled as θ1≈θ1 ′, the three pixels ((θ1, s1)) defined by the original angular direction coordinate θ and the radial direction coordinate s are used as the values given to the subpixels. , (Θ1, s2), (θ1, s3) pixels), the pixels of (θ1, s1), (θ1, s2), (θ1, s3) are 4 , 0.5 is given. In this way, a sinogram using the detection data corrected according to the positional deviation value can be created.

As described above, in the tomography apparatus according to the present embodiment, the information processing unit 12 as a regeneration unit regenerates a tomographic image by regenerating a sinogram in which detection results of the radiation detector 11 are arranged in a pixel shape. Is configured to do. The information processing unit 12 as correction unit is divided into an integer number of sub-pixels obtained each sinogram pixel from the ratio of the width of the semiconductor crystal body 24 as the installation position shift and the detection elements of the radiation detector 11 ₅ and, and further by a value within the mean of the values in the sub-pixels pixels included in each pixel, it corrects the detection result of the radiation detector 11 _5.

  FIG. 8 is a diagram showing a tomographic image obtained by the tomographic apparatus according to the present embodiment.

  FIG. 8A is a diagram showing a tomographic image regenerated from a sinogram created without correcting the misalignment of the detection data, and FIG. 8B is created by correcting the misalignment of the detection data. It is a figure which shows the tomographic image regenerated from a sinogram.

  Comparing FIGS. 8A and 8B, the tomographic image using the sinogram created using the corrected detection data has a circular central portion, and the resolution and accuracy can be improved by correcting the detection data. It can be seen that it is possible to regenerate a tomographic image from which the decrease in the frequency is removed.

  As described above, according to the tomography apparatus of the present embodiment, the tomographic image is regenerated using the sinogram created by correcting the misalignment of the detection data, so that the tomographic image from which the reduction in resolution and accuracy is removed is reproduced. Can be made.

  Unlike the PET apparatus, the CT apparatus or SPECT apparatus needs to rotate the radiation detection unit 11 with respect to the subject S. In particular, since the radiation detection unit 11 is heavy, there is a possibility that the displacement of the arrangement position due to rotation is added in addition to the positional deviation.

  On the other hand, in the case of a PET apparatus, it is not necessary to rotate the radiation detection unit 11 with respect to the subject S. Therefore, by measuring the positional deviation during assembly and correcting the detection data, There is an advantage that it is easy to suppress a decrease in resolution.

  In the present embodiment, the PET apparatus has been described as an example, but the present invention can also be applied to a SPECT (Single Photon Emission Computed Tomography) apparatus or an X-ray computed tomography apparatus.

  Although the tomographic apparatus of the exemplary embodiment of the present invention has been described above, the present invention is not limited to the specifically disclosed embodiment, and does not depart from the scope of the claims. Various modifications and changes are possible.

It is a block diagram which shows the structure of the PET apparatus of 1st Embodiment. Is a perspective view showing a structure of a semiconductor detector 20 included in the radiation detector 11 _1. It is a schematic plan view of a semiconductor detection part. It is a figure for demonstrating operation | movement of a semiconductor detection element. It is a schematic front perspective view of a semiconductor detection part. It is a figure which shows the procedure of the correction process of a sinogram and the reproduction | regeneration process of a tomographic image in the tomography apparatus of this Embodiment. It is a figure which shows notionally the technique of the correction process of the sinogram in the tomography apparatus of this Embodiment. It is a figure which shows the tomographic image obtained by the tomography apparatus of this Embodiment.

10 PET device 11, 11 ₁ to 11 ₈ radiation detector 12 information processing unit 13 display unit 14 control unit 15 input unit 16 base plate 20, 40, 50 semiconductor detection unit 21 the wiring board 22, 22 ₁ through 22 _n semiconductor detector Element array 22a First semiconductor detection element array 22b Second semiconductor detection element array 23, 23 ₁ to 23 ₆ Semiconductor detection element 24 Semiconductor crystal body 24a Incident surface 24b, 24c Groove 25 First electrode 26 Second electrode 27 Conductive adhesive layer 28a, 28b, 28 _{1 to} 28 _n guard member 29 connector 30 detection circuit 100 position data storage unit

Claims (3)

A plurality of detectors that are disposed on a base around the subject installation area and detect radiation emitted from the subject or radiation that passes through the subject;
A correction unit that corrects position information included in the detection result of the detector using an installation position shift of the detector;
Look containing a regenerating unit for regenerating an image on the basis of the detection result of said corrected by the correction unit,
The regeneration unit is configured to regenerate the image by regenerating data in which the detection results are arranged in pixels.
The correction unit is included in the detection result of the detector by dividing the pixel into sub-pixels according to the installation position shift and setting an average value of the values in the sub-pixels as the value in the pixel. A tomography device that corrects position information . The tomography apparatus according to claim 1 , wherein the regeneration unit uses a sinogram as data arranged in a pixel shape. Wherein the correction unit corrects the positional information included in the detection result of the detector using the ratio of the width of the detection elements included in the detector and the installation position shift of the detector, to claim 1 or 2 The tomography apparatus described.