JP2006239118A - X-ray ct system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To solve the problem that an improvement of an X-ray geometric efficiency is sought because of the problem about the area on a detector occupied by a detector X-ray collimator and a reflector accompanied by the increase in channels, the increase in rows and reduction and minuteness of each channel in an X-ray CT system having a two-dimensional area X-ray detector, and achieve picture quality improvement and exposure reduction of the scan by the X-ray detector without the detector X-ray collimator and the X-ray detector using a scintillator without the reflector in the X-ray CT system having the two-dimensional area X-ray detector. <P>SOLUTION: Compensation for the increase of an X-ray scattered ray due to the absence of the detector X-ray collimator is made by X-ray scatter compensation, and compensation for the increase of the crosstalk of the signals between channels and between rows due to the use of the scintillator without the reflector is made by crosstalk compensation. <P>COPYRIGHT: (C)2006,JPO&NCIPI

Description

  The present invention relates to an X-ray CT (Computed Tomography) apparatus, and relates to image quality improvement and exposure reduction in conventional scanning (axial scanning) or cine scanning or helical scanning.

  Conventionally, in an X-ray CT apparatus using a two-dimensional X-ray area detector having a matrix structure represented by a multi-row X-ray detector 24, a detector X-ray collimator 31 for preventing scattered X-rays is provided as shown in FIG. And a scintillator 24a which is an element that converts X-rays into light with a reflector 24c as a partition wall that prevents crosstalk of detection signals between channels and between columns, and a photodiode 24b that is an element that converts light into electrical signals X-ray detector with However, as the number of X-ray detectors increases, the number of rows increases, the size of each channel shrinks, and the size of the channel decreases, the proportion of the area occupied by the detector X-ray collimator 31 and reflector 24c is the aperture of each channel of the X-ray detector It has become a problem that it has become relatively larger than the area.

  The efficiency E in the X-ray geometry is determined by the following equation (1) as can be seen from FIG.

Here, l _x and l _y are intervals in the channel direction or column direction of the scintillator or photodiode. C _x and _cy are the width of the detector X-ray collimator width or reflector width in the channel direction or the column direction, and as shown in FIG. 14, the detector X-ray collimator width or the reflector width is large. Take the value of. In other words, in order to prevent the decrease in the efficiency E in the X-ray geometry, an improvement in X-ray utilization efficiency has been demanded. The following Patent Document 1 shows the existence of an object that covers the X-ray detector, and indicates that there may be a reduction in efficiency in X-ray geometry.
JP 2000-193750 A

  For this reason, the following problems existed until now.

  When the efficiency in terms of X-ray geometry decreases, the efficiency of the X-ray light conversion element and the light-to-electric signal conversion element decreases. As a result, the electrical signal output to the irradiated X-rays is reduced, so that a sufficient S / N signal cannot be obtained, and the problem of unnecessary X-ray exposure is increased. X-ray geometry of a two-dimensional area X-ray detector with a matrix structure using an X-ray light conversion element such as an X-ray detector without an X-ray collimator or a scintillator that converts X-rays without a reflector into light. The efficiency is better than the conventional X-ray detector. However, a two-dimensional area X-ray detector with a matrix structure using an X-ray light conversion element such as an X-ray detector without an X-ray collimator or a scintillator that converts X-rays without a reflector into light is scattered X-rays and crosstalk. There was a problem related to image quality such as signals.

  Accordingly, an object of the present invention is to provide a two-dimensional area X-ray detector having a matrix structure by an X-ray light conversion element such as an X-ray detector without a detector X-ray collimator, a scintillator that converts X-rays without a reflector into light. The purpose is to realize image quality improvement and exposure reduction of conventional scan (axial scan), cine scan or helical scan in the X-ray CT apparatus.

  The present invention takes measures against the increase of scattered X-rays due to the absence of the detector X-ray collimator by using X-ray scattering correction. In this case, the SN improvement of the X-ray signal is performed by increasing the X-ray dose due to the scattered X-ray, and the image quality of the tomographic image as an X-ray CT apparatus can be expected. Further, crosstalk correction is used to counter the increase in crosstalk of X-ray detector signals between channels and columns due to the absence of a scintillator reflector. As a result, the image quality of the tomographic image can be improved, and as a result, exposure reduction can also be realized.

  In a first aspect, the present invention relates to an X-ray generator and a two-dimensional X-ray area detector having a matrix structure that detects X-rays relative to the X-ray generator, around a rotation center therebetween. X-ray data collection means for collecting X-ray projection data transmitted through the subject in between while rotating, and an image for reconstructing the X-ray projection data collected from the X-ray data collection means In the X-ray CT apparatus, comprising: a reconstructing unit; an image display unit that displays the reconstructed tomographic image; and an imaging condition setting unit that sets imaging conditions for the tomographic imaging. The detector has a structure for capturing scattered X-rays, and the image reconstruction means corrects scattered X-rays.

  In the X-ray CT apparatus according to the first aspect described above, the increase in scattered X-rays due to the structure without the detector X-ray collimator, that is, the structure for capturing the scattered X-rays is corrected by correcting the X-ray scattered radiation. X-ray signal SN is improved by increasing the X-ray dose due to X-rays, improving image quality and reducing exposure.

  In a second aspect, the present invention relates to an X-ray generator, and a two-dimensional X-ray area detector having a matrix structure that detects X-rays relative to the X-ray generator, around a rotation center therebetween. X-ray data collection means for collecting X-ray projection data transmitted through the subject in between while rotating, and an image for reconstructing the X-ray projection data collected from the X-ray data collection means An X-ray CT apparatus comprising: a reconstruction unit; an image display unit that displays the tomographic image that has been reconstructed; and an imaging condition setting unit that sets an imaging condition for the tomographic imaging. The detector includes means for converting X-rays that occupy almost 100% of the X-ray irradiation surface into light, and means for converting light into an electrical signal. It is characterized by performing inter-crosstalk correction.

  In the X-ray CT apparatus according to the second aspect described above, X-ray occupying almost 100% of the X-ray irradiation surface of the X-ray detector due to the absence of a reflector such as a scintillator due to interchannel / column crosstalk correction. By correcting the increase in crosstalk between channels and columns in the X-ray detector signal of the scintillator, which is an element that converts light, it is possible to improve image quality and reduce exposure.

In a third aspect, the present invention relates to an X-ray generator and a two-dimensional X-ray area detector having a matrix structure that detects X-rays relative to the X-ray generator, around a rotation center between them. X-ray data collection means for collecting X-ray projection data transmitted through the subject in between while rotating, and an image for reconstructing the X-ray projection data collected from the X-ray data collection means An X-ray CT apparatus comprising: a reconstruction unit; an image display unit that displays the tomographic image that has been reconstructed; and an imaging condition setting unit that sets an imaging condition for the tomographic imaging. The detector has X-ray data collection means having means for converting X-rays occupying almost 100% of the X-ray irradiation surface into light and means for converting light into an electrical signal, and scattered X-rays. The image reconstructing means includes a scattered X-ray correction and It is characterized by correcting crosstalk between channels and columns.

  In the X-ray CT apparatus according to the third aspect, the increase in scattered X-rays due to the absence of the detector X-ray collimator is corrected by X-ray scattered ray correction, and the scintillator and the like are corrected by inter-channel / column crosstalk correction. By correcting the increase of crosstalk between channels and columns of X-ray detector signals due to the absence of reflectors, image quality can be improved and exposure can be reduced.

  In a fourth aspect, the present invention provides the X-ray CT apparatus according to any one of the first to third aspects, wherein the two-dimensional X-ray area detector is a multi-row X-ray detector. .

  The X-ray CT apparatus according to the fourth aspect can sufficiently improve image quality and reduce exposure even when a multi-row X-ray detector is used as the two-dimensional area X-ray detector.

  In a fifth aspect, the present invention provides the X-ray CT apparatus according to any one of the first to third aspects, wherein the two-dimensional X-ray area detector is a planar X typified by a flat panel X-ray detector. It is a line detector.

  In the X-ray CT apparatus according to the fifth aspect, even when a planar X-ray detector typified by a flat panel is used as the two-dimensional area X-ray detector, the image quality can be sufficiently improved and the exposure can be reduced.

  According to the present invention, the detector has an X-ray detector without an X-ray collimator, a two-dimensional area X-ray detector having a matrix structure by an X-ray light conversion element such as a scintillator that converts X-rays without a reflector into light. In the X-ray CT apparatus, it is possible to improve the image quality and reduce the exposure of conventional scanning (axial scanning), cine scanning, or helical scanning.

  Hereinafter, embodiments according to the present invention will be described in detail. Note that the present invention is not limited thereby.

  FIG. 1 is a configuration block diagram of an X-ray CT apparatus according to an embodiment of the present invention. The X-ray CT apparatus 100 includes an operation console 1, an imaging table 10, and a scanning gantry 20.

  The operation console 1 collects X-ray detector data collected by the scanning device gantry 20 and an input device 2 that receives input from the operator, a central processing device 3 that performs preprocessing, image reconstruction processing, post-processing, and the like. Data acquisition buffer 5, monitor 6 that displays tomograms reconstructed from projection data obtained by preprocessing X-ray detector data, program, X-ray detector data, projection data, and X-ray tomogram And a storage device 7 for storing.

  The imaging table 10 includes a cradle 12 on which a subject is placed and put into and out of the opening of the scanning gantry 20. The cradle 12 is moved up and down and linearly moved by the motor built in the imaging table 10.

  The scanning gantry 20 rotates around an X-ray tube 21, an X-ray controller 22, a collimator 23, a multi-row X-ray detector 24, a DAS (Data Acquisition System) 25, and the body axis of the subject. A rotation controller 26 for controlling the X-ray tube 21 and the like, and a controller 29 for exchanging control signals and the like with the operation console 1 and the imaging table 10. The scanning gantry tilt controller 27 can tilt the scanning gantry 20 forward and backward in the z direction by about ± 30 degrees.

  FIG. 2 is an explanatory diagram of the geometric arrangement of the X-ray tube 21 and the multi-row X-ray detector 24. FIG.

  The X-ray tube 21 and the multi-row X-ray detector 24 rotate around the rotation center IC. When the vertical direction is the y direction, the horizontal direction is the x direction, and the table traveling direction perpendicular thereto is the z direction, the rotation plane of the X-ray tube 21 and the multi-row X-ray detector 24 is the xy plane. The moving direction of the cradle 12 is the z direction.

  The X-ray tube 21 generates an X-ray beam called a cone beam CB. When the direction of the central axis of the cone beam CB is parallel to the y direction, the view angle is 0 degree.

  The multi-row X-ray detector 24 has, for example, 256 detector rows. Each detector row has, for example, 1024 detector channels. FIGS. 17, 18 and 19 show an X-ray detector without a detector X-ray collimator and without a scintillator reflector. FIG. 17 is a multi-row X-ray detector 24 of the present embodiment, FIG. 17 (a) is a side view of the multi-row X-ray detector 24 with the line direction as the line of sight, and FIG. It is sectional drawing which shows the X2 part of a). As shown in FIG. 17, the multi-row X-ray detector 24 of the present embodiment is a multi-row X-ray detector having neither a detector X-ray collimator nor a reflector. The multi-row X-ray detector 24 has a plurality of channels for detecting X-rays arranged in a matrix structure, and includes a scintillator 24a and a photodiode 24b. The scintillator 24a converts X-rays irradiated from the X-ray focal point F of the X-ray tube 21 and transmitted through the subject into light, and the photodiode 24b converts the light converted by the scintillator 24a as an electrical signal. In the multi-row X-ray detector 24, the scintillator 24b is integrally formed without being divided between a plurality of channels. That is, the scintillator 24b is formed so as to be continuous between a plurality of channels. FIG. 18 shows a flat panel type two-dimensional planar X-ray area detector 24 as a modified example, FIG. 18 (a) is a perspective view, and FIG. 18 (b) is an X3 portion of (a). FIG. FIG. 19 shows a variation using a flat panel type two-dimensional planar X-ray area detector 24 having neither a detector X-ray collimator nor a reflector, and FIG. 19 (a) is a perspective view. FIG. 19 (b) is a cross-sectional view showing the X4 portion of (a).

  Projection data collected by irradiation with X-rays is A / D converted from the multi-row X-ray detector 24 by the DAS 25 and input to the data collection buffer 5 via the slip ring 30. The data input to the data collection buffer 5 is processed by the central processing unit 3 according to the program in the storage device 7, reconstructed into a tomographic image, and displayed on the monitor 6.

  FIG. 3 is a flowchart showing an outline of the operation of the X-ray CT apparatus 100 of the present invention.

  In step S1, first, the X-ray tube 21 and the multi-row X-ray detector 24 are rotated around the subject, and a helical scan operation is performed while the cradle 12 is linearly moved along the table. X-ray detector data is collected by adding table linear movement z-direction position Ztable (view) to X-ray detector data D0 (view, j, i) represented by detector row number j and channel number i. To do.

  In step S2, X-ray detector data D0 (view, j, i) is preprocessed and converted into projection data. As shown in FIG. 4, the preprocessing includes step S21 offset correction, step S22 logarithmic conversion, step S23 X-ray dose correction, and step S24 sensitivity correction.

  In step S3, crosstalk correction is performed on the preprocessed projection data D1 (view, j, i). When the projection data subjected to crosstalk correction is D2 (view, j, i) and the crosstalk correction kernel is K (j, i), the following equation (2) is obtained.

  The crosstalk correction kernel K (j, i) uses a slit-shaped X-ray as shown in FIGS. 15 (a) and 15 (b) corresponding to the response function of the δ function in the channel direction and column direction in advance in the channel direction and column direction. Is obtained from correction X-ray detector data collected by scanning. Alternatively, the crosstalk correction kernel can also be obtained by using X-ray detector data acquired by scanning a pencil beam X-ray passing through a pinhole as shown in FIG. 15C in the channel direction and column direction. .

  For example, when a pencil beam of X-rays hits the j-line and i-channel X-ray detector channel, the j-line and i-channel X-ray detector channel is the center and the X-ray data leaked in the vicinity is j −1 column i−1 channel, i channel, i + 1 channel, j column i−1 channel, i + 1 channel, j + 1 column i−1 channel, i channel, i + 1 channel. Thereby, a matrix Diff (i, j) of X-ray data diffused by crosstalk is determined.

  For example, assuming that j1 column i channel data is D1 (j, i), an example of a diffused X-ray data matrix Diff (i, j) as shown in Equation (3) below is shown.

  This i channel, j column is scanned for all channels in all columns to obtain Diff (i, j), and the two-dimensional deconvolution function of Diff (i, j) is 1 of the crosstalk correction kernel. An example.

  Also, for example, a slit-shaped X-ray with a width of 1 channel is scanned in the channel direction to obtain projection data for all channels in all columns, and X leaked in the vicinity from i channel to i−1 channel, i + 1 channel in each column. The amount of crosstalk diffusion in the channel direction can be obtained from the line data. Similarly, X-rays that are slit-shaped with a width of one column are scanned in the column direction to obtain projection data for all channels in all columns, and X leaked to the vicinity of j−1 column and j + 1 column in each channel. The amount of crosstalk diffusion in the column direction can be obtained from the line data. A crosstalk correction may be performed as a crosstalk correction kernel in each direction by obtaining a one-dimensional deconvolution function in each direction from the crosstalk diffusion matrix in the channel direction and the column direction. Alternatively, a crosstalk diffusion matrix such as Equation 1 may be obtained from the crosstalk diffusion matrix in the channel direction and the column direction, and a two-dimensional deconvolution function may be obtained to perform crosstalk correction as a crosstalk correction kernel.

  In the above example for obtaining the crosstalk correction kernel, the leaked X-ray data (crosstalk data) of ± 1 pixel in the vicinity is used, but the leaked X in the wide range of ± 2 pixels or more in the vicinity is used. Crosstalk correction kernels may be obtained using line data (crosstalk data) to perform crosstalk correction.

  In step S4, beam hardening correction is performed on the projection data D2 (view, j, i) subjected to crosstalk correction. In beam hardening correction S4, if the projection data subjected to crosstalk correction S3 is Din (view, j, i) and the data after beam hardening correction S4 is Dout (view, j, i), For example, the ning correction S4 is expressed in a polynomial form as shown in the following equation (4).

  At this time, since independent beam hardening correction can be performed for each j column of the detector, if the tube voltage of each data acquisition system differs depending on the imaging conditions, the X-ray energy characteristics of the detector for each column Differences can be corrected.

  In step S5, X-ray scattering correction is performed on the projection data subjected to beam hardening correction.

  When the projection data corrected by beam hardening is D3 (view, j, i) and the projection data corrected by X-ray scattering is D4 (view, j, i), the following equation (5) is obtained. .

  Fscatter (x) can be obtained as one column, and a function to be corrected can be obtained in advance from the state of the scattered X-ray change between the state without the detector X-ray collimator and the state with the detector X-ray collimator.

  As shown in Figure 12, there is a detector X-ray collimator, and X3 detector data obtained from an X-ray detector with a scintillator with a reflector is D3 collimator. (view, j, i) and measure in advance. In addition, as shown in Fig. 17, there is no detector X-ray collimator, and X-ray detector data obtained from an X-ray detector with a scintillator with a reflector is converted to D3. (view, j, i). The relationship between D3 collimator (view, j, i) and D3 (view, j, i) is obtained under various imaging conditions, and the relationship of Fscatter is obtained as shown by the following equation (6). Thus, the above X-ray scattering correction can be realized.

  In step S6, a z-filter convolution process for applying a filter in the z-direction (column direction) to the projection data that has been corrected for X-ray scattering is performed.

  In step S6, the projection data of the multi-row X-ray detector Det (ch, row) (ch = 1 to CH, row = 1 to ROW) corrected for X-ray scattering in each view angle and each data acquisition system, In the column direction, for example, a filter having a column direction filter size of 5 columns as shown in the following formula (7) is applied. However, it is set as Formula (8).

(W ₁ (ch), w ₂ (ch), w ₃ (ch), w ₄ (ch), w ₅ (ch)) (7)

  The corrected detector data Dcor (ch, row) is expressed by the following equation (9).

  The maximum value of the channel is CH, and the maximum value of the column is ROW.

  Further, when the column direction filter coefficient is changed for each channel, the slice thickness can be controlled according to the distance from the reconstruction center. In general, in slice images, the slice thickness is thicker in the periphery than in the reconstruction center, so the column direction filter coefficient is changed between the center and the periphery, and the column direction filter coefficient is changed in the column direction near the center channel. If the width of the filter coefficient is changed widely, the slice thickness can be made uniform in the peripheral part and the image reconstruction center part by changing the width of the column direction filter coefficient in the vicinity of the peripheral channel.

  In this way, by controlling the column direction filter coefficients of the central channel and the peripheral channel of the multi-row X-ray detector 24, the slice thickness can also be controlled at the central portion and the peripheral portion. When the slice thickness is slightly reduced with the row direction filter, both artifacts and noise are greatly improved. Thereby, artifact improvement and noise improvement can also be controlled. That is, it is possible to control the tomographic image reconstructed, that is, the image quality in the xy plane. In another embodiment, a thin slice thickness tomographic image can be realized by using a column direction (z direction) filter coefficient as a deconvolution filter.

  In step S7, reconstruction function superimposition processing is performed. That is, the Fourier transform is performed, the reconstruction function is multiplied, and the inverse Fourier transform is performed. In reconstruction function superimposition processing S5, if the data after z filter convolution processing is Din, the data after reconstruction function convolution processing is Dout, and the reconstruction function to be superimposed is Kernel (j), the reconstruction function convolution processing is as follows: (11)

  That is, since the reconstruction function kernel (j) can perform an independent reconstruction function superimposing process for each j column of the detector, the difference in noise characteristics and resolution characteristics for each column can be corrected.

  In step S8, three-dimensional backprojection processing is performed on the projection data DO (view, j, i) subjected to reconstruction function superimposition processing to obtain backprojection data D3 (x, y). In the present invention, helical scanning is performed, but the image to be reconstructed is reconstructed into a three-dimensional image on a plane perpendicular to the z axis and on the xy plane. The following reconstruction area P is assumed to be parallel to the xy plane. This three-dimensional backprojection process will be described later with reference to FIG.

  In step S9, post-processing such as image filter superimposition and CT value conversion is performed on the backprojection data D3 (x, y) to obtain a tomographic image.

  In post-processing image filter convolution, the data after 3D backprojection is Din (x, y, z), the data after image filter convolution is Dout (x, y, z), and the image filter is Filter (z) Then, it is shown as Expression (12).

  That is, since independent image filter convolution processing can be performed for each j column of the detector, the difference in noise characteristics and resolution characteristics for each column can be corrected.

  The obtained tomographic image is displayed on the monitor 6.

  FIG. 5 is a flowchart showing details of the three-dimensional backprojection process (step S8 in FIG. 4).

  In the present invention, helical scanning is performed, but the image to be reconstructed is reconstructed into a three-dimensional image on a plane perpendicular to the z axis and on the xy plane. The following reconstruction area P is assumed to be parallel to the xy plane.

  In step S81, attention is paid to one view in all views necessary for image reconstruction of a tomogram (that is, a view for 360 degrees or a view for "180 degrees + fan angle"). Projection data Dr corresponding to each pixel is extracted.

  As shown in FIGS. 6 (a) and 6 (b), a 512 × 512 pixel square region parallel to the xy plane is used as a reconstruction region P, and pixel rows L0 and y = 63 parallel to the x axis where y = 0 Pixel column L63, pixel column L127 with y = 127, pixel column L191 with y = 191, pixel column L255 with y = 255, pixel column L319 with y = 319, pixel column L383 with y = 383, pixel column with y = 447 When the pixel column L511 of L447, y = 511 is taken as a column, these pixel columns L0 to L511 are projected on the surface of the multi-row X-ray detector 24 in the X-ray transmission direction on lines T0 to T511 as shown in FIG. If these projection data are extracted, they become projection data Dr of the pixel columns L0 to L511.

  The X-ray transmission direction is determined by the X-ray focal point F of the X-ray tube 21 and the geometric position of each pixel and the multi-row X-ray detector 24, but the X-ray detector data D0 (view, j, i) Because the z coordinate z (view) is attached to the X-ray detector data as the table linear movement z-direction position Ztable (view), X-rays are also used for the projection data D0 (view, j, i) during acceleration / deceleration The transmission direction can be determined accurately.

  For example, a part of the line goes out of the channel direction of the multi-row X-ray detector 24, such as a line T0 in which the pixel row L0 is projected on the surface of the multi-row X-ray detector 24 in the X-ray transmission direction. In this case, the corresponding projection data Dr is set to “0”. Further, if the projection is out of the z direction, the projection data Dr is extrapolated.

  Thus, as shown in FIG. 8, projection data Dr (view, x, y) corresponding to each pixel in the reconstruction area P can be extracted.

  Returning to FIG. 5, in step S82, the projection data Dr (view, x, y) is multiplied by the cone beam reconstruction weighting coefficient to create projection data D2 (view, x, y) as shown in FIG.

  Here, the cone beam reconstruction weighting coefficient w (i, j) is as follows. In the case of fan beam image reconstruction, in general, when view = βa, a straight line connecting the focal point of the X-ray tube 21 and the pixel g (x, y) on the reconstruction area P (on the xy plane) is the center of the X-ray beam. When the angle formed with respect to the axis Bc is γ and the opposite view is view = βb, the following equation (13) is obtained.

  βb = βa + 180 ° −2γ (13)

  If the angles between the X-ray beam passing through the pixel g (x, y) on the reconstruction area P and the opposite X-ray beam and the reconstruction plane P are αa and αb, the cone beam reconstruction weighting coefficient depending on these ωa and ωb are multiplied and added to obtain backprojection pixel data D2 (0, x, y). This is shown by the following formula (14).

  D2 (0, x, y) = ωa · D2 (0, x, y) _a + ωb · D2 (0, x, y) _b (14)

  Note that the sum of the cone beam reconstruction weighting coefficients between the opposed beams is expressed by the following formula (15).

  ωa + ωb = 1 (15)

  By multiplying and adding cone beam reconstruction weighting coefficients ωa and ωb, cone angle artifacts can be reduced.

  For example, the cone beam reconstruction weighting coefficients ωa and ωb can be obtained by the following equations.

  When 1/2 of the fan beam angle is γmax, the following equations (16) to (21) are obtained.

  Here, for example, q = 1.

  For example, as an example of ga and gb, when max [] is a function that takes a larger value, the following equations (22) and (23) are obtained.

In the case of fan beam image reconstruction, each pixel on the reconstruction area P is further multiplied by a distance coefficient. For the distance coefficient, the distance from the focus of the X-ray tube 21 to the detector row j and channel i of the multi-row X-ray detector 24 corresponding to the projection data Dr is r0, and the distance from the focus of the X-ray tube 21 corresponds to the projection data Dr. (R1 / r0) ² where r1 is the distance to the pixel on the reconstruction area P.

  In the case of parallel beam image reconstruction, each pixel on the reconstruction area P may be multiplied by only the cone beam reconstruction weight coefficient w (i, j).

In step S83, as shown in FIG. 10, the projection data D2 (view, x, y) is added in correspondence with the pixels to the backprojection data D3 (x, y) that has been cleared in advance.
In step S84, steps S81 to S83 are repeated for all views necessary for image reconstruction of the tomographic image (ie, views for 360 degrees or views for “180 degrees + fan angle”), as shown in FIG. Then, back projection data D3 (x, y) is obtained.

  As shown in FIGS. 11A and 11B, the reconstruction area P may be a circular area.

  In the X-ray CT apparatus 100 of the present embodiment described above, a two-dimensional matrix structure with an X-ray light conversion element such as an X-ray detector without a detector X-ray collimator, a scintillator that converts X-rays without a reflector into light, etc. Image quality improvement and exposure reduction of conventional scan (axial scan), cine scan or helical scan of X-ray CT apparatus with area X-ray detector can be realized.

  Specifically, the X-ray CT apparatus 100 of the present embodiment includes an X-ray tube 21 that irradiates a subject with X-rays, and a plurality of detections that detect X-rays that have been irradiated from the X-ray tube 21 and transmitted through the subject. Multi-row X-ray detector 24 in which detector channels (detecting elements) are arranged in a matrix structure, DAS 25 that collects X-ray data detected by multi-row X-ray detector 24, and data collected by DAS 25 And a central processing unit 3 for reconstructing a tomographic image of the subject based on the above. Here, each channel of the multi-row X-ray detector 24 has a scintillator 24a that converts X-rays transmitted through the subject into light, and a photodiode 24b that converts light converted by the scintillator 24a into an electrical signal. The scintillator 24a is integrally formed without being divided between the plurality of detector channels. In other words, the multi-row X-ray detector 24 is formed in such a structure that, in a plurality of detector channels, light from the scintillator 24a corresponding to one channel can be detected by the photodiode 24b corresponding to the other channel. ing. The DAS 25 collects the electrical signals converted by the photodiode 24b as data. In addition, the X-ray CT apparatus 100 of the present embodiment stores a program for causing the central processing unit 3 to correct data collected by the DAS 25 using the crosstalk correction kernel K (j, i). It has a device 7. That is, the relationship between the data collected by the DAS 25 and the crosstalk data included in the data is calculated in advance in association with each channel. Alternatively, a relational expression between data collected by the DAS 25 by periodic calibration and crosstalk data included in the data may be obtained. Then, based on the relational expression about the crosstalk data obtained by these methods, a program for causing the central processing unit 3 to correct the crosstalk data from the data collected by the DAS 25 is stored in the storage device 7. I am letting. The central processing unit 3 corrects the data collected by the DAS 25 using the program stored in the storage device 7, and then reconstructs a tomographic image of the subject based on the corrected data. That is, the central processing unit 3 corrects the data collected by the DAS 25 so that crosstalk data due to crosstalk between the channels of the multi-row X-ray detector 24 is removed or restored to the original data. Thereafter, a tomographic image of the subject is reconstructed based on the corrected data. Thus, in this embodiment, the scintillator 24c of the multi-row X-ray detector 24 is integrally formed, and the reflector 24c that divides the scintillator 24a between channels and prevents crosstalk is not formed. The aperture ratio, which is the ratio of the area for detecting X-rays in the X-ray detector 24, is improved, and the utilization efficiency of X-rays can be improved. For this reason, this embodiment can improve the S / N ratio of an image. Further, the data detected by the multi-row X-ray detector 24 and collected by the DAS 25 may include crosstalk data due to crosstalk between channels because the reflector 24c is not formed. However, after correcting the collected data based on the crosstalk correction kernel K (j, i), the tomographic image of the subject is reconstructed based on the corrected data. Can be removed. Therefore, according to the present embodiment, the image quality can be improved.

  Furthermore, in the present embodiment, the multi-row X-ray detector 24 does not have a detector X-ray collimator 31 that shields scattered X-rays in the channel direction irradiated on each channel on the detection surface. For this reason, in this embodiment, irradiation of scattered X-rays in the channel direction is not limited, and each channel is formed to detect scattered X-rays. The storage device 7 stores a program for causing the central processing unit 3 to correct the data collected by the DAS 25 using the scattered X-ray correction function Fscatter (x). That is, in the present embodiment, a relational expression between data collected by the DAS 25 and scattered X-ray data included in the data due to scattered X-rays is calculated in advance in association with each channel, and A program that causes the central processing unit 3 to correct the scattered X-ray data from the data collected by the DAS 25 during the main scan based on the relational expression for the scattered X-ray data is stored in the storage device 7. Then, the central processing unit 3 performs scattered X-ray correction using a program stored in the storage device 7, and reconstructs a tomographic image of the subject based on the scattered X-ray corrected data. . That is, the central processing unit 3 corrects the data collected by the DAS 25 so as to remove the scattered X-ray data, and then reconstructs a tomographic image of the subject based on the corrected data. Thus, in this embodiment, since the detector X-ray collimator 31 that shields scattered X-rays is not formed on the detector surface dp of the multi-row X-ray detector 24, the multi-row X-ray detector 24 It is possible to improve the utilization efficiency of X-rays by increasing the ratio of the area for detecting X-rays. For this reason, this embodiment can improve the S / N ratio of an image. Further, the data detected by the multi-row X-ray detector 24 and collected by the DAS 25 may include scattered X-ray data, but the data is corrected based on the scattered X-ray correction function, Since the tomographic image of the subject is reconstructed based on the corrected data, the image can be reconstructed by removing the influence of scattered X-rays. Therefore, according to the present embodiment, the image quality of the tomographic image can be improved.

  In implementing the present invention, the present invention is not limited to the above-described embodiment, and various modifications can be employed.

  For example, the image reconstruction method may be a conventionally known Feldkamp method or a three-dimensional image reconstruction method also referred to as cone beam reconstruction. Furthermore, other three-dimensional image reconstruction methods may be used. Alternatively, the same effect can be obtained by two-dimensional image reconstruction that is not conventional cone beam reconstruction.

  In the above embodiment, column direction (z direction) filters having different coefficients for each column are superimposed, so that differences in image quality due to differences in X-ray cone angle, etc., particularly in conventional scan (axial scan). Adjustment is made to achieve uniform slice thickness, artifact, and noise image quality in each column, but various filter coefficients can be considered, but all can produce the same effect.

  The above-described embodiment can be used in an X-ray CT-PET apparatus, an X-ray CT-SPECT apparatus, or the like combined with an industrial X-ray CT apparatus or other apparatuses in addition to a medical X-ray CT apparatus.

  The above embodiment shows the case of a multi-row X-ray detector, but the case of a single-row X-ray detector is also effective.

  In the above embodiment, one row of crosstalk correction and one row of X-ray scattering correction are shown, but the same effect can be obtained by other crosstalk correction and X-ray scattering correction. Further, either one of crosstalk correction and X-ray scattering correction may be used. For example, when the reflector 24c is formed in the multi-row X-ray detector 24 as shown in FIG. Only scattering correction may be performed.

  In the above embodiment, one column for obtaining the crosstalk correction kernel and one column for obtaining the X-ray scattering correction function are shown, but the same effect can be obtained by other methods.

  In the above embodiment, image reconstruction is performed in the process flow as shown in FIG. 3, but the same effect can be obtained in a similar process flow. Further, the same effect can be obtained even if the positions of the crosstalk correction and X-ray scattering correction processing are not as shown in FIG.

FIG. 1 is a block diagram showing an X-ray CT apparatus according to an embodiment of the present invention. FIG. 2 is an explanatory diagram showing rotation of the X-ray generator (X-ray tube) and the multi-row X-ray detector. FIG. 3 is a flowchart showing a schematic operation of the X-ray CT apparatus according to one embodiment of the present invention. FIG. 4 is a flowchart showing details of the preprocessing. FIG. 5 is a flowchart showing details of the 3D image reconstruction process. FIG. 6 is a conceptual diagram showing a state in which lines on the reconstruction area are projected in the X-ray transmission direction. FIG. 7 is a conceptual diagram showing lines projected on the detector surface. FIG. 8 is a conceptual diagram showing a state in which the projection data Dr (view, x, y) is projected onto the reconstruction area. FIG. 9 is a conceptual diagram showing backprojected pixel data D2 of each pixel on the reconstruction area. FIG. 10 is an explanatory diagram illustrating a state in which the back projection pixel data D2 is added to all the views in correspondence with the pixels to obtain the back projection data D3. FIG. 11 is a conceptual diagram showing a state in which a line on a circular reconstruction area is projected in the X-ray transmission direction. FIG. 12 is a diagram showing an X-ray detector having a conventional detector X-ray collimator and a scintillator with a reflector. Here, (a) is a side view with the line direction as the line of sight, and (b) is a cross-sectional view showing the X0 portion of (a). FIG. 13 is a diagram showing X-ray geometric efficiency. FIG. 14 is a diagram illustrating a detector X-ray collimator width and a reflector width. In FIG. 15, (a) is a diagram showing scanning with a slit-shaped X-ray in the column direction. (B) is a figure which shows scanning a slit-shaped X-ray in a channel direction. (C) is a diagram showing scanning in the column direction and the channel direction with a pencil beam X-ray passing through the pinhole. FIG. 16 is a diagram showing an X-ray detector having a scintillator with a reflector without a detector X-ray collimator. Here, (a) is a side view with the line direction as the line of sight, and (b) is a cross-sectional view showing the X1 portion of (a). FIG. 17 is a view showing an X-ray detector having a scintillator without a reflector without a detector X-ray collimator. Here, (a) is a side view with the line direction as the line of sight, and (b) is a cross-sectional view showing the X2 portion of (a). FIG. 18 is a diagram showing an X-ray detector without a flat panel type detector X-ray collimator and having a scintillator without a reflector. Here, (a) is a perspective view, and (b) is a cross-sectional view showing the X3 portion of (a). FIG. 19 is a diagram showing an X-ray detector having a scintillator without a reflector without a detector X-ray collimator combining a plurality of flat panels. Here, (a) is a perspective view, and (b) is a cross-sectional view showing the X4 portion of (a).

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 ... Operation console 2 ... Input device 3 ... Central processing unit 5 ... Data collection buffer 6 ... Monitor 7 ... Storage device 10 ... Imaging table 12 ... Cradle 15 ... Rotating part 20 ... Scanning gantry 21 ... X-ray tube 22 ... X-ray controller 23 ... Collimator 24 ... Multi-row X-ray detector 24a ... Scintillator 24b ... Photodiode 25 ... DAS (data collection device)
26 ... Rotating part controller 27 ... Gantry tilt controller 29 ... Control controller 30 ... Slip ring dp ... Detector surface P ... Reconstruction area pp ... Projection surface
IC: Center of rotation (ISO)

Claims (5)

An X-ray generator and a matrix-structured two-dimensional X-ray area detector that detects X-rays relative to the X-ray generator are in between while rotating around the center of rotation. X-ray data collection means for collecting X-ray projection data transmitted through the subject;
Image reconstruction means for reconstructing an image of the X-ray projection data collected from the X-ray data collection means;
Image display means for displaying the tomographic image reconstructed;
In an X-ray CT apparatus having an imaging condition setting means for setting imaging conditions for the tomographic imaging,
The two-dimensional X-ray area detector is a structure that captures scattered X-rays,
The image reconstruction means performs scattered X-ray correction.
X-ray CT system characterized by that. An X-ray generator and a matrix-structured two-dimensional X-ray area detector that detects X-rays relative to the X-ray generator are in between while rotating around the center of rotation. X-ray data collection means for collecting X-ray projection data transmitted through the subject;
Image reconstruction means for reconstructing an image of the X-ray projection data collected from the X-ray data collection means;
Image display means for displaying the tomographic image reconstructed;
In an X-ray CT apparatus having an imaging condition setting means for setting imaging conditions for the tomographic imaging,
The two-dimensional X-ray area detector
Means for converting X-rays that occupy almost 100% of the X-ray irradiated surface into light;
Means for converting light into an electrical signal,
The X-ray CT apparatus, wherein the image reconstruction means corrects crosstalk between channels and columns. An X-ray generator and a matrix-structured two-dimensional X-ray area detector that detects X-rays relative to the X-ray generator are in between while rotating around the center of rotation. X-ray data collection means for collecting X-ray projection data transmitted through the subject;
Image reconstruction means for reconstructing an image of the X-ray projection data collected from the X-ray data collection means;
Image display means for displaying the tomographic image reconstructed;
In an X-ray CT apparatus having imaging condition setting means for setting imaging conditions for the tomographic imaging,
The two-dimensional X-ray area detector
Means for converting X-rays that occupy almost 100% of the X-ray irradiated surface into light;
X-ray data collection means having means for converting light into an electrical signal, and a structure for capturing scattered X-rays,
The X-ray CT apparatus, wherein the image reconstruction means performs scattered X-ray correction and interchannel / column crosstalk correction. In the X-ray CT apparatus according to any one of claims 1 to 3,
The X-ray CT apparatus, wherein the two-dimensional X-ray area detector is a multi-row X-ray detector. In the X-ray CT apparatus according to any one of claims 1 to 3,
2. The X-ray CT apparatus, wherein the two-dimensional X-ray area detector is a planar X-ray detector typified by a flat panel X-ray detector.

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