JP4884765B2 - X-ray CT system - Google Patents

Description

  The present invention relates to an X-ray CT (Computed Tomography) apparatus used for medical or industrial purposes, and relates to a conventional scan (conventional scan), a cine scan, a helical scan, or a variable scan. The present invention relates to exposure reduction and image quality improvement in a pitch helical scan and a helical shuttle scan.

  Conventionally, in an X-ray CT apparatus using a two-dimensional X-ray area detector having a matrix structure represented by an X-ray CT apparatus or a flat panel (flat panel), as shown in FIG. Each X-ray detector channel of the X-ray area detector is read one by one when collecting data. In this case, the crosstalk rate indicating the crosstalk ratio of each X-ray detector channel is as shown in FIG. In particular, the problem of crosstalk is that the amount of crosstalk increases when the width of the reflector existing between the X-ray detector channels is thin or when there is no reflector. In this case, the resolution of the XY plane on the tomographic image is low. It was a problem from the viewpoint of image quality that deteriorated.

However, in an X-ray CT apparatus using a two-dimensional X-ray area detector typified by an X-ray CT apparatus or a flat panel, the problem of unnecessary X-ray exposure is increasing.
As the width of the X-ray detector expands in the z-direction, the area of the X-ray detector increases, the number of channels of the data acquisition device that performs A / D conversion increases, and each X-ray detector channel also becomes finer. . In this case, if there is a reflector present between each X-ray detector channel, the X-ray capture efficiency is lowered, so that the exposure of the subject tends to increase. In order to avoid this, the X-ray detector is proceeding in the direction of reducing the width of the reflector or eliminating the reflector. The problem in this case is the problem of data crosstalk between the X-ray detector channels.

One cause of crosstalk occurring between the respective X-ray detector channels is that, as shown in FIG. 19, from the scintillator of each X-ray detector channel to the adjacent photodiode (photo-diode). Light may leak in and enter.
US Pat. No. 6,115,448

  Accordingly, an object of the present invention is to provide a conventional scan (axial scan), cine scan, or helical scan of an X-ray CT apparatus having a two-dimensional X-ray area detector having a matrix structure represented by a flat panel X-ray detector. Alternatively, X-ray CT that realizes a data acquisition method or preprocessing method or image reconstruction method in which crosstalk between the X-ray detectors is reduced and S / N is improved during variable-pitch helical scan data acquisition. To provide an apparatus.

  In the present invention, in order not to reduce the resolution of the X-ray detector, data is collected at intervals of a size dc in the channel direction or a size dr in the column direction of one X-ray detector. Further, in order to improve the crosstalk rate, detection signals are added together in the channel direction and the column direction as shown in FIG.

  The effect of crosstalk on the resolution varies depending on each image reconstruction function. In addition, when crosstalk correction is performed, characteristics vary depending on the image reconstruction function even when the correction coefficient is changed. The crosstalk characteristics also change when an image is reconstructed by bundling a plurality of X-ray detector channels in the z direction, which is the orthogonal direction.

  In consideration of the above, by preparing not only one crosstalk correction coefficient but also various cases with different shooting conditions, depending on the conditions of data collection and image reconstruction, using the optimum coefficient, By providing an X-ray detector or an X-ray CT apparatus capable of obtaining a tomographic image having optimum resolution and noise characteristics (S / N) according to the application by performing crosstalk correction Solve the problem.

  In a first aspect, the present invention relates to an X-ray generator that generates X-rays and a two-dimensional X-ray detector that detects the X-rays with an X-ray detector channel of a two-dimensional X-ray area detector facing the direction of generation. An X-ray area detector and the two-dimensional X-ray area detector are arranged opposite to each other around a rotation center where a subject is arranged, and the X-ray area detector is rotated around the rotation center while generating the X-ray. X-ray data collection means for collecting X-ray projection data transmitted through the specimen, image reconstruction means for reconstructing the X-ray projection data and reconstructing a tomographic image, and image display for displaying the tomographic image Means and an imaging condition setting means for setting imaging conditions used in the acquisition, wherein the X-ray data acquisition means is in the longitudinal direction of the two-dimensional X-ray area detector The X-ray detector channel in the channel direction When the magnitude to be performed is dc, the integer n is a number equal to or greater than 2, and the X-ray detection signals of the X-ray detectors are added n by n in the channel direction, and the added X-ray detector is added. An X-ray CT apparatus is provided, wherein data is collected as one X-ray projection data with the interval of dc in the channel direction.

In the X-ray CT apparatus according to the first aspect, since the X-ray detection signals are bundled n times in the channel direction, the S / N is improved by n ^1/2 times, and the data collection interval is dc in the channel direction. Is the spatial resolution in the channel direction of n · dc / 2.

  In a second aspect, the present invention relates to an X-ray generator that generates X-rays, and a two-dimensional X-ray detector that detects the X-rays with an X-ray detector channel of a two-dimensional X-ray area detector facing the direction of generation. An X-ray area detector and the two-dimensional X-ray area detector are arranged opposite to each other around a rotation center where a subject is arranged, and the X-ray area detector is rotated around the rotation center while generating the X-ray. X-ray data collection means for collecting X-ray projection data transmitted through the specimen, image reconstruction means for reconstructing the X-ray projection data and reconstructing a tomographic image, and image display for displaying the tomographic image And an imaging condition setting means for setting an imaging condition used in the acquisition, wherein the X-ray data acquisition means is substantially in the longitudinal direction of the two-dimensional X-ray area detector. X-ray detector channel in orthogonal column direction When the magnitude is dr, the integer m is a number of 2 or more, and the X-ray detection signals of the X-ray detectors are added m by m in the column direction, and the added X-ray detector is added. An X-ray CT apparatus is provided, wherein one piece of data is collected as the X-ray projection data with an interval of dr in the column direction.

In the X-ray CT apparatus according to the second aspect, since the X-ray detection information is bundled m columns in the column direction, the S / N is improved by m ^1/2 times, and the data collection interval is dr in the column direction. Is the spatial resolution in the column direction of m · dr / 2.

  In a third aspect, the present invention provides an X-ray generator that generates X-rays, and a two-dimensional X-ray detector that detects the X-rays using an X-ray detector channel of a two-dimensional X-ray area detector that faces the direction of the generation. An X-ray area detector and the two-dimensional X-ray area detector are arranged opposite to each other around a rotation center where a subject is arranged, and the X-ray area detector is rotated around the rotation center while generating the X-ray. X-ray data collection means for collecting X-ray projection data transmitted through the specimen, image reconstruction means for reconstructing the X-ray projection data and reconstructing a tomographic image, and image display for displaying the tomographic image Means and an imaging condition setting means for setting imaging conditions used in the acquisition, wherein the X-ray data acquisition means is in the longitudinal direction of the two-dimensional X-ray area detector The X-ray detector channel in the channel direction When the magnitude of the X-ray detector channel in the column direction substantially orthogonal to the longitudinal direction of the two-dimensional X-ray area detector is dr, the integers n and m are 2 or more. As a number, X-ray detection signals of the X-ray detectors are added n by n in the channel direction and m in the column direction, and the added X-ray detector data is added to one X-ray detector data. There is provided an X-ray CT apparatus characterized by collecting line projection data at intervals of the dc and the dr in the channel direction and the column direction.

In the X-ray CT apparatus according to the third aspect, if the n-channel X-ray detection information is bundled in the channel direction, the S / N is improved by n ^1/2 times, and the x-ray detection information is bundled in m columns in the column direction. lever m ^1/2 times only the better the S / n, when the data collection interval in the channel direction of the dc, there is spatial resolution in the channel direction of the n · dc / 2, dr data collection interval in the column direction In this case, the spatial resolution in the column direction is m · dr / 2.

  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 or a matrix structure 2. An X-ray CT apparatus comprising a dimensional X-ray area detector is provided.

  Since the X-ray CT apparatus according to the fourth aspect may be a two-dimensional X-ray area detector extending in a two-dimensional direction, it is a multi-row X-ray detector or a matrix structure two-dimensional X-ray area detector. Also good.

  In a fifth aspect, the present invention provides the X-ray CT apparatus according to any one of the first to fourth aspects, wherein the X-ray data collection means performs at least the number of the n and the m that perform the addition. An X-ray CT apparatus is provided in which one number is two.

In the X-ray CT apparatus according to the fifth aspect, at least one of n or m shown in the first to fourth aspects is 2, so that two channels are bundled in the channel direction by 2 ^1/2. If the S / N is improved by a factor of 2 and two rows are bundled in the column direction, the S / N is improved by a factor of 2 ^1/2 and the data collection interval is dc in the channel direction. Since the interval is dr in the column direction, both the spatial resolution in the column direction is set to the same resolution as when data is not bundled.

  In a sixth aspect, the present invention provides the X-ray CT apparatus according to any one of the first to fifth aspects, wherein the X-ray data collection means sets the interval for performing collection in the channel direction to the dc An X-ray CT apparatus characterized in that the value is an integer multiple of.

In the X-ray CT apparatus according to the sixth aspect, since the data collection interval in the channel direction is an integral multiple of dc, the data collection amount is reduced by that amount and the data processing speed is increased.
In a seventh aspect, the present invention provides the X-ray CT apparatus according to any one of the second to fifth aspects, wherein the X-ray data collection means sets the interval at which the column direction collection is performed, Provided is an X-ray CT apparatus characterized by having a value of dr.

In the X-ray CT apparatus according to the seventh aspect, since the data collection interval in the column direction is an integral multiple of dr, the data collection amount is reduced by that amount and the data processing speed is increased.
In an eighth aspect, the present invention provides the X-ray CT apparatus according to any one of the first to seventh aspects, wherein the image reconstruction means includes the channel direction of the two-dimensional X-ray area detector or the There is provided an X-ray CT apparatus characterized in that crosstalk correction is turned on / off based on the collection interval in the column direction or the number of additions.

  In the X-ray CT apparatus according to the eighth aspect, the image quality is controlled by controlling on / off of the crosstalk correction based on the data collection interval in the channel direction, the column direction, and the number of additions in the two-dimensional X-ray area detector. To optimize.

  In a ninth aspect, the present invention provides the X-ray CT apparatus according to any one of the first to eighth aspects, wherein the image reconstruction unit includes the channel direction of the two-dimensional X-ray area detector or the An X-ray CT apparatus is provided, wherein a coefficient of crosstalk correction is changed based on the collection interval in the column direction or the number of additions.

  In the X-ray CT apparatus according to the ninth aspect, the coefficient of crosstalk correction is based on the channel direction in the multi-row X-ray detector or the two-dimensional X-ray area detector, the data collection interval in the column direction, and the number of additions. Is used to control the effect of crosstalk correction to optimize image quality.

  In a tenth aspect, the present invention provides the X-ray CT apparatus according to any one of the first to ninth aspects, wherein the image reconstruction means includes the channel direction of the two-dimensional X-ray area detector or the An X-ray CT apparatus is provided, wherein the image reconstruction function set by the imaging condition setting means is changed based on the collection interval in the column direction or the number of additions.

  In the X-ray CT apparatus according to the tenth aspect, or based on the channel direction, the data collection interval in the column direction, and the number of additions in the two-dimensional X-ray area detector, the image reconstruction designated by the imaging condition setting means Fine-tune functions to optimize image quality.

  In an eleventh aspect, the present invention provides the X-ray CT apparatus according to any one of the first to tenth aspects, wherein the image reconstruction means includes the channel direction of the two-dimensional X-ray area detector or the The X-ray CT apparatus is characterized in that the z-direction filter coefficient is changed based on the collection interval in the column direction or the number of additions.

  In the X-ray CT apparatus according to the eleventh aspect, in the column direction filter processing based on the channel direction in the multi-row X-ray detector or the two-dimensional X-ray area detector, the data collection interval in the column direction, and the number of additions. By controlling the z-direction filter coefficients, noise artifacts can be controlled and image quality is optimized.

  In a twelfth aspect, the present invention provides the X-ray CT apparatus according to any one of the first to eleventh aspects, in which the X-ray data collection unit is configured to perform the collection interval or the based on the imaging conditions. An X-ray CT apparatus is provided, wherein the number of additions is calculated, and the image reconstruction unit determines crosstalk correction, the z-direction filter, and an image reconstruction function based on the imaging conditions. .

  In the X-ray CT apparatus according to the twelfth aspect, the data collection interval in the channel direction and the column direction in the two-dimensional X-ray area detector and the number of additions are determined based on the imaging conditions determined by the imaging condition setting means. Based on the data collection method, crosstalk correction, z-direction filter, and fine adjustment of the image reconstruction function are optimized to obtain an image with optimum image quality.

  In a thirteenth aspect, the present invention provides the X-ray CT apparatus according to the twelfth aspect, in which the X-ray data collection means includes a noise index value of the imaging condition, and based on the noise index value, X-ray characterized in that an acquisition interval or the number of additions is calculated, and the image reconstruction means determines a crosstalk correction, the z-direction filter, and an image reconstruction function based on the imaging conditions. A CT apparatus is provided.

  In the X-ray CT apparatus according to the thirteenth aspect, the image quality is optimized based on the noise index value determined by the imaging condition setting means.

  According to the X-ray CT apparatus of the present invention, or a conventional scan (axial scan) or cine scan of an X-ray CT apparatus having a two-dimensional X-ray area detector having a matrix structure represented by a flat panel X-ray detector. Or, when acquiring data of helical scan or variable pitch helical scan, the crosstalk between the X-ray detectors is reduced, and a data acquisition method or preprocessing method or image reconstruction method that improves S / N is realized. There is an effect that an X-ray CT apparatus can be realized.

Hereinafter, the present invention will be described in more detail with reference to embodiments shown in the drawings. Note that the present invention is not limited thereby.
FIG. 1 is a 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. Note that the xyz coordinates shown in FIG. 1 represent the same coordinates as the xyz coordinates in the following drawings, and indicate the positional relationship between the drawings.

  The operation console 1 includes an input device 2 that is an imaging condition setting unit that receives an input from an operator, a central processing unit 3 that is an image reconstruction unit that performs preprocessing, image reconstruction processing, post-processing, and the like, and a scanning gantry. A data collection buffer (buffer) 5 that collects the X-ray detector data collected in 20 and an image display means for displaying a tomographic image reconstructed from projection data obtained by preprocessing the X-ray detector data. A monitor 6 and a storage device 7 for storing a program, X-ray detector data, projection data, and an X-ray tomographic image are provided.

  The photographing condition is input from the input device 2 as the photographing condition setting means and stored in the storage device 7. FIG. 14 shows an example of the shooting condition input screen. 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 moved linearly by a motor built in the imaging table 10.

  The scanning gantry 20 includes an X-ray tube 21 which is an X-ray generator, an X-ray controller 22, a collimator 23, an X-ray beam forming filter 28, and a two-dimensional X-ray area detection. A multi-row X-ray detector 24, a DAS (Data Acquisition System) 25, an X-ray tube 21 rotating around the body axis of the subject, a control signal, etc. Is provided with a control controller 29 for communicating with the operation console 1 and the imaging table 10, and a slip ring 30.

  Here, the X-ray tube 21, the collimator 23, the X-ray beam forming filter 28, the X-ray controller 22, the DAS 25, the rotating unit controller 26, and the control controller 29 constitute X-ray data collection means.

  The X-ray beam forming filter 28 has the thinnest filter thickness in the X-ray direction toward the center of rotation, which is the imaging center, and the filter thickness increases toward the periphery, so that X-rays can be absorbed more. X-ray filter. For this reason, exposure of the body surface of the subject having a cross-sectional shape close to a circle or an ellipse can be reduced. Further, the scanning gantry tilt controller 27 can tilt the scanning gantry 20 forward and backward in the z direction by about ± 30 degrees.

  The X-ray tube 21 and the multi-row X-ray detector 24 that is a two-dimensional X-ray area detector rotate around the rotation center IC. When the vertical direction is the y direction, the horizontal direction is the x direction, and the table and cradle traveling direction perpendicular to these are the z direction, the rotation plane of the X-ray tube 21 and the multi-row X-ray detector 24 is the xy plane. is there. The moving direction of the cradle 12 is the z direction.

  2 and 3 are explanatory views of the geometric arrangement of the X-ray tube 21 and the multi-row X-ray detector 24, which is a two-dimensional X-ray area detector, as viewed from the xy plane or the yz plane. The X-ray tube 21 generates an X-ray beam called a cone beam CB. When the central axis direction of the cone beam CB is parallel to the y direction, the view angle is set to 0 degree. The multi-row X-ray detector 24 has, for example, 256 X-ray detector rows in the z direction. Each X-ray detector array has, for example, 1024 X-ray detector channels in the channel direction.

  In FIG. 2, the X-ray beam emitted from the X-ray focal point of the X-ray tube 21 is caused by the X-ray beam forming filter 28 to cause more X-rays at the center of the reconstruction area P and more at the periphery of the reconstruction area P. After the X-ray dose is spatially controlled so that a small amount of X-rays are irradiated, the X-rays are absorbed by the subject existing inside the reconstruction area P, and the transmitted X-rays are converted into the multi-row X-ray detector 24. Is collected as X-ray detector data.

  In FIG. 3, the X-ray beam emitted from the X-ray focal point of the X-ray tube 21 is controlled in the slice thickness direction of the tomographic image by the X-ray collimator 23, that is, the X-ray beam width is reduced at the rotation center axis IC. The X-ray is absorbed by the subject existing near the rotation center axis IC, and the transmitted X-ray is collected as X-ray detector data by the multi-row X-ray detector 24.

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

In the following, examples of crosstalk reduction processing and S / N improvement processing will be shown for three embodiments.
Embodiment 1: Example in the case of the multi-row X-ray detector 24 Embodiment 2: Example in the case of the two-dimensional X-ray area detector 34 Embodiment 3: Control of the crosstalk reduction processing depending on the imaging conditions Example (Embodiment 1)
FIG. 4 is a flowchart showing an outline of the operation of the X-ray CT apparatus of the first embodiment.

  In step P1, the subject is placed on the cradle 12 and aligned. The subject placed on the cradle 12 aligns the slice light center position of the scanning gantry 20 with the reference point of each part.

  In step P2, scout image collection is performed. Scout images are usually taken at 0 ° and 90 °, but depending on the part, for example, the head may be a 90 ° scout image only. Details of scout image shooting will be described later.

  In step P3, shooting conditions are set. The normal photographing conditions are set while displaying the position and size of the tomographic image to be photographed on the scout image. In this case, the region of interest is set on the scout image together with the display of the X-ray dose information as a whole for the helical scan, the variable pitch helical scan, the conventional scan (axial scan), or the cine scan. Displays radiation dose information. In the cine scan, when the number of rotations or time is entered, X-ray dose information for the input number of rotations or the input time in the region of interest is displayed.

  In step P4, tomographic imaging is performed. Details of tomographic imaging will be described later. FIG. 5 is a flowchart showing an outline of operations of tomographic imaging and scout imaging of the X-ray CT apparatus 100 of the present invention.

  In step S1, in the case of helical scanning, the X-ray detection is performed while rotating the X-ray tube 21 and the multi-row X-ray detector 24 around the subject and moving the cradle 12 on the imaging table 10 linearly. The device data is collected, and the X-ray detector data D0 (view, j, i) represented by the view angle view, the detector row number j, and the channel number i is moved to the table linear movement z-direction position Ztable ( view) is added to collect X-ray detector data. In the case of variable-pitch helical scan, data collection is performed not only in the range of a constant speed in the helical scan but also during acceleration and deceleration.

  In the case of conventional scan (axial scan) or cine scan, X-ray detector data is obtained by rotating the data acquisition system one or more times while the cradle 12 on the imaging table 10 is fixed at a certain z-direction position. Data collection. If necessary, after moving to the next position in the z direction, the data acquisition system is rotated once or a plurality of times to collect data of X-ray detector data.

  In scout imaging, the X-ray tube 21 and the multi-row X-ray detector 24 are fixed, and the X-ray detector data is collected while the cradle 12 on the imaging table 10 is linearly moved. .

  In step S2, the X-ray detector data D0 (view, j, i) is preprocessed and converted into projection data. As shown in FIG. 6, the preprocessing includes step S21 offset correction, step S22 logarithmic conversion, step S23 X-ray dose correction, and step S24 sensitivity correction. In the case of scout imaging, if preprocessed X-ray detector data is displayed by matching the pixel size in the channel direction and the pixel size (size) in the z direction, which is the cradle linear movement direction, to the display pixel size of the monitor 6. Completed as a scout statue.

  Assuming that the projection data for which the pre-processing in step S2 has been completed is D01 (view, j, i), the crosstalk reduction processing in step S0 is performed on this data using the following formula, and crosstalk reduction and S / N are performed. Make improvements.

Or

Or

  In Equation (1), data is bundled in two channels by two channels in the channel direction and column direction. The data interval is dc in the channel direction and dr in the column direction. However, the width of one channel of the X-ray detector is dc in the channel direction and dr in the column direction. FIG. 17A and FIG. 17B show diagrams when this processing is performed.

In the formula (2), data is bundled by two channels in the channel direction. The data interval is dc in the channel direction and dr in the column direction.
In Equation (3), two columns of data are bundled in the column direction. The data interval is dc in the channel direction and dr in the column direction.

  Here, it will be shown using FIG. 16 and FIG. 18 that the crosstalk rate decreases in the case of Equation (1). First, FIG. 16 shows a conventional crosstalk rate. The i-channel and j-row X-ray detector data is d (i, j), and the crosstalk amount leaking to the adjacent X-ray detector channel in the channel and column directions is Δd.

  The crosstalk rate in this case is as shown in the following formula (4).

  In addition, when the crosstalk reduction processing shown in Equation (1) is performed, since 2 channels and 2 columns are bundled in the channel direction and the column direction, the signal 4 · d (i, j) for a total of 4 channels is obtained. . However, assuming that the X-ray detector channels of d (i + 1, j), d (i, j + 1), and d (i + 1, j + 1) also have a signal amount equivalent to d (i, j), the expression (1) shows. When the crosstalk reduction processing is performed, an approximation of the following formula (5) is assumed as a signal amount after being bundled.

  Further, the crosstalk amount leaking from the total of four channels is 4 · 2 · Δd = 8 · Δd because each side is doubled. The crosstalk rate in this case is as shown in the following formula (6).

That is, the crosstalk rate in the case of the formula (6) is halved compared to the crosstalk rate in the case of the formula (4). In this way, the crosstalk rate is improved.
Similarly, for S / N, the signal of one channel before the crosstalk reduction processing is d (i, j), and the signal of the bundled one channel after the crosstalk reduction processing is 4 · d (i, j). The noise of the signal is improved to 1/4 ^1/2 = ^1/2 times. Regarding the spatial resolution, the data collection interval does not change at the interval of dc in the channel direction and at the interval of dr in the column direction, similarly before and after the crosstalk reduction processing. That is, even after the crosstalk reduction processing, the data collection aperture 2 · dc in the channel direction and the data collection aperture 2 · dr in the column direction make the data collection interval dc in the channel direction and dr in the column direction, which is ideal in the sampling theorem. Data collection. For this reason, the spatial resolution does not deteriorate.

  In step S3, beam hardening correction is performed on the preprocessed projection data D1 (view, j, i). In the beam hardening correction S3, the projection data subjected to the sensitivity correction S24 in the pre-processing S2 is D1 (view, j, i), and the data after the beam hardening correction S3 is D11 (view, j, i). The beam hardening correction S3 is expressed, for example, in a polynomial form as shown in the following formula (7).

  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 varies depending on the imaging conditions, the X-ray energy characteristic of the detector for each column Differences can be corrected.

  In step S4, a z-filter convolution process for applying a filter in the z-direction (column direction) is performed on the projection data D11 (view, j, i) subjected to beam hardening correction. That is, the projection data D11 (view, j, i) (where i = 1 to CH, j) of the multi-row X-ray detector 24 subjected to beam hardening correction after pre-processing in each view angle and each data acquisition system. = 1 to ROW), a filter having a column direction filter size of 5 columns is applied in the column direction, for example, as shown in the following equations (8) and (9).

However,

And
The corrected detector data D12 (view, j, i) is expressed by the following equation (10).

It becomes. If the maximum channel value is CH and the maximum column value is ROW, the following equations (11) and (12) are obtained.

  Further, when the column direction filter coefficient is changed for each channel, the slice thickness can be controlled in accordance with the distance from the image 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 in the central portion and the peripheral portion. When the slice thickness is slightly reduced by the column direction filter, both artifact and noise are greatly improved. Thereby, artifact improvement and noise improvement can also be controlled. That is, the tomographic image reconstructed in three dimensions, that is, the image quality in the xy plane can be controlled.

  As another embodiment, a thin slice thickness tomographic image can be realized by using a column direction (z direction) filter coefficient as a deconvolution filter. Further, the fan beam X-ray projection data is converted into parallel beam X-ray projection data as necessary.

  In step S5, reconstruction function superimposition processing is performed. That is, Fourier transform, multiplication by an image reconstruction function, and inverse Fourier transform are performed. In the reconstruction function superimposing process S5, if the data after the z filter convolution process is D12, the data after the reconstruction function convolution process is D13, the image reconstruction function to be superimposed is Kernel (j), and the convolution calculation is *, The constituent function superimposition processing is expressed as the following formula (13).

In other words, the image reconstruction function kernel (j) can perform an independent reconstruction function convolution process for each j column of the detector, so that the difference in noise characteristic and resolution characteristic for each column can be corrected.
In step S6, three-dimensional backprojection processing is performed on the projection data D13 (view, j, i) subjected to reconstruction function superimposition processing to obtain backprojection data D3 (x, y, z). The image reconstructed is a three-dimensional image reconstructed 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 S7, post-processing such as image filter superposition and CT value conversion is performed on the backprojection data D3 (x, y, z) to obtain a tomographic image D31 (x, y). In post-processing image filter convolution processing, the tomographic image after three-dimensional backprojection is D31 (x, y, z), the data after image filter convolution is D32 (x, y, z), and the image filter is Filter (z ), The following formula (14) is obtained.

That is, since an independent image filter convolution process can be performed for each j column of the detector, a difference in noise characteristics and resolution characteristics for each column can be corrected. The obtained tomographic image is displayed on the monitor 6.

FIG. 7 is a flowchart showing details of the three-dimensional backprojection process shown in step S6 of FIG.
In the first embodiment, 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 S61, attention is paid to one view among all the views necessary for image reconstruction of the tomographic image (that is, a view for 360 degrees or a view for 180 degrees and a fan angle). Projection data Dr corresponding to each pixel is extracted. As shown in FIGS. 8A and 8B, a square region of 512 × 512 pixels parallel to the xy plane is set as a reconstruction region P, and pixel rows L0, y parallel to the x axis where y = 0. = 63 pixel column L63, y = 127 pixel column L127, y = 191 pixel column L191, y = 255 pixel column L255, y = 319 pixel column L319, y = 383 pixel column L383, y = 447 9 is a line T0 as shown in FIG. 9 in which these pixel columns L0 to L511 are projected onto the surface of the multi-row X-ray detector 24 in the X-ray transmission direction. When the projection data on .about.T511 are extracted, they become projection data Dr (view, x, y) of the pixel columns L0 to L511. However, x and y correspond to each pixel (x, y) of the tomographic image.

  The X-ray transmission direction is determined by the X-ray focal point of the X-ray tube 21 and the geometric position of each pixel and the multi-row X-ray detector 24, but z of the X-ray detector data D0 (view, j, i). Since the coordinate z (view) is attached to the X-ray detector data as the table linear movement z-direction position Ztable (view), the X-ray detector data D0 (view, j, i) during acceleration / deceleration is also known. In the data acquisition geometric system of the X-ray focus and multi-row X-ray detector 24, the X-ray transmission direction can be accurately obtained.

  Note that, for example, a part of the line is outside 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. If it appears, the corresponding projection data Dr (view, x, y) is set to “0”. Further, if the projection is out of the z direction, the projection data Dr (view, x, y) is extrapolated. In this manner, as shown in FIG. 10, projection data Dr (view, x, y) corresponding to each pixel in the reconstruction area P can be extracted.

  Returning to FIG. 7, in step S62, 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, generally, 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) with view = βa 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 formula (15) is obtained.

If the angles formed by the X-ray beam passing through the pixel g (x, y) on the reconstruction area P and the opposite X-ray beam to the reconstruction plane P are αa and αb, the cone beam reconstruction weighting coefficient depending on these angles ωa and ωb are multiplied and added to obtain backprojection pixel data D2 (0, x, y). In this case, Equation (16) is obtained.

  Note that the sum of the cone beam reconstruction weighting coefficients between the opposed beams is expressed by Equation (17).

Cone angle artifacts can be reduced by multiplying and adding cone beam reconstruction weighting coefficients ωa and ωb. For example, the cone beam reconstruction weighting coefficients ωa and ωb can be obtained by the following equations. Note that ga is a weighting coefficient for the view βa, and gb is a weighting coefficient for the view βb. When 1/2 of the fan beam angle is γmax, the following formula (18) to formula (23) are obtained.

(Here, for example, q = 1)
For example, as an example of ga and gb, when max [...] Is a function that takes the larger value, the following formulas (24) and (25) 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 the channel i of the multi-row X-ray detector 24 corresponding to the projection data Dr is r0, and the distance coefficient corresponds to the projection data Dr from the focus of the X-ray tube 21. When the distance to the pixel on the reconstruction area P to be set is r1, (r1 / r0) ² .
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 S63, as shown in FIG. 12, 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 S64, steps S61 to S63 are repeated for all views necessary for image reconstruction of tomographic images (that is, views for 360 degrees or "180 degrees + fan angle"), as shown in FIG. Then, back projection data D3 (x, y) is obtained. As shown in FIGS. 13A and 13B, the reconstruction area P may not be a square area of 512 × 512 pixels, but a circular area having a diameter of 512 pixels.

As described above, the crosstalk reduction processing S0 in step S0 of the first embodiment is performed after the preprocessing in step S2, but is performed in the z filter convolution processing in step S4 after the beam hardening processing in step S3. The same effect can be obtained even before the fan-para conversion, which is the conversion from the fan beam to the parallel beam, after the fan-para conversion of the z filter convolution process in step S4, and after the reconstruction function convolution process of step S5.
(Embodiment 2)
Next, in the second embodiment, a crosstalk reduction process when a similar two-dimensional X-ray area detector 34 is used instead of the multi-row X-ray detector 24 will be described. 20 shows an X-ray CT apparatus having a two-dimensional X-ray area detector 34 using one flat X-ray detector, and FIG. 21 shows a two-dimensional X-ray area detection using a plurality of flat X-ray detectors. 2 shows an X-ray CT apparatus having a detector 34.

  In a two-dimensional X-ray area detector 34 having a matrix structure typified by a flat panel X-ray detector, each X-ray detector channel is too fine to determine the position of X-rays to be irradiated. There are many cases where adjustment (alignment) cannot be completed. Therefore, the channel direction and the column direction are determined from the X-ray irradiation position irradiated with the collimator without adjusting the position, and each X-ray detector channel is bundled to a required size to detect each X-ray. By relocating the instrument channels, a data acquisition system can be defined.

  FIG. 22 shows a two-dimensional X-ray area detector 34 composed of original fine X-ray detector channels. The horizontal direction is the channel direction, and the vertical direction is the column direction. FIG. 23 shows a two-dimensional X-ray area detector 34 consisting of X-ray detector channels rearranged in appropriate X-ray detector channels. FIG. 24 shows an enlarged view of a newly rearranged two-dimensional X-ray detector channel on each channel and each row of X-ray detectors of the original two-dimensional X-ray area detector 34.

FIG. 25 is a flowchart showing an outline of the tomographic and scout imaging operations of the X-ray CT apparatus 100 of the present invention.
In step S1, in the helical scan, the X-ray detector 21 is rotated while the X-ray tube 21 and the two-dimensional X-ray area detector 34 are rotated around the subject and the cradle 12 on the imaging table 10 is moved linearly. The data is collected, and the table linear movement z-direction position Ztable (view) is added to the X-ray detector data D0 (view, j, i) represented by the view angle view, the detector row number j, and the channel number i. ) To collect X-ray detector data. In the variable pitch helical scan, data collection is performed not only at a constant speed in the helical scan but also during acceleration and deceleration.

  In the conventional scan (axial scan) or cine scan, the data acquisition system is rotated one or more times while the cradle 12 on the imaging table 10 is fixed at a certain z-direction position to collect data of X-ray detector data. Do. If necessary, after moving to the next position in the z direction, the data acquisition system is rotated once or a plurality of times to collect data of X-ray detector data. In scout imaging, the X-ray tube 21 and the two-dimensional X-ray area detector 34 are fixed, and the X-ray detector data is collected while the cradle 12 on the imaging table 10 is linearly moved. To do.

  When performing image reconstruction by performing three-dimensional backprojection or two-dimensional backprojection, such as conventional scan (axial scan), cine scan, helical scan, or variable pitch helical scan, an appropriate one as shown in FIG. It is effective from the viewpoint of reducing the data processing amount of the following preprocessing and reconstruction function superimposition processing to perform the processing to be bundled in the X-ray detector channel at the time of data collection. The process of bundling the appropriate X-ray detector channel, that is, the process of adding data, may be performed on the data acquisition device (DAS) 25, may be performed on the data acquisition buffer 5, or may be performed after the data acquisition. You may perform by the process of the software (software) on the processing apparatus 3. FIG.

  In the bundling process, the coordinates (X, Y) on the two-dimensional X-ray area detector 34 after the bundling are less than the coordinates (x, y) on the two-dimensional X-ray area detector 34 before the bundling. Thus, it is obtained by affine transformation. However, a, b, c, d, e, and f are real constants.

In general, in the planar type two-dimensional X-ray area detector 34, if neither the (X, Y) coordinate system nor the (x, y) coordinate system is distorted, it is obtained by the affine transformation of the angle θ rotation as follows. .

  When k> 1, it means that the number of channels and the number of columns of the two-dimensional X-ray area detector 34 after being bundled are reduced. In this case, rotation affine transformation may be performed after the X-ray detector data is bundled in advance as follows.

  If the original X-ray projection data is d (x, y), the bundled X-ray projection data is d (X, Y), and the intermediate X-ray projection data is d ′ (x ′, y ′),

Then, d (x, y) is compressed to d '(x', y ') in the x direction and 1 / k in the y direction. If k = 2,

It becomes. As a result, the data is bundled and the noise and S / N of the data are improved.
Thereafter, the following rotation affine transformation may be performed.

  In step S2, the X-ray detector data D0 (view, j, i) is preprocessed and converted into projection data. As shown in FIG. 6, the preprocessing includes step S21 offset correction, step S22 logarithmic conversion, step S23 X-ray dose correction, and step S24 sensitivity correction. In the case of scout image capturing, the preprocessed X-ray detector data is displayed as a scout image by displaying the pixel size in the channel direction and the pixel size in the z direction, which is the cradle linear movement direction, in accordance with the display pixel size of the monitor 6. Completion.

Thereafter, the crosstalk reduction correction in step S0 shown in the first embodiment is performed. The beam hardening correction and subsequent steps after step S3 are the same as those in the first embodiment.
(Embodiment 3)
In the third embodiment, an example of controlling whether or not to use the crosstalk reduction process depending on the shooting condition is shown. In the crosstalk reduction processing shown in the first embodiment or the second embodiment, data is bundled in the channel direction or the column direction, so that the effect is great when used together with low-frequency image reconstruction. In this case, the crosstalk correction can be eliminated or the coefficient of the crosstalk correction can be weakened.

That is, in the photographing condition setting means as shown in FIG. 14, when a low frequency image reconstruction function is selected, the crosstalk correction is eliminated and the image reconstruction is performed or the weak crosstalk correction is applied. It is sufficient to reconstruct the image.
Also, if the imaging condition setting means has been set to reconstruct a thick slice tomogram by applying a wide z-direction filter in the z-direction, crosstalk correction is performed using crosstalk reduction processing. It is sufficient to reconstruct an image without losing the image, or to reconstruct an image with weak crosstalk correction.

  As described above, the presence or absence of crosstalk correction and the strength of the crosstalk can be reduced simply by performing the crosstalk reduction processing depending on the image reconstruction function of the imaging conditions determined by the imaging condition setting means and the z filter processing determined from the slice thickness. And optimal image reconstruction can be performed.

  In the X-ray CT apparatus 100 described above, according to the X-ray CT apparatus or the X-ray CT imaging method of the present invention, or a two-dimensional X-ray area detector having a matrix structure represented by a flat panel X-ray detector. In conventional scan (axial scan), cine scan, helical scan, or variable pitch helical scan of an X-ray CT apparatus, exposure reduction and image quality improvement can be realized.

  Note that the image reconstruction method in this embodiment may be a three-dimensional image reconstruction method based on a conventionally known Feldkamp method. Furthermore, other three-dimensional image reconstruction methods may be used. Alternatively, two-dimensional image reconstruction may be used.

  Also, in this embodiment, column direction (z direction) filters with different coefficients are superimposed on each column to adjust image quality variation, and achieve uniform slice thickness, artifact, and noise image quality in each column. However, various z-direction filter coefficients are conceivable for this, but any of them can produce the same effect.

  In this embodiment, it is written based on a medical X-ray CT apparatus, but it 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 another apparatus. .

In the present embodiment, the image reconstruction of the tomographic image is described, but the same effect can be obtained also in the image reconstruction of the scout image.
In this embodiment, the weighted addition coefficient is “1” and added in the channel direction and the column direction, respectively, as in Expression 1, Expression 2, and Expression 3. However, weighted addition processing is performed by applying the weighting coefficient. Also good.

1 is a block diagram showing an X-ray CT apparatus according to an embodiment of the present invention. It is explanatory drawing which looked at X-ray generator (X-ray tube) and xy plane. It is explanatory drawing which looked at X-ray generator (X-ray tube) and yz plane. It is a flowchart which shows the flow of subject imaging | photography. It is a flowchart which shows schematic operation | movement of the X-ray CT apparatus which concerns on one embodiment of this invention. It is a flowchart which shows the detail of pre-processing. It is a flowchart which shows the detail of a three-dimensional image reconstruction process. It is a conceptual diagram which shows the state which projects the line on a reconstruction area | region to a X-ray transmissive direction. It is a conceptual diagram which shows the line projected on the detector surface. It is a conceptual diagram which shows the state which projected the projection data Dr (view, x, y) on the reconstruction area. It is a conceptual diagram which shows the backprojection pixel data D2 of each pixel on a reconstruction area. It is explanatory drawing which shows the state which obtains backprojection data D3 by adding all the views to backprojection pixel data D2 corresponding to a pixel. It is a conceptual diagram which shows the state which projects the line on a circular reconstruction area | region to a X-ray transmissive direction. It is a figure which shows the imaging condition input screen of a X-ray CT apparatus. It is a figure which shows the conventional X-ray detector. It is a figure which shows the crosstalk rate of each channel of the conventional X-ray detector. FIG. 5A is a diagram illustrating data collection according to the first embodiment, and FIG. 5B is a diagram illustrating crosstalk reduction processing according to the first embodiment. It is a figure which shows the crosstalk rate of 1 channel of the X-ray detector by this Embodiment 1. FIG. It is explanatory drawing explaining the crosstalk which arises between channels. 1 is a diagram showing an X-ray CT apparatus having a two-dimensional X-ray area detector using one planar X-ray detector. FIG. A diagram showing an X-ray CT apparatus having a two-dimensional X-ray area detector combining a plurality of planar X-ray detectors, and a two-dimensional X-ray area detection combining a plurality of planar X-ray detectors It is a figure (b) which shows the irradiation area | region of the X-ray by the collimator on a container. It is a figure which shows the two-dimensional X-ray area detector which consists of a fine X-ray detector channel. FIG. 2 shows a two-dimensional X-ray area detector consisting of X-ray detector channels rearranged in appropriate X-ray detector channels. It is a figure which shows rearrangement of the X-ray detector channel in a two-dimensional X-ray area detector. FIG. 10 is a flowchart of image reconstruction processing in the second embodiment.

Explanation 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 Unit 20 Scanning Gantry 21 X-ray Tube 22 X-ray Controller 23 Collimator 24 Multi-row X-ray Detector 25 DAS ( Data collection device)
26 Rotating part controller 27 Scanning gantry tilt controller 28 X-ray beam forming filter 29 Control controller 30 Slip ring 34 Two-dimensional X-ray area detector

Claims (13)

An X-ray generator ;
A two-dimensional X-ray area detector having a plurality of X-ray detector channels ;
The X-ray generator and the two-dimensional X-ray area detector are arranged opposite to each other with the imaging object interposed therebetween, rotated around the imaging object, generate X-rays from the X-ray generator, and pass through the imaging object . X-ray data collection means for collecting X-ray projection data to be performed;
Image reconstruction means for reconstructing a tomographic image based on the X-ray projection data ;
And imaging condition setting means for setting the shooting conditions used in the previous Symbol collection,
An X-ray CT apparatus comprising:
The X-ray data acquisition means, the number of time, 2 or more integer n to the two-dimensional X-ray area detector of the longitudinal and is definitive in the channel direction the X-ray detector channel width (arrangement interval) and dc The X-ray detection signal of the X-ray detector channel is added n times in the channel direction, and the added X-ray detector data is used as one X-ray projection data in the channel direction. An X-ray CT apparatus that collects data with an interval of dc. An X-ray generator ;
A two-dimensional X-ray area detector having a plurality of X-ray detector channels ;
The X-ray generator and the two-dimensional X-ray area detector are arranged opposite to each other with the imaging object interposed therebetween, rotated around the imaging object, generate X-rays from the X-ray generator, and pass through the imaging object . X-ray data collection means for collecting X-ray projection data to be performed;
Image reconstruction means for reconstructing a tomographic image based on the X-ray projection data ;
And imaging condition setting means for setting the shooting conditions used in the previous Symbol collection,
An X-ray CT apparatus comprising:
The X-ray data acquisition means, when the two-dimensional X-ray area detector of a longitudinal direction substantially perpendicular to definitive in the column direction the X-ray detector channel width (arrangement interval) and dr, 2 or more integer m The number of X-ray detection signals of the X-ray detector channels is added m by one in the column direction, and the added X-ray detector data as one X-ray projection data is added to the column direction. X-ray CT apparatus, which collects at intervals of dr. An X-ray generator ;
A two-dimensional X-ray area detector having a plurality of X-ray detector channels ;
The X-ray generator and the two-dimensional X-ray area detector are arranged opposite to each other with the imaging object interposed therebetween, rotated around the imaging object, generate X-rays from the X-ray generator, and pass through the imaging object . X-ray data collection means for collecting X-ray projection data to be performed;
Image reconstruction means for reconstructing a tomographic image based on the X-ray projection data ;
And imaging condition setting means for setting the shooting conditions used in the previous Symbol collection,
An X-ray CT apparatus comprising:
The X-ray data acquisition means, the two-dimensional X-ray area detector of the longitudinal and is definitive in the channel direction the X-ray detector channel width (arrangement interval) and dc, of the two-dimensional X-ray area detector When the width (arrangement interval) of the X-ray detector channels in the column direction substantially perpendicular to the longitudinal direction is dr , the integers n and m are numbers of 2 or more, n in the channel direction, and the column direction The X-ray detection signals of the X-ray detectors are added in increments of m, and the added X-ray detector data as one X-ray projection data is added to the dc in the channel direction and the column direction. And an X-ray CT apparatus that collects data with an interval of dr.   4. The X-ray CT apparatus according to claim 1, wherein the two-dimensional X-ray area detector includes a multi-row X-ray detector or a matrix-structured two-dimensional X-ray area detector. 5. X-ray CT apparatus that is characterized.   5. The X-ray CT apparatus according to claim 1, wherein the X-ray data collection unit has at least one of the numbers of n and m for performing the addition as two. X-ray CT apparatus.   6. The X-ray CT apparatus according to claim 1, wherein the X-ray data collection unit sets an interval when performing collection in the channel direction to a value that is an integral multiple of the dc. X-ray CT apparatus that is characterized.   6. The X-ray CT apparatus according to claim 2, wherein the X-ray data collection unit sets an interval when performing the collection in the column direction to a value that is an integral multiple of the dr. X-ray CT apparatus that is characterized.   8. The X-ray CT apparatus according to claim 1, wherein the image reconstruction unit includes the collection interval or the addition in the channel direction or the column direction of the two-dimensional X-ray area detector. 9. An X-ray CT apparatus characterized in that crosstalk correction is turned on and off based on the number of steps performed.   9. The X-ray CT apparatus according to claim 1, wherein the image reconstruction unit includes the collection interval or the addition in the channel direction or the column direction of the two-dimensional X-ray area detector. An X-ray CT apparatus characterized in that a coefficient of crosstalk correction is changed based on the number to perform.   10. The X-ray CT apparatus according to claim 1, wherein the image reconstruction unit includes the collection interval or the addition in the channel direction or the column direction of the two-dimensional X-ray area detector. The X-ray CT apparatus is characterized in that the image reconstruction function set by the imaging condition setting means is changed based on the number to be performed.   11. The X-ray CT apparatus according to claim 1, wherein the image reconstruction unit includes the collection interval or the addition in the channel direction or the column direction of the two-dimensional X-ray area detector. The X-ray CT apparatus is characterized in that the z-direction filter coefficient is changed based on the number to perform.   12. The X-ray CT apparatus according to claim 1, wherein the X-ray data collection unit calculates the collection interval or the number of additions based on the imaging conditions, and An X-ray CT apparatus characterized in that the reconstruction means determines crosstalk correction, the z-direction filter, and an image reconstruction function based on the imaging conditions.   13. The X-ray CT apparatus according to claim 12, wherein the X-ray data collection means includes a noise index value of the imaging condition, and calculates the collection interval or the number to be added based on the noise index value. The X-ray CT apparatus, wherein the image reconstruction unit determines crosstalk correction, the z-direction filter, and an image reconstruction function based on the imaging condition.

Images