JPH09327454A - Differential image photographing method and x-ray ct apparatus - Google Patents

Abstract

PROBLEM TO BE SOLVED: To obtain an accurate contrast image by producing a differential image based on substantial difference between data of an identical position of a subject obtained by a plurality of continuous scanning. SOLUTION: Fluoroscopic data of a subject are collected through a detector array of a two-row detector array by helical scanning and stored in a memory 7. An image reorganization device of a central processing unit 3(CPU) reorganizes a CT image of the subject using the measurement data stored in the memory 7. Differential calculation is carried out by the CPU 3 between a stored image in an image buffer and an image of the identical section after one scanning time. The differential calculation is carried out on every corresponding picture elements to form the differential image. The differential image represents a change amount of an identical section within one scanning time. Thereby, an image of a part not filled with a contrasting agent is completely removed to enable to obtain an accurate contrast image.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a differential image capturing method and an X-ray CT (Computed Tomography) apparatus, and more particularly, to a differential image capturing method and an X-ray when a blood vessel or the like is imaged using a contrast agent or the like. Regarding a CT device.

[0002]

2. Description of the Related Art When an image of, for example, a blood vessel of a subject is to be imaged using an X-ray CT apparatus, first, the subject before injection of a contrast agent is scanned to obtain a tomographic image (plane image) thereof.
Then, a contrast agent is injected, and after a predetermined time (about several tens of seconds), the same site is scanned to capture a tomographic image (contrast image), and the difference between the two images (subtractio
n)). As a result, the portion common to both images is erased and only the portion where the contrast agent has spread, that is, the blood vessel image is drawn.

[0003]

In this case, since there is a time lag between the two scans, it is inevitable that the tissue in the subject will move due to breathing motions or the like during that time. Therefore, when the difference is calculated, a difference image due to the change in the tissue position is also generated. This difference image is an image irrelevant to the contrast agent, and interferes with correct interpretation of the image.

The present invention has been made to solve the above problems, and an object thereof is to realize a differential image pickup method and an X-ray CT apparatus for obtaining an accurate contrast image.

[0005]

[Means for Solving the Problems]

[1] A first invention for solving the problem is an X-ray CT apparatus for generating an image based on data obtained by helically scanning an object with an X-ray irradiation source and an X-ray detector facing the X-ray irradiation source. The differential image capturing method according to claim 1, wherein the differential image is generated based on a substantial difference between data of each time obtained by continuously scanning the same site of the subject a plurality of times. This is an imaging method.

According to the first invention for solving the problem, the difference image is generated based on the substantial difference between the data of each time obtained by continuously scanning the same region of the subject a plurality of times. Since this is done, the time difference between the data for which the difference is obtained can be shortened, and thereby a difference image capturing method for obtaining an accurate contrast image can be realized.

In the first invention for solving the problem, "substantial difference" means at least a difference between data of each time or a difference of images respectively generated based on the data of each time. Shall mean.

[2] A second invention for solving the problem is an X-ray irradiation source, and an N-type detector array in which a large number of X-ray detectors are arranged in one row facing the X-ray irradiation source. (≧ 2) multi-row detector array arranged in parallel, and the X-ray irradiation source around the subject while linearly moving the subject along the juxtaposed direction of the multi-row detector array X having a data collecting means for collecting data through the multi-row detector array while continuously rotating the X and an image generating means for generating an image based on the data collected by the data collecting means.
A line CT apparatus, a difference for generating a difference image based on a substantial difference between the data of each group obtained through the detector array of each row of the multi-row detector array for the same site of the subject An X-ray CT apparatus having an image generating means.

According to the second invention for solving the problem, the substantial difference between the data of each group obtained through the detector array of each row of the multi-row detector array for the same site of the subject is obtained. Since the difference image is generated based on the above, it is possible to reduce the time difference between the data for which the difference is obtained, thereby realizing the X-ray CT apparatus that obtains an accurate contrast image.

In the second invention for solving the problem, the difference image generating means generates the difference image based on the difference between the data of each group, which speeds up the generation of the difference image. Is preferred.

In the second invention for solving the problem, the difference image generating means may generate a difference image based on a difference between the images generated based on the data of each time. This is preferable in that a finer difference image is obtained by finely adjusting the image.

[0012]

BEST MODE FOR CARRYING OUT THE INVENTION Hereinafter, embodiments of the present invention will be described in detail with reference to the drawings. Note that the present invention is not limited to the embodiment.

[Overall Configuration] FIG. 1 shows a block diagram of an X-ray CT apparatus. This device is an example of an embodiment of the present invention. Note that an example of an embodiment relating to the device of the present invention is shown by the configuration of the present device. Further, an example of an embodiment relating to the method of the present invention is shown by the operation of the present apparatus.

In FIG. 1, an X-ray CT apparatus 100 includes an operation console 1 and an imaging table 1.
0 and a scanning gantry 20. The operation console 1 includes an input device 2 for inputting an operator's instruction and information, and a central processing unit 3 for executing a helical scan process, an image reconstruction (image generation) process, and the like.
And the control signal and the like are transmitted to the imaging table 10 and the scanning gantry 20.
A control interface (interface) 4 for outputting data to the scanning gantry 20, a data collection buffer (buffer) 5 for collecting data acquired by the scanning gantry 20, and a CRT (cathod-ra) for displaying images and the like.
y tube) 6 and a storage device 7 for storing various data and programs.

The imaging table 10 is designed so that a subject (not shown) can be placed and moved in the body axis direction. The scanning gantry 20 includes an X-ray tube 30, a collimator 50 that forms an X-ray beam, and a two-row detector array.
(array) 60, X-ray controller 21 for adjusting X-ray irradiation timing and intensity, and collimator 50
Controller 22 for adjusting the X-ray passage aperture of the
And a data collection unit 23 that collects data detected by the two-row detector array 60, and an X-ray tube 30 around the body axis of the subject.
And a rotation controller 24 for rotating the two-row detector array 60 and the like.

Here, the X-ray tube 30 is an example of an embodiment of the X-ray irradiation source in the present invention. Two-row detector array 6
Reference numeral 0 is an example of an embodiment of the X-ray detector and the multi-row detector array in the present invention. It should be noted that the detector array is not limited to two rows, but may be multi-rows of three rows or more.

The scanning gantry 20 is an example of an embodiment of the data collecting means in the present invention. The central processing unit 3 is an example of an embodiment of the image generating means and the difference image generating means in the present invention.

FIG. 2 shows a conceptual configuration of the two-row detector array 60. The two-row detector array 60 is an integration of a first-row detector array 61 and a second-row detector array 62. The detector array 61 in the first row forms a multi-channel detector array in which a large number (for example, 1000) of X-ray detectors 61 (i) are arranged in an arc shape.

Similarly, the detector array 62 in the second row also forms a multi-channel detector array in which a large number of X-ray detectors 62 (i) are arranged in an arc shape. Here, i is a channel number, and i = 1 to I.

The X-ray detectors 61 (i) and 62 (i) are constituted by solid-state detectors such as scintillation X-ray detectors and semiconductor X-ray detectors. FIG.
FIG. 4 is an explanatory diagram of a mutual relationship between the X-ray tube 30, the collimator 50, and the two-row detector array 60. In addition, (a) of FIG.
Is a front view, and (b) is a side view.

The X-rays emitted from the X-ray tube 30 are shaped into a flat fan-shaped X-ray beam Xr by the collimator 50, and the detector array 61 of the first row of the two-row detector array 60 is formed.
And the detector array 62 in the second row is evenly illuminated.

Here, the straight line L connecting the focal point of the X-ray tube 30 and the center of the two-row detector array 60 is called an angle reference axis. Further, the focus of the X-ray tube 30 and the individual X-ray detectors 61 (i),
The angle formed by the straight line connecting 62 (i) with respect to the angle reference axis L is called a channel angle γ.

In the central X-ray detectors 61 (I / 2) and 62 (I / 2) of the two-row detector array 60, the channel angle is γ = 0. In the X-ray detectors 61 (1) and 62 (1) at the left end of the two-row detector array 60 in the figure, the channel angle is γ = −γm. In the X-ray detectors 61 (I) and 62 (I) at the right end of the two-row detector array 60 in the figure, the channel angle is γ = + γm.

The channel number i and the channel angle γ are 1
Since it corresponds to one-to-one, for convenience of explanation, in the following, X
X-ray detector 61
It is expressed as (γ), 62 (γ).

The subject is carried in with the body axis orthogonal to the fan surface of the X-ray beam Xr. This state is shown in FIG. As shown in the figure, the projection image of the subject sliced by the fan-shaped X-ray beam Xr is projected on the two-row detector array 60, and the data is measured.

The thickness of the X-ray beam Xr at the ISO center of the object OB gives the slice thickness T. X-ray beam X
r is applied to the detector arrays 61 and 62 in such a manner that the slice thickness T is divided into two equal parts. That is, the slice thickness for each detector array is th = T / 2.

The X-ray tube 30, the collimator 50, and the two-row detector array 60 maintain the relationship as described above, and the object O
A helical scan is performed by continuously rotating around B while the imaging table 10 on which the subject OB is placed is continuously linearly fed in the body axis direction. Helical scan 1
Projection data of the subject is acquired at multiple (eg, 1000) view angles per rotation.

[Collecting View Data] FIG. 5 is an explanatory diagram of view angles. An angle β formed by the angle reference axis L and the vertical axis at one angular position where the X-ray tube 30 and the two-row detector array 60 are rotated is called an absolute view angle. When the absolute view angle at the desired image generation position is β = β ₀ ,
θ = β−β _{0 is referred} to as a relative view angle.

The data collected at the relative view angle θ by the X-ray detector with the channel angle γ is represented by D (γ, θ).
For example, if γ = 0, it is D (0, θ), if γ = −γm, it is D (−γm, θ), and if γ = + γm, it is D (+ γ
m, θ).

FIG. 6 is an explanatory diagram of the data of the opposite view. In FIG. 6, data D (-γm,
When attention is paid to (θ), the data of the opposing view closest to the linear movement axis of the helical scan is data D (+ γ) obtained in front of the image generation position, as shown in (b).
m, θ ′) or data D (+ γm, θ ″) obtained behind the image generation position as shown in (c). Here, θ ′ = θ−π−2 (−γm,), θ ″ =
θ + π−2 (−γm).

Generally, the data of the opposite view of the data D (γ, θ) is D (−γ, θ−π) in front of the image generation position.
−2γ), and D (−γ, θ + π−2γ) on the rear side. FIG. 7 is a diagram illustrating the position of the relative view angle θ on the linear movement axis z of the helical scan.

As shown in FIG. 7B, the position on the z-axis corresponding to the middle of the detector array 61 and the detector array 62 at the absolute view angle β ₀ is used as a reference, and z = 0.
And

Further, the thickness (slice thickness) of the fan-shaped X-ray beam Xr incident on each detector array 61, 62 is th, and the linear movement distance for each rotation of the helical scan is d, and p = d / When th is a helical pitch, in FIG. 7, the helical pitch is p = 1.

As shown in FIG. 7A, at the relative view angle θ = −π, the data D1 (γ, −π) obtained by the first detector array 61 is the data of the slice plane at z = 0. And the data D obtained by the second detector array 62
2 (γ, −π) is z = −th slice plane data.

However, the z position of each slice plane is represented by the z position of the center of thickness of each slice in the ISO center of the object OB. Further, as shown in FIG. 7B, at the relative view angle θ = 0, the data D1 (γ, 0) obtained by the first detector array 61 is z = th.
The data D2 (γ, 0) obtained by the second detector array 62 is z = −th / 2.
Data of the slice plane.

Further, as shown in FIG. 7C, the data D1 (γ, π) obtained by the first detector array 61 at the relative view angle θ = π is the data of the slice plane at z = th. And the data D2 (γ, π) obtained by the second detector array 62 is the data of the slice plane at z = 0.

When the state of FIG. 7C is reached, the slice plane of the data obtained by the second detector array 62 is shown in FIG.
This coincides with the slice plane of the data obtained by the first detector array 61 in (a). After that, the second detector array 62 moves helically along the same trajectory as that of the first detector array 61 and measures the data.

The state is shown in FIG. That is, the second detector array 62 is delayed by one scan
The detector array 61 measures the data at the same position where the data was detected.

[Contrast Imaging] FIG. 9 is a block diagram showing the concept of contrast imaging by the present apparatus. A contrast agent is injected into, for example, a blood vessel of the subject OB, and contrast imaging by a helical scan is started. The helical scan is started at an appropriate time after the injection of the contrast agent. This time is determined based on the prediction of the time for the contrast agent to reach the imaging site.

By the helical scan, the X-ray transmission data of the object OB is collected through the detector arrays 61 and 62 of the two-row detector array 60 and stored in the storage device 7.

The image reconstructing device 31 reconstructs a tomographic image of the subject using the measurement data stored in the storage device 7.
The image reconstructing device 31 is included in the central processing unit 3.

The position on the z-axis (image generation position) of the slice whose tomographic image is to be reconstructed, the reconstruction function,
Reconstruction parameters such as image reconstruction interval (reconstruction pitch) and slice thickness of reconstructed image
Is set in advance by the operator through the input device 2, and image reconstruction is performed based on that. Image reconstruction can be performed using, for example, filtered back projection (filte
red back-projection) method.

Prior to image reconstruction, the multiple view data needed to generate the image is calculated. View data is calculated from the data collected by the data collection unit.
By selecting two data that are the same view data or the opposite view data as the desired view and are closest to the image generation position on the z-axis, and weight-add these data according to the distance from the image generation position. Done. In addition, the measurement data set by the same detector array is used for calculation of the view data.

A tomographic image is reconstructed by filtered back projection using the plurality of view data thus calculated. First, it is performed using the data calculated from the measurement data set of the detector array 61, and the image data is temporarily stored in the image buffer 32. The image buffer 32 is provided inside the central processing unit 3. Next, a tomographic image at the same slice position is reconstructed using the data calculated from the measurement data set of the detector array 62. This image is an image of the same section one scan time later.

A subtractor 33 performs a difference operation between this image and the image stored in the image buffer 32. Subtracter 33
Is provided inside the central processing unit 3. The difference calculation is performed for each corresponding pixel to form a difference image. The difference image represents the amount of change within one scan time for the same cross section. This difference image is stored in the storage device 7 through the image processing device 34. The image processing device 34 is provided inside the central processing unit 3.

Difference images are similarly formed and stored in the storage device 7 at other positions on the z-axis. The difference images are appropriately given to the CRT through the image processing device and displayed as a visible image.

The concentration of the contrast agent in the blood changes over time as shown in FIG. 10, for example. That is, the concentration of the contrast agent reaches a certain peak that increases with time and then gradually decreases. In the case of an artery, both the increase rate and the peak value are high as in the curve A, and in the case of the vein, both the increase rate and the peak value are low as in the curve B.

In the course of such a concentration change, for example, the contrast agent concentration is a1 based on the measurement data of the detector array 61.
When the image in the state of (b1) is reconstructed, the image reconstructed based on the measurement data of the detector array 62 becomes, for example, the a2 (b2) density state after one scan time has elapsed.

Therefore, in the difference image, the artery is represented by the CT value corresponding to the density difference a2-a1 and the vein is represented by the CT value corresponding to the density difference b2-b1. Further, the portion where the contrast agent does not flow is offset by the difference calculation.

Here, the time difference between the two images for which the difference is obtained is the rotation time of one scan, that is, one pitch of the helical scan, and is usually 1 second, for example, 0.5 if it is short.
Seconds, at most 2 seconds. Therefore, since it is easy to stop the body movement of the subject OB by holding the breath during that time, it is possible to eliminate the change in the position of the tissue, thereby completely erasing the image of the portion where the contrast agent does not flow in, and obtaining an accurate contrast image. Images can be obtained.

Further, since it is not necessary to pick up a plane image, it is not necessary to pick up an image of a subject twice twice as in the conventional case, and the efficiency is greatly improved. The CT value of the blood vessel image in the contrast image captured by this apparatus corresponds to the change in the concentration of the contrast agent within a certain period of time, and therefore represents the blood flow velocity. That is, the present apparatus performs blood flow velocity imaging. This is a feature of the present invention that cannot be realized by the conventional difference between the plain image and the contrast image.

With such characteristics, the arterial phase and the venous phase can be clearly discriminated in one contrast image, and for a malignant tumor part such as cancer, the blood flow velocity of the main body and the peripheral part thereof can be obtained. The lesion area can be clearly discriminated based on the difference.

Incidentally, in the generation of the difference image, instead of obtaining the difference of the reconstructed image, the difference of the view data may be obtained before the back projection and the back projection may be performed using this difference data.
This is due to the linearity of the back projection operation,
The addition and subtraction of data is based on the fact that the same result can be obtained before or after back projection.

In this case, for example, as shown in FIG. 11, the view data calculation device 35 first uses the measurement data set of the detector array 61 collected in the storage device 7 to obtain the data of a plurality of views at the image generation position. Is calculated. This view data is stored in the data buffer 36.

Next, using the measurement data set of the detector array 62 collected in the storage device 7, data of a plurality of views at the same image generation position is calculated. This view data is subtracted in the subtracter 33 between the same views as the data in the data buffer 36 to form difference view data.

The image reconstructing device 31 reprojects the image by backprojecting the differential view data. The image reconstructed because the view data is difference data becomes a difference image. The difference image is stored in the storage device 7 through the image processing device 34, is also appropriately read and displayed on the CRT 6.

This method is preferable in that the image reconstruction only needs to be performed once and the generation of the difference image is speeded up. On the other hand, the method of subtracting the images described above is preferable in that a finer difference image is obtained by finely adjusting each image.

In the above, the example of performing the helical scan has been described. However, the imaging table 10 is stopped and a normal axial scan is performed, and the imaging table 10 is fed by the slice thickness th each time one scan is completed. May be.

Also in this case, since the measurement data sets having a time difference of one scan can be obtained by the detector arrays 61 and 62, the difference image can be generated in the same manner as described above. This method is preferable because it does not require view data calculation as in the case of helical scan.

Although an example in which the detector array has two rows has been described, when the number of rows is three or more, the time difference between the measurement data of the first row and the measurement data of the subsequent rows is an integer of the scan time. Since it increases twice, it is possible to obtain a difference image for a desired time difference by selecting a combination of data pairs for which a difference is obtained.

[0061]

As described above in detail, according to the first invention for solving the problem, the substantial portion between the data of each time obtained by continuously scanning the same region of the subject a plurality of times. Since the difference image is generated based on the effective difference, the time difference between the data for which the difference is obtained can be shortened, and thereby the difference image capturing method for obtaining an accurate contrast image can be realized.

Further, according to the second invention for solving the problem, there is a substantial difference between the data of each group obtained through the detector array of each row of the multi-row detector array for the same site of the subject. Since the difference image is generated based on the difference, the time difference between the data for which the difference is obtained can be shortened,
Thereby, an X-ray CT apparatus that obtains an accurate contrast image can be realized.

[Brief description of drawings]

FIG. 1 is a block diagram of an apparatus according to an example of an embodiment of the present invention.

FIG. 2 is a conceptual configuration diagram of a two-row detector array in the apparatus according to the example of the embodiment of the present invention.

FIG. 3 is a conceptual configuration diagram of an X-ray irradiation / detection system in an apparatus according to an example of an embodiment of the present invention.

FIG. 4 is a conceptual configuration diagram of an X-ray irradiation / detection system in an apparatus according to an example of an embodiment of the present invention.

FIG. 5 is an explanatory diagram of views and data in the apparatus according to the embodiment of the present invention;

FIG. 6 is an explanatory diagram of views and data in the apparatus according to the embodiment of the present invention;

FIG. 7 is an explanatory diagram of a helical scan in the apparatus according to the embodiment of the present invention;

FIG. 8 is a conceptual explanatory diagram of data collection in the apparatus according to the example of the embodiment of the present invention.

FIG. 9 is a conceptual explanatory diagram of differential image generation in the apparatus according to the example of the embodiment of the present invention.

FIG. 10 is an explanatory diagram of contrast imaging in the apparatus according to the example of the embodiment of the present invention.

FIG. 11 is a conceptual explanatory diagram of differential image generation in the apparatus according to the example of the embodiment of the present invention.

[Explanation of symbols]

 100 X-ray CT device 1 Operation console 2 Input device 3 Central processing unit 4 Control interface 5 Data acquisition buffer 6 CRT 7 Storage device 10 Imaging table 20 Scanning gantry 21 X-ray controller 22 Collimator controller 23 Data acquisition unit 24 Rotation controller 30 X-ray Tube 50 Collimator 60 Two-row detector array OB Subject 31 Image reconstructor 32 Image buffer 33 Subtractor 34 Image processor 35 View data calculator 36 Data buffer

Claims (2)

[Claims] 1. A differential image capturing method in an X-ray CT apparatus for generating an image based on data obtained by helically scanning an object with an X-ray irradiation source and an X-ray detector facing the X-ray irradiation source. A differential image capturing method, wherein a differential image is generated based on a substantial difference between data of respective times obtained by continuously scanning the same site of a sample a plurality of times. 2. An X-ray irradiation source, and an N-type detector array which is opposed to the X-ray irradiation source and has a large number of X-ray detectors arranged in one row.
(≧ 2) multi-row detector array arranged in parallel, and the X-ray irradiation source around the subject while linearly moving the subject along the juxtaposed direction of the multi-row detector array An X-ray CT apparatus having a data collecting unit that collects data through the multi-row detector array while continuously rotating the lens and an image generating unit that generates an image based on the data collected by the data collecting unit. And a difference image generating means for generating a difference image based on the substantial difference between the data of each group obtained through the detector array of each row of the multi-row detector array for the same part of the subject. An X-ray CT apparatus characterized by: