WO2012057023A1

WO2012057023A1 - Radiographic imaging system and control method for same

Info

Publication number: WO2012057023A1
Application number: PCT/JP2011/074294
Authority: WO
Inventors: 拓司多田
Original assignee: 富士フイルム株式会社
Priority date: 2010-10-28
Filing date: 2011-10-21
Publication date: 2012-05-03

Abstract

The present invention enhances the precision of corrected data acquired in pre-imaging. When an imaging site in a patient to be imaged is selected, a first X-ray dose (I1) is computed on the basis of the X-ray transmittance of the imaging site. A determination is made as to whether pixel data are saturated in the case that the first X-ray dose (I1) is a multiple of α (for example, 100). When the pixel data are not saturated, the α-multiple of the first X-ray dose (I1) is used as a second X-ray dose (I2). When the pixel data are saturated, the α/ν multiple (where ν is a positive number 2 or greater) of the first X-ray dose (I1) is used as the second X-ray dose (I2). Pre-imaging is performed with the exposure conditions being determined on the basis of the second X-ray dose (I2). In the case that the second X-ray dose (I2) is determined by multiplying the first X-ray dose (I1) by α/ν, X-ray irradiation and image detection are executed ν times or more at each scanning position of pre-imaging.

Description

Radiographic imaging system and control method thereof

The present invention relates to a radiographic imaging system that detects an image based on a phase change of radiation and a control method thereof.

Radiation, such as X-rays, is known to have a greater phase change due to the subject than an intensity change. Focusing on this characteristic, research on X-ray phase imaging for obtaining a high-contrast image (hereinafter referred to as a phase contrast image) from a subject having a low X-ray absorption capacity based on the phase change of X-rays has been actively conducted. Yes.

An X-ray imaging system using two grids (diffraction gratings) and an X-ray image detector is known as a kind of X-ray phase imaging (see, for example, Patent Document 1 and Non-Patent Document 1). In this X-ray imaging system, a first grid is disposed behind the subject as viewed from the X-ray source, a second grid is disposed downstream from the first grid by a Talbot distance, and the X-ray is behind the first grid. It is configured by arranging an image detector. The Talbot distance is a distance at which X-rays that have passed through the first grid form a self-image (stripe image) of the first grid due to the Talbot interference effect. This self-image is modulated by the phase change of the X-rays caused by the subject and refraction. By detecting this modulation amount, the phase change of the X-ray is imaged.

The fringe scanning method is known as a method for detecting the modulation amount. In the fringe scanning method, the second grid is intermittently moved with respect to the first grid, and photographing is performed while the second grid is stopped. This intermittent movement is performed by a constant scanning pitch obtained by equally dividing the lattice pitch in a direction substantially parallel to the surface of the first grid and substantially perpendicular to the lattice line direction of the first grid. From the intensity change of each pixel value obtained by the X-ray image detector, a phase differential image representing the distribution of the X-ray refraction angle in the subject is obtained. A phase contrast image is obtained by integrating the phase differential image. This fringe scanning method is also used in an imaging apparatus using laser light (see, for example, Non-Patent Document 2).

In the fringe scanning method, if there is a manufacturing error or misalignment in the first and second grids, noise corresponding to the manufacturing error or misalignment occurs in the phase differential image. In Patent Document 1, in order to remove this noise, apart from the above-described series of imaging in a state where the subject is arranged (hereinafter referred to as main imaging), the series of imaging (hereinafter, referred to as “subject” is not described). It has been proposed to subtract a phase differential image (correction data) obtained by pre-imaging from a phase differential image obtained by actual imaging.

Japanese Patent No. 4445397

However, since the phase change of the X-ray by the subject is minute, the phase differential image inhibits the imaging of the subject even with minute noise. For this reason, it is desired to obtain correction data with high accuracy and to remove noise with high accuracy. Japanese Patent Application Laid-Open No. 2004-228561 does not describe any improvement in accuracy of correction data.

It is an object of the present invention to provide a radiographic image capturing system and a control method thereof that can improve the accuracy of correction data for correcting a phase differential image.

In order to achieve the above object, a radiographic imaging system of the present invention includes a radiation source, a first grid, an intensity modulation unit, a radiographic image detector, a phase differential image generation unit, and a correction data generation unit. A correction processing unit and a control unit. The radiation source emits radiation at the time of pre-imaging without arranging the subject and at the time of actual imaging with the subject arranged. The first grid passes the radiation and generates a first periodic pattern image. The intensity modulation unit performs intensity modulation on the first periodic pattern image to generate a second periodic pattern image. The radiological image detector detects the second periodic pattern image and generates image data. The phase differential image generation unit generates a phase differential image based on the image data generated during the main photographing. The correction data generation unit generates correction data based on the image data generated during the pre-photographing. The correction processing unit subtracts the correction data from the phase differential image. The control unit determines an exposure condition during the pre-imaging based on a second radiation dose obtained by multiplying the first radiation dose during the main imaging by α times (where α> 1). To cause the radiation source to perform radiation.

It is preferable that the control unit determines whether or not the pixel data is saturated when the first radiation dose is multiplied by α. If saturation occurs, the control unit determines the second radiation dose to be a value obtained by multiplying the first radiation dose by α / ν (where ν is a positive number of 2 or more), and during the pre-imaging. , Causing the radiation source and the radiation image detector to execute the radiation and the image detection respectively ν times or more.

It is preferable that the correction data generation unit adds each image data generated by the radiation image detector when the radiation and the image detection are performed ν times or more during the pre-imaging.

The intensity modulation unit applies intensity modulation at a plurality of scanning positions having different phases with respect to the first periodic pattern image to generate a plurality of second periodic pattern images, and the radiation image detector includes the plurality of radiation image detectors. The second periodic pattern image is detected to generate a plurality of image data, and the phase differential image generation unit and the correction data generation unit change the intensity of pixel data with respect to the scanning position based on the plurality of image data The phase differential image and the correction data are respectively generated by calculating the phase shift amount of the intensity modulation signal representing the control signal, and the control unit obtains a value obtained by multiplying the first radiation dose by α / ν. When the radiation dose is used, it is preferable that the radiation and the image detection are executed at least ν times at each scanning position at the time of the pre-imaging.

The correction data generation means adds ν or more image data obtained at each scanning position when the radiation and the image detection are performed at ν times or more at each scanning position during pre-imaging. Is preferred.

It is preferable that an operation unit that enables selection of the type of imaging region is further provided, and the control unit determines the first radiation dose according to the imaging region selected by the operation unit. Further, the control unit may determine the first radiation dose based on a thickest imaging part having the largest thickness among a plurality of imaging parts that can be selected by the operation unit. Further, a photo timer may be further provided, and the control unit may set a set value of the photo timer as the first radiation dose.

The α is preferably 100 or more.

The intensity modulation unit includes a second grid having a periodic pattern in the same direction as the first periodic pattern image, and a scanning mechanism that moves one of the first and second grids at a predetermined pitch. Is preferred.

The first grid is an absorption grid, and it is preferable to project incident radiation onto the second grid geometrically. Further, the first grid may be an absorption grid or a phase grid, and may cause a Talbot interference effect on incident radiation and project it onto the second grid.

The control method of the present invention includes a radiation source that emits radiation during pre-imaging without placing a subject and during the main photography with the subject, and a first periodic pattern image that passes through the radiation and generates a first periodic pattern image. Grid, an intensity modulation unit that applies intensity modulation to the first periodic pattern image to generate a second periodic pattern image, and radiation that detects the second periodic pattern image and generates image data An image detector, a phase differential image generation unit that generates a phase differential image based on the image data generated during the main photographing, and a correction that generates correction data based on the image data generated during the pre-photographing The present invention is used in a radiographic system including a data generation unit and a correction processing unit that subtracts the correction data from the phase differential image. In the control method of the present invention, the exposure condition at the time of the pre-imaging is determined based on the second radiation dose obtained by multiplying the first radiation dose at the time of the main imaging by α (where α> 1).

According to the present invention, the exposure condition during pre-imaging is determined based on the second radiation dose obtained by multiplying the first radiation dose during the main imaging by α times (where α> 1). Since the radiation source performs X-ray emission, the accuracy of correction data acquired by pre-imaging can be improved.

It is a schematic diagram which shows the structure of a X-ray imaging system. It is a block diagram which shows the structure of an image process part. It is a block diagram which shows the structure of a X-ray image detector. It is explanatory drawing explaining the structure of a 1st and 2nd grid. It is explanatory drawing explaining a fringe scanning method. It is a graph which shows an intensity | strength modulation signal. It is a flowchart which shows the determination method of the exposure conditions at the time of pre imaging | photography. It is explanatory drawing explaining the addition process of pixel data.

In FIG. 1, radiation, for example, an X-ray imaging system 10 includes an X-ray source 11, an imaging unit 12, a memory 13, an image processing unit 14, an image recording unit 15, an imaging control unit 16, a console 17, and a system control unit 18. Prepare. As is well known, the X-ray source 11 has a rotary anode type X-ray tube (not shown) and a collimator (not shown) for limiting the X-ray irradiation field, and the subject H is irradiated with X-rays. Irradiate.

The imaging unit 12 includes an X-ray image detector 20, a first grid 21, a second grid 22, and a scanning mechanism 23. The 1st and

2nd grids

21 and 22 are absorption type grids, and are arranged facing the X-ray source 11 in the Z direction which is the X-ray irradiation direction. A space is provided between the X-ray source 11 and the first grid 21 so that the subject H can be arranged. As is well known, the X-ray image detector 20 is a flat panel detector using a semiconductor circuit, and is disposed behind the second grid 22 so that the surface thereof is orthogonal to the Z direction.

The first grid 21 has a plurality of X-ray absorption parts 21a and X-ray transmission parts 21b extended in the Y direction, which is one direction in a plane orthogonal to the Z direction. The X-ray absorption part 21a and the X-ray transmission part 21b are alternately arranged along the X direction orthogonal to the Z direction and the Y direction, and constitute a striped grid. The second grid 22 has a plurality of X-ray absorbing portions 22 a and X-ray transmitting portions 22 b that are extended in the Y direction and arranged alternately along the X direction, like the first grid 21. The

X-ray absorbers

21a and 22a are formed of a metal having X-ray absorption such as gold (Au) and platinum (Pt). The X-ray transmissive

portions

21b and 22b are formed of a material having X-ray permeability such as silicon (Si) or resin or a gap.

The first grid 21 transmits a X-ray emitted from the X-ray source 11 to generate a first periodic pattern image (hereinafter referred to as a G1 image). The second grid 22 generates a second periodic pattern image (hereinafter referred to as a G2 image) by partially shielding (intensity modulating) the G1 image.

The memory 13 temporarily stores a plurality of image data obtained by the photographing unit 12. The image processing unit 14 generates a phase differential image and a phase contrast image based on a plurality of image data stored in the memory 13. The image recording unit 15 records the phase contrast image generated by the image processing unit 14. The imaging control unit 16 controls the X-ray source 11 and the imaging unit 12. The console 17 has a well-known operation unit, monitor, and the like, and enables input of shooting conditions, shooting instructions, etc., and display of image information such as shooting information and images. The system control unit 18 comprehensively controls each unit according to an input signal from the input unit of the console 17.

The X-ray imaging system 10 is an X-ray diagnostic apparatus that images a patient as the subject H. The type of the imaging region (chest, breast, lower limb, hand, shoulder joint, etc.) of the patient to be imaged can be selected by the operation unit of the console 17.

The scanning mechanism 23 intermittently moves the second grid 22 in the X direction, and sequentially changes the relative position of the second grid 22 with respect to the first grid 21. The scanning mechanism 23 has a piezoelectric actuator or an electrostatic actuator. The scanning mechanism 23 is driven based on the control of the imaging control unit 16 at the time of stripe scanning described later. The memory 13 stores image data generated by the X-ray image detector 20 in each scanning step of fringe scanning. The second grid 22 and the scanning mechanism 23 constitute an intensity modulation unit.

2, the image processing unit 14 includes a phase differential image generation unit 30, a correction data generation unit 31, a correction data storage unit 32, a correction processing unit 33, and a phase contrast image generation unit 34. The phase differential image generation unit 30 generates a phase differential image of the subject H based on a plurality of image data obtained by the X-ray image detector 20 in each scanning step at the time of main imaging performed by arranging the subject H. Generate.

The correction data generation unit 31 generates a phase differential image based on a plurality of image data obtained by the X-ray image detector 20 in each scanning step at the time of pre-imaging performed without arranging the subject H. Is stored in the correction data storage unit 32 as correction data. The correction data storage unit 32 is configured by a nonvolatile memory such as a flash memory.

The correction processing unit 33 subtracts the correction data stored in the correction data storage unit 32 from the phase differential image generated by the phase differential image generation unit 30 for each pixel, and a phase differential image after subtraction (hereinafter, corrected). (Referred to as phase differential image) is input to the phase contrast image generator 34. The phase contrast image generation unit 34 generates a phase contrast image by integrating the corrected phase differential image along the X direction.

The phase contrast image generated by the phase contrast image generation unit 34 is recorded in the image recording unit 15 and then output to the console 17 and displayed on the monitor.

3, the X-ray image detector 20 includes an image receiving unit 41, a scanning circuit 42, and a readout circuit 43. The image receiving unit 41 is a two-dimensional array of a plurality of pixels 40 that convert X-rays into electric charges and store them on an active matrix substrate (not shown) along the X and Y directions. The scanning circuit 42 controls the timing for reading out charges from the pixels 40. The readout circuit 43 converts the charges read from the pixels 40 into image data and outputs the image data. The scanning circuit 42 and each pixel 40 are connected to each other by a scanning line 44 for each row. The readout circuit 43 and each pixel 40 are connected to each other by a signal line 45 for each column. The arrangement pitch of the pixels 40 is about 100 μm in each of the X direction and the Y direction.

As is well known, the pixel 40 directly converts X-rays into electric charges by a conversion layer (not shown) such as amorphous selenium, and a capacitor (not shown) connected to the converted charge at the lower electrode of the conversion layer. ) Is a direct conversion type X-ray detection element. Each pixel 40 is provided with a TFT switch (not shown). The gate electrode of the TFT switch is connected to the scanning line 44, the source electrode is connected to the capacitor, and the drain electrode is connected to the signal line 45. When the TFT switch is turned on by the drive pulse applied from the scanning circuit 42, the charge accumulated in the capacitor is read out to the signal line 45.

The pixel 40 temporarily converts X-rays into visible light with a scintillator (not shown) formed of gadolinium oxide (Gd ₂ O ₃ ), cesium iodide (CsI), or the like, and converts the converted visible light into photo An indirect conversion type X-ray detection element that converts the charge into a charge by a diode (not shown) and stores the charge may be used. In addition to the radiation image detector based on the TFT panel, a radiation image detector based on a solid-state imaging device such as a CCD sensor or a CMOS sensor may be used as the X-ray image detector 20.

The readout circuit 43 includes an integration amplifier, an A / D converter, a correction circuit (none of which is shown), and the like. The integrating amplifier integrates the charges output from each pixel 40 via the signal line 45 and converts them into a voltage signal (image signal). The A / D converter converts the image signal converted by the integrating amplifier into digital image data. The correction circuit performs dark current correction, gain correction, linearity correction, and the like on the image data, and inputs the corrected image data to the memory 13.

In FIG. 4, the X-ray emitted from the X-ray source 11 has a cone beam shape with the X-ray focal point 11a as the light emission point, and therefore the G1 image generated by the first grid 21 is from the X-ray focal point 11a. It is enlarged in proportion to the distance. The lattice pitch p ₂ and the width d _{2 in} the X direction of the X-ray absorber 22 a of the second grid 22 are the distance L ₁ between the X-ray focal point 11 a and the first grid 21, the first grid 21 and the first grid 21. Using the distance L ₂ between the two grids 22 and the lattice pitch p ₁ and the width d ₁ of the X-ray absorbing portion 21a of the first grid 21, settings are made as shown in the following equations (1) and (2). Has been.

For example, the grating pitch p ₂ is 5 μm, and the width d ₂ is half that of 2.5 μm. The thickness of the X-ray absorber 21a in the Z direction is set to, for example, about 100 μm in consideration of corneal X-ray vignetting emitted from the X-ray source 11. It is not always necessary to satisfy the formula (2), and the intervals d ₁ and d ₂ may be set independently.

The 1st and

2nd grids

21 and 22 are comprised so that the X-rays which passed X-ray

low absorption part

21b, 22b may be projected linearly (geometrical optical). Specifically, it is realized by setting the distances d ₁ and d ₂ to a value sufficiently larger than the peak wavelength of the X-rays emitted from the X-ray source 11, and most of the X-rays included in the emitted X-rays are diffracted. Without passing through, the

X-ray transmission parts

21b and 22b pass through while maintaining straightness. For example, when tungsten is used as the rotating anode of the aforementioned X-ray tube and the tube voltage is 50 kV, the peak wavelength of the X-ray is about 0.4 mm. In this case, the distances d ₁ and d ₂ may be about 1 μm to 10 μm. The grating pitches p ₁ and p ₂ are about 2 μm to 20 μm.

If the Talbot effect occurs in the first grid 21, the distance L _2, as it is constrained to Talbot distance, in the present embodiment, the first grid 21 to geometrical optics projecting an X-ray because it is constituted, the distance L ₂ can be set independently of the Talbot distance.

The imaging unit 12 of the present embodiment does not constitute a Talbot interferometer, but the Talbot distance Z _m when the Talbot interference effect is assumed to occur in the first grid 21 is the lattice pitch p ₁ , p ₂ , using the X-ray wavelength (peak wavelength) λ and a positive integer m, represented by the following formula (3).

Expression (3) is an expression that represents the Talbot distance when the X-rays emitted from the X-ray source 11 are in the shape of a cone beam. “Atsushi Momose, et al. No. 10, October 2008, 8077 ”.

In the present embodiment, it is possible to set the distance L ₂ as described above independently of the Talbot distance, for the purpose of thinning in the Z direction of the imaging unit 12, the distance L _2, in the case of m = 1 set to the minimum value shorter than a Talbot distance Z _1. That is, the distance L ₂ is set to a range satisfying the following expression (4).

A G2 image generated by intensity-modulating the G1 image generated by the first grid 21 with the second grid 22 is captured by the X-ray image detector 20. There may be a slight difference between the pattern period of the G1 image at the position of the second grid 22 and the lattice pitch p2 of the _second grid 22 due to an arrangement error or the like. In this case, the G2 image Moire fringes occur. Even when moire fringes occur in the G2 image, there is no particular problem if the period of the moire fringes in the X direction or Y direction is larger than the arrangement pitch of the pixels 40.

When the subject H is disposed between the X-ray source 11 and the first grid 21, the G2 image is modulated by the subject H. This modulation amount is proportional to the refraction angle of X-rays by the subject H. The phase differential image corresponds to the distribution of the refraction angle of X-rays.

Next, a fringe scanning method for generating a phase differential image will be described. Here, the coordinates in the X, Y, and Z directions are the x, y, and z coordinates. FIG. 4 illustrates one path of X-rays that are refracted according to the phase shift distribution Φ (x) in the X direction of the subject H. Reference numeral 50 indicates a path along which the X-ray goes straight when the subject H does not exist. X-rays traveling along the path 50 pass through the first and

second grids

21 and 22 and enter the X-ray image detector 20. Reference numeral 51 denotes an X-ray path refracted by the subject H when the subject H exists. X-rays traveling along the path 51 pass through the first grid 21 and are then absorbed by the X-ray absorbing portion 22 a of the second grid 22.

The phase shift distribution Φ (x) of the subject H is expressed by the following formula (5) using the refractive index distribution n (x, z) of the subject H. Here, the y-coordinate is omitted for simplification of description.

The G1 image projected from the first grid 21 to the position of the second grid 22 is displaced in the X direction by an amount corresponding to the refraction angle φ due to refraction of X-rays at the subject H. This displacement amount Δx is approximately expressed by the following equation (6) based on the fact that the refraction angle φ of X-rays is very small.

Here, the refraction angle φ is expressed by the following equation (7) using the X-ray wavelength λ and the phase shift distribution Φ (x).

Thus, the displacement amount Δx is related to the phase shift distribution Φ (x) of the subject H. This displacement amount Δx is calculated by the following equation (8) based on the phase shift amount ψ (the phase shift amount with and without the subject H) of the intensity modulation signal generated for each pixel 40. Are related. The intensity modulation signal is a waveform signal that represents a change in intensity of pixel data in accordance with a change in the relative position (scanning position) between the first grid 21 and the second grid 22 in the X direction.

Therefore, the two-dimensional distribution of the phase shift amount ψ of the intensity modulation signal corresponds to the differential image of the phase shift distribution Φ (x). Here, a two-dimensional distribution of the phase shift amount ψ is a phase differential image. By integrating the phase differential image along the X direction, a phase contrast image representing the phase shift distribution is generated.

In the fringe scanning method, in order to obtain the intensity modulation signal, one of the first and

second grids

21 and 22 is intermittently moved by a predetermined pitch in the X direction relative to the other, and is stopped. Take a photo. In the present embodiment, the first grid 21 is fixed, and the second grid 22 is intermittently moved in the X direction by the scanning mechanism 23. Moire fringes generated in the G2 image move as the second grid 22 moves, and when the translational distance reaches the grating period (grating pitch p ₂ ) of the second grid 22, it matches the original pattern.

FIG. 5 schematically shows a state where the second grid 22 is intermittently moved at a constant scanning pitch (p ₂ / M) obtained by dividing the lattice pitch p ₂ into M pieces. The scanning mechanism 23 sequentially moves the second grid 22 to the M scanning positions of k = 0, 1, 2,..., M−1. M is an integer greater than or equal to 2, for example, M = 5.

In the position of k = 0, mainly the X-ray component (non-refractive component) that has not been refracted by the subject H passes through the second grid 22. When the second grid 22 is moved in order of k = 1, 2,..., X-rays passing through the second grid 22 are refracted by the subject H while the non-refractive component decreases. Component (refractive component) increases. In particular, at the position of k = M / 2, the X-ray that passes through the second grid 22 is substantially only a refractive component. Beyond the position of k = M / 2, the X-ray passing through the second grid 22 has a reduced refractive component while an increased non-refractive component.

When imaging is performed by the X-ray image detector 20 at each scanning position k = 0, 1, 2,..., M−1, M pixel data are obtained for each pixel 40. The M pixel data forms an intensity modulation signal.

Hereinafter, a method of calculating the phase shift amount ψ of the intensity modulation signal will be described. The pixel data I _k (x) at each scanning position k is generally represented by the following equation (9).

Here, A ₀ represents the intensity of incident X-rays. _An is a value corresponding to the contrast of the intensity modulation signal (where n is a positive integer). I is an imaginary unit. φ (x) represents the refraction angle φ as a function of the x-coordinate.

In the present embodiment, since the scanning pitch (p ₂ / M) is constant, the following expression (10) is established. When this relational expression is applied to the expression (9), the refraction angle φ (x) is expressed by the following expression (11).

Here, arg [] means extraction of a declination, and corresponds to the phase shift amount ψ (x) of the intensity modulation signal I _k (x) as shown in the following equation (12).

In the above description, the y coordinate is not considered, but the two-dimensional distribution ψ (x, y) of the phase shift amount is obtained by considering the y coordinate. This two-dimensional distribution ψ (x, y) corresponds to the phase differential image. This phase differential image ψ (x, y) is expressed by the following equation (13) using an arctangent function.

FIG. 6 exemplifies intensity modulation signals I _k (x, y) obtained at the time of main photographing and photographing scanning, respectively. The phase shift amount ψ ₁ (x, y) during the main photographing is calculated by the phase differential image generation unit 30 based on the equation (13). The phase shift amount ψ ₂ (x, y) at the time of one pre-photographing is similarly calculated by the correction data generation unit 31 based on the formula (13). The phase shift amount ψ ₂ (x, y) at the time of pre-photographing is caused by a manufacturing error or a displacement of the first and

second grids

21 and 22. Correction processing unit 33, a phase shift amount ψ ₁ (x, y) from the phase shift amount ψ ₂ (x, y) is subtracted.

The phase differential image is susceptible to noise because the X-ray refraction angle φ of the subject H is very small. This noise is roughly classified into electronic noise and quantum noise. Electronic noise is caused by the X-ray image detector 20. Quantum noise depends on the X-ray dose and becomes apparent at low X-ray doses. Due to these noises, errors occur in the calculated values of the phase shift amounts ψ ₁ (x, y), ψ ₂ (x, y).

The noise variation (standard deviation) N included in the phase shift amount (ψ ₁ (x, y) −ψ ₂ (x, y)) after subtraction by the correction processing unit 33 is the phase shift amount ψ ₁ (x, y ) Variation N1 and phase shift amount ψ ₂ (x, y) variation N2 are expressed by the following equation (14). Hereinafter, N is referred to as total variation, N1 is referred to as first variation, and N2 is referred to as second variation.

The first and second variations N1 and N2 are expressed by the following formula (15) and the following formula (16) using the first and second X-ray doses I1 and I2, respectively. Here, the first X-ray dose I1 is the maximum X-ray dose detected by the X-ray image detector 20 at each scanning position k at the time of main imaging. The second X-ray dose I2 is the maximum X-ray dose detected by the X-ray image detector 20 at each scanning position k during pre-imaging.

The relationship between the noise variation of the phase shift amount and the maximum X-ray dose shown in Equations (15) and (16) was found by the present applicant through simulation. Specifically, the applicant assigns a random error depending on the size of each pixel data constituting the intensity modulation signal, and then calculates the phase shift amount 50 times based on the equation (13). Repeatedly, the variation (standard deviation) in the amount of phase shift was evaluated. Then, this variation was evaluated by changing the maximum value of the intensity modulation signal (corresponding to the maximum X-ray dose detected by the X-ray image detector 20) stepwise. As a result, the present applicant has found that the variation in the noise of the phase shift amount is proportional to the reciprocal of the square root of the maximum X-ray dose. Note that the present applicant has the relationship of the expressions (15) and (16) when the first and second X-ray doses I1 and I2 are the average X-ray doses detected by the X-ray image detector 20. It is found that it holds.

From equations (14) to (16), it can be seen that the first and second X-ray doses I1 and I2 can be increased to reduce the overall variation N. Since it is not desirable to increase the first X-ray dose I1 from the viewpoint of exposure of the subject H, in this embodiment, the overall variation N is reduced by increasing the second X-ray dose I2.

Table 1 shows a change in the total variation N when the second X-ray dose I2 is controlled and the second variation N2 is changed.

From Table 1, if the second X-ray dose I2 is about four times the first X-ray dose I1 and the second variation N2 is about half of the first variation N1, the error caused by the second variation N2 Is about 10% of the total variation N. Further, if the second X-ray dose I2 is about 100 times the first X-ray dose I1 and the second variation N2 is about 1/10 of the first variation N1, an error caused by the second variation N2 Can be almost ignored. Therefore, the X-ray source 11 is controlled, and the second X-ray dose I2 is preferably at least 4 times, more preferably 100 times or more, with respect to the first X-ray dose I1.

On the other hand, if the second X-ray dose I2 is increased too much, there is a possibility that the accumulated charge amount of the pixel 40 of the X-ray image detector 20 is saturated and the pixel data is saturated. For this reason, it is preferable to set the second X-ray dose I2 in a range where the pixel data is not saturated.

Next, a method for determining exposure conditions during pre-photographing will be described with reference to the flowchart shown in FIG. First, when an imaging part of a patient to be imaged (subject H) is selected by the operation unit of the console 17 (step S10: YES), the system control unit 18 X-ray transmittance of the selected imaging part. Based on the above, the first X-ray dose I1 optimal for imaging is calculated (step S11).

Next, the system control unit 18 sets the magnification α of the second X-ray dose I2 with respect to the first X-ray dose I1, for example, 100 times (step S12). In this case, is the pixel data of the pixel 40 saturated? Is determined based on the characteristics of the X-ray image detector 20 (step S13). When it is determined that the pixel data is not saturated (step S13: NO), α times the first X-ray dose I1 is determined as the second X-ray dose I2 (step S14).

On the other hand, when it is determined that the pixel data is saturated (step S13: YES), a value (α / ν) obtained by dividing the magnification α by the positive integer ν is set as a new magnification β (step S15). Β times the X-ray dose I1 of 1 is determined as the second X-ray dose I2 (step S16). The integer ν is the smallest positive integer within a range where the pixel data is not saturated.

Then, the system control unit 18 exposes the exposure condition (the X-ray source 11 of the X-ray source 11) so that the X-ray image detector 20 detects the second X-ray dose I2 determined in step S14 or step S16. The tube voltage, tube current, and exposure time are determined (step S17). Since the X-ray dose depends on the mAs value represented by the product of the tube current and the exposure time, the tube current or the exposure time may be adjusted so as to obtain the determined second X-ray dose I2.

When the second X-ray dose I2 is determined in step S14, the system control unit 18 performs the X-ray emission by the X-ray source 11 and the detection operation by the X-ray image detector 20 at each pre-scanning scanning position k. Is performed only once. On the other hand, when the system control unit 18 determines the second X-ray dose I2 in step S16, the X-ray emission from the X-ray source 11 and the detection by the X-ray image detector 20 at each scanning position k of pre-imaging. The operation is performed ν times. Therefore, in the former case, one piece of image data is obtained at each scanning position k, whereas in the latter case, ν pieces of image data are obtained at each scanning position k.

When ν pieces of image data are obtained at each scanning position k, the correction data generation unit 31 converts the pixel data I _k (x, y) of each pixel 40 into the same scanning position k as shown in FIG. Addition is performed every time to generate one intensity modulation signal. In this case, the correction data described above is calculated using the intensity modulation signal obtained by the addition process. Since the total X-ray dose by ν exposures using the second X-ray dose I2 determined in step S16 is substantially equal to the second X-ray dose I2 determined in step S14, either step S14 or step S16 Thus, even when the second X-ray dose I2 is determined, the noise of the correction data becomes a negligible level. For this reason, the image quality of the corrected phase differential image generated by the correction processing unit 33 is improved.

Note that it is conceivable that the S / N deteriorates by adding a plurality of pixel data I _k (x, y) in the correction data generation unit 31, but this S / N deterioration can be ignored. This is usually because the electronic noise N _e is smaller than the quantum noise N _x (N _x >> N _e ), and if the total X-ray dose is the same, the total noise amount is independent of the number of X-ray radiation divisions. This is because it is almost constant. For example, the total amount of noise when X-ray radiation is performed once with an X-ray dose of 100 mR is expressed as (10N _x ² + N _e ² ) ^1/2, and 10 times of X-ray radiation is performed with an X-ray dose of 10 mR. The total amount of noise when performed is represented as {10 (N _x ² + N _e ² )} ^1/2 . If N _x >> N _e , the two are almost the same.

Next, the operation of the X-ray imaging system 10 will be described. First, when an imaging region is selected by the operation unit of the console 17, the exposure condition at the time of pre-imaging is determined as described above. Next, when an instruction to start pre-imaging is input from the operation unit of the console 17, the second grid 22 is intermittently moved to each scanning position k by the scanning mechanism 23, and X-rays are obtained at each scanning position k. X-ray emission from the source 11 and detection operation by the X-ray image detector 20 are performed, and image data is generated. At each scanning position k, imaging is performed once or ν times according to the determined second X-ray dose I2, and one or ν image data is generated.

Each image data is stored in the memory 13, and correction data is generated by the correction data generation unit 31 of the image processing unit 14. In addition, when ν times of photographing are performed at each scanning position k of the pre-photographing, the pixel data I _k (x, y) at each scanning position k is added by the correction data generation unit 31 and then the intensity is added. A modulated signal is generated. The correction data generated by the correction data generation unit 31 is stored in the correction data storage unit 32 until the next pre-photographing is performed.

Next, when an instruction to start main imaging is input from the console 17 with the subject H placed, the exposure conditions for main imaging are set, and the second grid 22 is set as in the case of pre-imaging. While being moved to each scanning position k, image data is generated by the X-ray emission from the X-ray source 11 and the X-ray image detector 20 at each scanning position k. At the time of main imaging, X-ray emission is performed with the first X-ray dose I1. In this case, the number of X-ray emissions at each scanning position k is one, and one piece of image data is generated at each scanning position k.

The image data is stored in the memory 13 and a phase differential image is generated by the phase differential image generation unit 30 of the image processing unit 14. The phase differential image is input to the correction processing unit 33, and the correction processing unit 33 subtracts the correction data stored in the correction data storage unit 32 from the phase differential image.

Then, based on the corrected phase differential image generated by the correction processing unit 33, a phase contrast image generation unit 34 generates a phase contrast image. The phase contrast image is recorded in the image recording unit 15 and then displayed on the monitor of the console 17.

Hereinafter, other embodiments of the present invention will be described. In the following embodiments, detailed description of the same configuration as that of the above embodiment is omitted.

In the above embodiment, each time an imaging region is selected on the console 17, the exposure conditions for pre-imaging are determined. Of the plurality of imaging regions that can be selected on the console 17, the thickness is the largest (X-rays). The exposure conditions for the pre-imaging may be determined based on the thickest imaging region where the transmittance is low. In this case, since the correction data obtained by the correction data generation unit 31 is applied even when an imaging part other than the thickest imaging part is selected, it is necessary to perform pre-imaging every time the imaging part is changed on the console 17. There is no.

In the above embodiment, the second X-ray dose I2 is determined based on the first X-ray dose I1 calculated based on the imaging region. However, it is described in US Patent Application Publication No. 2006/0023839 and the like. A known phototimer (automatic exposure device) may be provided, and the second X-ray dose I2 may be determined by setting the phototimer setting value (reference X-ray dose for stopping X-ray emission) as the first X-ray dose I1. .

In the above embodiment, when the second X-ray dose I2 is determined in any of Steps S16, ν times of imaging are performed at each scanning position k. However, ν times or more are captured at each scanning position k. May be performed.

In the above embodiment, the first and

second grids

21 and 22 are provided between the X-ray source 11 and the X-ray image detector 20, but the X-ray source 11 and the first and second grids are further provided. A known source grid (source grid) may be provided between the grid 21 and the grid 21.

In the above-described embodiment, the X-rays incident on the first grid 21 are configured to be geometrically optically projected onto the second grid 22, but instead, this is disclosed in Japanese Patent No. 4445397 (US Patent). As described in Japanese Patent No. 7180979), the X-rays incident on the first grid form a self-image of the first grid by the Talbot interference effect and are projected onto the position of the second grid. Also good. In this case, it is necessary to set the distance between the first and second grids to the Talbot distance.

In this case, the first grid may be a phase grid instead of the absorption grid. The phase-type grid has a thickness of the X-ray absorption part and the X-ray transmission part so that a phase difference of “π” or “π / 2” occurs between the X-ray absorption part and the X-ray transmission part. It is configured by setting materials.

In the above embodiment, the subject H is disposed between the X-ray source 11 and the first grid 21, but the subject H is disposed between the first grid 21 and the second grid 22. You may arrange. In this case as well, a phase differential image and a phase contrast image can be similarly generated.

In the above embodiment, the phase contrast image is displayed on the monitor, but the corrected phase differential image generated by the correction processing unit 33 may be displayed on the monitor.

In the above embodiment, the second grid 22 and the X-ray image detector 20 are provided separately. However, instead of the X-ray image detector 20, JP-A-2009-133823 (US Pat. No. 7,746,811). The second grid 22 can be omitted by using the X-ray image detector disclosed in the specification.

This X-ray image detector is a direct conversion type X-ray image detector including a conversion layer that converts X-rays into electric charges and a charge collection electrode that collects electric charges converted in the conversion layers. The charge collection electrode of each pixel is configured by arranging a plurality of linear electrode groups formed by electrically connecting linear electrodes arranged at a constant period so that their phases are different from each other. ing. In this case, the charge collection electrode constitutes an intensity modulation unit.

In the above embodiment, the phase differential image is generated based on the fringe scanning method. However, the phase differential image is generated based on the Fourier transform method described in International Publication WO2010 / 050833 (US Pat. No. 8,0097,979). It may be generated. This Fourier transform method obtains a Fourier spectrum of moire fringes generated in image data by Fourier transforming single image data obtained by an X-ray image detector, and obtains a spectrum corresponding to the carrier frequency from this Fourier spectrum. This is a method of generating a phase differential image by performing inverse Fourier transform after separation. In this case, since it is not necessary to move the first and

second grids

21 and 22, the scanning mechanism 23 can be omitted.

Further, there is a method in which the scanning mechanism 23 is omitted and a phase differential image is generated based on single image data obtained by the X-ray image detector 20 via the first and

second grids

21 and 22. As this method, there is a pixel division method filed by the present applicant as Japanese Patent Application No. 2010-256241. In this pixel division method, the first grid 21 and the second grid 22 are slightly rotated around the Z direction, and moire fringes having a period in the Y direction are generated in the G2 image. The single image data obtained by the X-ray image detector 20 is divided into groups of pixel rows (pixels arranged in the X direction) having different phases with respect to the moire fringes, and the plurality of divided image data are A phase differential image is generated in the same procedure as the above-described fringe scanning method, assuming that the images are based on a plurality of different G2 images by fringe scanning. In this pixel division method, the intensity modulation signal described above is expressed as a change in intensity of pixel values for one cycle of moire fringes generated in single image data.

The above embodiments may be combined with each other within a consistent range. The present invention is applicable not only to a radiographic imaging system for medical diagnosis but also to other radiographic systems for industrial use and nondestructive inspection. The present invention can also be applied to a radiographic imaging system that uses gamma rays or the like in addition to X-rays.

DESCRIPTION OF SYMBOLS 10 X-ray imaging system 11 X-ray source 21 1st grid 21a X-ray absorption part 21b X-ray low absorption part 22 2nd grid 22a X-ray absorption part 22b X-ray low absorption part

Claims

A radiation source that emits radiation at the time of pre-imaging without arranging the subject and at the main imaging with the subject arranged;
A first grid that passes the radiation to generate a first periodic pattern image;
An intensity modulation unit for applying intensity modulation to the first periodic pattern image to generate a second periodic pattern image;
A radiation image detector for detecting the second periodic pattern image and generating image data;
A phase differential image generation unit that generates a phase differential image based on the image data generated during the main photographing;
A correction data generation unit that generates correction data based on the image data generated during the pre-shooting;
A correction processing unit for subtracting the correction data from the phase differential image;
Based on a second radiation dose obtained by multiplying the first radiation dose at the time of the main imaging by α times (where α> 1), an exposure condition at the time of the pre-imaging is determined. A control unit for executing radiation;
A radiation imaging system comprising:
The control unit determines whether or not pixel data is saturated when the first radiation dose is multiplied by α. If the pixel data is saturated, the control unit increases the first radiation dose by α / ν (however, ν (2 is a positive number of 2 or more) is defined as the second radiation dose, and the radiation source and the radiation image detector are caused to execute the radiation and the image detection at least ν times during the pre-imaging. The radiation imaging system according to claim 1.
The correction data generation unit performs addition processing on each piece of image data generated by the radiation image detector when the radiation and the image detection are executed at least ν times during the pre-imaging. The radiation imaging system according to item 2 of the range.
The intensity modulation unit generates intensity second modulation pattern at a plurality of scanning positions having different phases with respect to the first period pattern image to generate a plurality of second period pattern images,
The radiation image detector detects the plurality of second periodic pattern images and generates a plurality of image data,
The phase differential image generation unit and the correction data generation unit calculate the phase differential image and the phase differential image by calculating a phase shift amount of an intensity modulation signal representing an intensity change of pixel data with respect to the scanning position based on the plurality of image data. Generating each of the correction data;
In the case where the value obtained by multiplying the first radiation dose by α / ν is used as the second radiation dose, the control unit performs the radiation and the image detection at ν times or more at each scanning position during the pre-imaging. The radiation imaging system according to any one of claims 1 to 3, wherein the radiation imaging system is executed.
The correction data generation means adds ν or more image data obtained at each scanning position when the radiation and the image detection are performed at ν times or more at each scanning position during pre-imaging. The radiation imaging system according to claim 4, wherein:
It further includes an operation unit that enables selection of the type of imaging region,
The radiographic system according to any one of claims 1 to 5, wherein the control unit determines the first radiation dose according to an imaging region selected by the operation unit. .
The said control part determines the said 1st radiation dose based on the thickest imaging | photography part with the largest thickness among the some imaging | photography parts selectable by the said operation part. The radiation imaging system according to any one of claims 5 to 5.
A photo timer,
The radiographic system according to any one of claims 1 to 5, wherein the control unit sets the set value of the phototimer as the first radiation dose.
The radiation imaging system according to any one of claims 1 to 8, wherein α is 100 or more.
The intensity modulation unit includes a second grid having a periodic pattern in the same direction as the first periodic pattern image, and a scanning mechanism that moves one of the first and second grids at a predetermined pitch. The radiation imaging system according to any one of claims 4 to 9, wherein:
The radiation imaging system according to claim 10, wherein the first grid is an absorption grid, and projects incident radiation onto the second grid geometrically.
11. The first grid according to claim 10, wherein the first grid is an absorption grid or a phase grid, and causes a Talbot interference effect on incident radiation to be projected onto the second grid. Radiography system.
A radiation source that emits radiation at the time of pre-imaging without arranging the subject and at the main imaging with the subject arranged;
A first grid that passes the radiation to generate a first periodic pattern image;
An intensity modulation unit for applying intensity modulation to the first periodic pattern image to generate a second periodic pattern image;
A radiation image detector for detecting the second periodic pattern image and generating image data;
A phase differential image generation unit that generates a phase differential image based on the image data generated during the main photographing;
A correction data generation unit that generates correction data based on the image data generated during the pre-shooting;
A correction processing unit for subtracting the correction data from the phase differential image;
A method for controlling a radiation imaging system comprising:
A radiation imaging system control method, wherein an exposure condition for the pre-imaging is determined based on a second radiation dose obtained by multiplying the first radiation dose during the main imaging by α times (where α> 1). .