JP2011206113A - Radiographic imaging system - Google Patents

Abstract

PROBLEM TO BE SOLVED: To reduce the manufacturing cost and to improve the quality of a phase contrast image.SOLUTION: The radiographic imaging system includes an X-ray source, an absorbent grid transmitting X rays to generate a striped image (a G1 image), and an optical reading type X-ray image detector (an FPD). The FPD records the G1 image as an electrostatic latent image, and by scanning the recorded electrostatic latent image with reading light in the direction (an x direction) of a periodic pattern of the G1 image at a shorter pitch than a period P, the electrostatic latent image is read as image data. A sampling processing part generates an image group consisting of a plurality of intensity-modulated images by sampling the image data at a plurality of relative positions (k=0-3) in the same period as the periodic pattern of the G1 image with the different phases from the periodic pattern. A phase differential image is generated by computing the quantity of displacement of phases of intensity modulation signals obtained from the image group by a subject, and the phase contrast image is obtained by integrating the phase differential images in the x direction.

Description

  The present invention relates to a radiation imaging system that images a subject with radiation such as X-rays, and more particularly to a radiation imaging system that performs X-ray phase imaging using a grating.

  X-rays are used as a probe for seeing through the inside of a subject because they have characteristics such as attenuation depending on the atomic numbers of elements constituting the substance and the density and thickness of the substance. X-ray imaging is widely used in fields such as medical diagnosis and non-destructive inspection.

  In a general X-ray imaging system, a subject is placed between an X-ray source that emits X-rays and an X-ray image detector that detects X-rays, and a transmission image of the subject is taken. In this case, the X-rays emitted from the X-ray source toward the X-ray image detector correspond to the difference in the characteristics (atomic number, density, thickness) of the substances existing on the path to the X-ray image detector. After receiving a certain amount of attenuation (absorption), it enters each pixel of the X-ray image detector. As a result, an X-ray absorption image of the subject is detected and imaged by the X-ray image detector. As an X-ray image detector, in addition to a combination of an X-ray intensifying screen and a film and a stimulable phosphor, a flat panel detector (FPD) using a semiconductor circuit is widely used.

  However, since the X-ray absorption ability is lower as a substance composed of an element having a smaller atomic number, a sufficient softness or contrast (contrast) of an X-ray absorption image cannot be obtained in a soft body tissue or soft material. There is. For example, most of the components of the cartilage part constituting the joint of the human body and the joint fluid in the vicinity thereof are water, and there is little difference in the amount of X-ray absorption between them, so that it is difficult to obtain a difference in light and shade.

  Against this background, in recent years, instead of X-ray intensity changes by the subject, an image based on the X-ray phase change (angle change) by the subject (hereinafter referred to as a phase contrast image) is obtained. Research on line phase imaging has been actively conducted. In general, when X-rays are incident on an object, the phase exhibits a higher interaction than the intensity of the X-rays. Therefore, X-ray phase imaging using a phase difference is a weakly absorbing object with low X-ray absorption capability. Can also obtain a high-contrast image. As a kind of such X-ray phase imaging, an X-ray imaging system using an X-ray Talbot interferometer composed of two transmission diffraction gratings and an X-ray image detector has been devised (for example, Patent Document 1). reference).

  The X-ray Talbot interferometer has a first diffraction grating disposed behind the subject, and a second diffraction grating downstream by a specific distance (Talbot interference distance) determined by the grating pitch of the first diffraction grating and the X-ray wavelength. Are arranged, and an X-ray image detector is arranged behind them. The Talbot interference distance is a distance at which X-rays that have passed through the first diffraction grating form a self-image (stripe image) due to the Talbot interference effect. This self-image is the same as the X-ray source and the first diffraction grating. Are modulated by the interaction (phase change) between the subject and the X-rays arranged between the two.

  In the X-ray Talbot interferometer, the phase contrast image of the subject is obtained from the change (phase shift) caused by the subject of the fringe image whose intensity is modulated by superimposing the self-image of the first diffraction grating and the second diffraction grating. To do. In the fringe scanning method, the second diffraction grating is substantially parallel to the surface of the first diffraction grating with respect to the first diffraction grating and substantially in the grating direction (strip direction) of the first diffraction grating. In the vertical direction, a plurality of times of imaging are performed while translational movement (scanning) is performed at a scanning pitch obtained by equally dividing the lattice pitch, and the phase of intensity change with respect to the scanning of the pixel data of each pixel obtained by the X-ray image detector is detected. This is a method of obtaining a phase differential image (corresponding to the angular distribution of X-rays refracted by the subject) from the amount of deviation (phase deviation amount with and without the subject H). The phase contrast image of the subject is obtained by integrating along the fringe scanning direction. Since the intensity of the pixel data is periodically modulated with respect to the scan, the set of pixel data for the scan is hereinafter referred to as an “intensity modulation signal”.

  However, in an X-ray imaging system using an X-ray Talbot interferometer, it is necessary to provide two gratings, a first diffraction grating and a second diffraction grating, and relatively translate these two gratings. Since the scanning mechanism to be used is necessary, there is a problem that the cost is increased. In addition, since the X-ray imaging system performs imaging at each scanning position moved by the scanning mechanism, X-ray exposure is performed for the number of scanning steps, resulting in a large exposure dose of the subject. There is.

  Therefore, the present applicant has proposed to configure an X-ray image detector so that a plurality of fringe images having different phases can be simultaneously detected by one imaging (see Patent Document 2). As a result, the second diffraction grating and the scanning mechanism can be eliminated, so that the cost can be reduced, and since only one imaging is required, the exposure dose can be reduced.

  Specifically, the X-ray image detector includes a plurality of linear electrode groups having different phases with respect to the grating pattern of the first diffraction grating in one pixel, so that a plurality of phases having different phases from each other are provided. A fringe image can be acquired, and the intensity-modulated signal can be obtained by one shooting. Each linear electrode group is a charge collection electrode that collects charges generated in the X-ray conversion layer.

JP 2008-200320 A JP 2009-133823 A

  However, in the X-ray image detector described in Patent Document 2, in order to arrange a plurality of linear electrode groups in one pixel, it is necessary to finely refine the linear electrodes. In addition, this miniaturization may cause a decrease in manufacturing yield due to dust or the like, resulting in an increase in manufacturing cost.

  Furthermore, if the linear electrodes are miniaturized as described above, the amount of collected charge per linear electrode is reduced, and therefore, there is a problem that it is easily affected by noise and the image quality of the phase contrast image is reduced.

  The present invention has been made in view of the above problems, and in a radiographic system that enables phase imaging using one grating and one radiographic image detector, a reduction in manufacturing cost and phase contrast images can be obtained. An object of the present invention is to provide a radiation imaging system that can achieve high image quality.

  In order to achieve the above object, a radiation imaging system of the present invention records a radiation source that emits radiation, a grating that passes the radiation to generate a fringe image, and the fringe image as an electrostatic latent image, Scanning the electrostatic latent image recorded in the periodic pattern direction of the fringe image with scanning light at a pitch shorter than the period, thereby reading the electrostatic latent image as image data; Sampling processing for generating an image group composed of a plurality of intensity-modulated images by sampling the image data at a plurality of relative positions having the same period as the periodic pattern and having different phases with respect to the periodic pattern. And a phase of an object disposed between the radiation source and the radiation image detector based on the image group generated by the sampling processing unit. Image processing means for generating contrast images, characterized by comprising a.

  In addition, the image processing unit calculates a phase shift amount of an intensity modulation signal for each pixel based on the image group generated by the sampling processing unit, and generates a phase differential image; and Preferably, the phase differential image includes phase contrast image generation means for generating a phase contrast image by integrating the phase differential image in the periodic pattern direction.

  Moreover, it is preferable that the said grating | lattice is an absorption-type grating | lattice, and projects the radiation from the said radiation source on a radiographic image detector as a fringe image.

  The grating is preferably a phase-type grating and projects radiation from the radiation source as a fringe image onto a radiation image detector by a Talbot interference effect.

  The radiation image detector includes a first electrode layer that transmits radiation, a recording photoconductive layer that generates a charge pair by incidence of radiation transmitted through the first electrode layer 30, and a recording photoconductive layer. The charge transport layer that acts as an insulator for the charge of one polarity of the charge pair generated in step 2 and the charge transport layer that acts as a conductor for the charge of the other polarity, A reading photoconductive layer that generates a pair, a second electrode layer having a plurality of linear electrodes, and extending in a direction orthogonal to the linear electrodes, and driven in the extending direction of the linear electrodes, A reading line light source that irradiates the reading photoconductive layer to the reading photoconductive layer through the second electrode layer, and the electrostatic latent image is formed at an interface between the recording photoconductive layer and the charge transport layer. It is preferable that a power storage unit for storing

  The linear electrode is connected to a charge amplifier for reading out charges, and includes a transparent linear electrode that transmits the reading light, and a light-shielding linear electrode that is provided with a ground potential and shields the reading light. It is preferable.

  The radiation imaging system of the present invention uses a light reading type radiation image detector and performs the above-described fringe scanning by scanning the reading light, so there is no need to make the linear electrodes finer as in the prior art, It is possible to reduce the manufacturing cost and improve the image quality of the phase contrast image.

1 is a schematic diagram illustrating a configuration of an X-ray imaging system according to a first embodiment of the present invention. It is a schematic diagram which shows the structure of a flat panel detector (FPD). It is a schematic diagram for demonstrating imaging | photography operation | movement of FPD. It is a schematic diagram for demonstrating the reading process of FPD. It is a schematic diagram for demonstrating the fringe scanning method. It is a schematic diagram for demonstrating the scanning method of a reading line light source. It is a schematic diagram for demonstrating the effect | action of a sampling process part. It is a graph for demonstrating the phase shift of an intensity | strength modulation signal. It is a schematic diagram for demonstrating the integration area in the case of integrating pixel data. It is a block diagram which shows the modification of a sampling process circuit.

(First embodiment)
In FIG. 1, an X-ray imaging system 10 according to the first embodiment of the present invention is disposed so as to face an X-ray source 11 that irradiates a subject H with X-rays, and is opposed to the X-ray source 11. An imaging unit 12 that detects X-rays transmitted through the specimen H and generates image data, a memory 13 that stores image data read from the imaging unit 12, and a phase contrast based on the image data stored in the memory 13 An image processing unit 14 that generates an image, an image recording unit 15 that records a phase contrast image generated by the image processing unit 14, an imaging control unit 16 that controls the X-ray source 11 and the imaging unit 12, and an operation And a system control unit 18 that comprehensively controls the entire X-ray imaging system 10 based on an operation signal input from the console 17.

  The X-ray source 11 includes a high voltage generator, an X-ray tube, a collimator (all not shown), and the like, and irradiates the subject H with X-rays based on the control of the imaging control unit 16. For example, the X-ray tube is of a rotating anode type, emits an electron beam from a filament in accordance with a voltage from a high voltage generator, and collides the electron beam with a rotating anode rotating at a predetermined speed, thereby causing an X-ray. Is generated. The rotating anode rotates in order to reduce deterioration due to the electron beam continuing to hit the fixed position, and the collision part of the electron beam becomes an X-ray focal point that emits X-rays. The collimator limits the irradiation field so as to shield a portion of the X-ray emitted from the X-ray tube that does not contribute to the examination region of the subject H.

  The imaging unit 12 is provided with a flat panel detector (FPD) 20 and an absorption grating 21 for detecting phase change (angle change) of X-rays by the subject H and performing phase imaging. The FPD 20 is arranged so that the detection surface is orthogonal to a direction along the optical axis A of X-rays irradiated from the X-ray source 11 (hereinafter referred to as z direction).

The absorptive lattice 21 includes a plurality of X-ray shielding portions 21a extending in one direction in the plane orthogonal to the z direction (hereinafter referred to as the y direction) and a direction orthogonal to the z direction and the y direction (hereinafter referred to as the x direction). ) to those arranged at a predetermined pitch p _1. As a material of the X-ray shielding part 21a, a metal excellent in X-ray absorption is preferable, for example, gold or lead is preferable.

  The image processing unit 14 performs sampling processing on image data stored in the memory 13 to generate a plurality of images, and generates a phase differential image based on the plurality of images obtained by the sampling processing unit 22. A phase differential image generation unit 23 and a phase contrast image generation unit 24 that generates a phase contrast image by integrating the phase differential image generated by the phase differential image generation unit 23 along the x direction are provided. The phase contrast image generated by the phase contrast image generation unit 24 is recorded in the image recording unit 15 and then output to the console 17 and displayed on a monitor (not shown).

  In addition to the monitor, the console 17 includes an input device (not shown) through which an operator inputs a shooting instruction and the content of the instruction. As this input device, for example, a switch, a touch panel, a mouse, a keyboard or the like is used, and an X-ray imaging condition such as a tube voltage of the X-ray tube or an X-ray irradiation time, an imaging timing, etc. are input by operation of the input device. The The monitor is composed of a liquid crystal display or a CRT display, and displays characters such as X-ray imaging conditions and the phase contrast image.

  In FIG. 2, an FPD 20 is an optical reading type X-ray image detector, and a charge pair (electronic -The photoconductive layer for recording 31 that generates (hole pairs) and the charge of one polarity among the charge pairs generated in the photoconductive layer for recording 31 act as an insulator and the charge of the other polarity , A charge transport layer 32 that acts as a conductor, a read photoconductive layer 33 that generates charge pairs (electron-hole pairs) upon irradiation with the read light LT, and a second electrode layer 34. And a glass substrate 35 and a readout line light source 36 for generating linear readout light LT. At the interface between the recording photoconductive layer 31 and the charge transport layer 32, a power storage unit 37 is formed that accumulates one of the charge pairs generated in the recording photoconductive layer 31. Each of the layers is formed on the glass substrate 35 in order from the second electrode layer 34.

The first electrode layer 30 is made of gold (Au) having a thickness of about 100 nm. As the first electrode layer 30, Al or the like can be used instead of Au, and further, Nesa film (SnO ₂ ), ITO (Indium Tin Oxide), IZO (Indium Zinc Oxide), amorphous light It is also possible to use IDIXO (Idemitsu Indium X-metal Oxide; Idemitsu Kosan Co., Ltd.) which is a permeable oxide film.

  The recording photoconductive layer 31 is made of amorphous selenium (a-Se) having a thickness of about 10 μm to 1500 μm. Amorphous selenium has properties such as relatively high quantum efficiency for X-rays and high dark resistance.

The charge transport layer 32, and the mobility of the charge charged on the first electrode layer 30 during the X-ray imaging, a large difference in mobility of the charge and vice versa polarity (e.g. 10 ² or more, preferably 10 ³ or more) is preferable. For example, poly N-vinylcarbazole (PVK), N, N′-diphenyl-N, N′-bis (3-methylphenyl)-[1,1′-biphenyl] -4,4′-diamine (TPD) An organic compound such as a discotic liquid crystal, a TPD polymer (polycarbonate, polystyrene, PVK) dispersion, a semiconductor material such as a-Se or As ₂ Se ₃ doped with 10 ppm to 200 ppm of Cl, and a thickness of 0. About 2 μm to 2 μm is appropriate.

  The reading photoconductive layer 33 is made of a-Se having a thickness of about 5 μm to 20 μm. In addition, as the photoconductive layer 33 for reading, instead of a-Se, Se-Te, Se-As-Te, metal-free phthalocyanine, metal phthalocyanine, MgPc (Magnesium phtalocyanine), VoPc (phase II of Vanadyl phthalocyanine), CuPc (Cupper phtalocyanine) or the like can be used.

  The second electrode layer 34 includes a plurality of transparent linear electrodes 34a that transmit the reading light LT, and a plurality of light shielding linear electrodes 34b that shield the reading light LT. The transparent linear electrodes 34a and the light-shielding linear electrodes 34b extend from one end of the detection surface to the other end in the x direction, and are alternately arranged in parallel at a predetermined interval in the y direction.

  The transparent linear electrode 34a is made of a material that transmits the reading light LT and has conductivity. The transparent linear electrode 34a is made of, for example, ITO, IZO or IDIXO having a thickness of about 100 nm to 200 nm.

  The light shielding linear electrode 34b is formed of a material that shields the reading light LT and has conductivity. The light shielding linear electrode 34b transmits erasing light for erasing charges remaining in the power storage unit 37 after a reading process described later. For this reason, the light-shielding linear electrode 34b is preferably made of a combination of a color filter that shields the reading light LT and transmits erasing light, and the transparent conductive material. The thickness of the transparent conductive material is about 100 nm to 200 nm.

  The readout line light source 36 extends from one end of the detection surface to the other end in the y direction, and is orthogonal to all the transparent linear electrodes 34a. The reading line light source 36 is disposed below the glass substrate 35 and irradiates the reading photoconductive layer 33 with the reading light LT via the glass substrate 35 and the second electrode layer 34 during a reading process described later. Further, the readout line light source 36 is configured to translate in the x direction by a moving mechanism (not shown), and irradiation and translational movement of the reading light LT are controlled by the above-described imaging control unit 16.

  In addition, the FPD 20 is provided with a power supply circuit 38 that applies a predetermined negative voltage or ground potential to the first electrode layer 30 under the control of the imaging control unit 16. As will be described in detail later, the power supply circuit 38 supplies a negative voltage to the first electrode layer 30 during recording of the X-ray image, and supplies a ground potential to the first electrode layer 30 during reading processing.

  Further, the FPD 20 is provided with a charge amplifier 39 for reading out charges generated in the reading photoconductive layer 33 during reading processing. The charge amplifier 39 is connected to each transparent linear electrode 34a. One light shielding linear electrode 34b is all grounded. The charge amplifier 39 is a circuit for accumulating charges, and includes a reset switch 39a for resetting the accumulated charges. The imaging switch 16 controls the reset switch 39a.

  The reading line light source 36 is moved (scanned) by a predetermined pitch in the x direction at the time of reading processing, and irradiates the reading light LT at each moved position (hereinafter referred to as a scanning position), and according to the reading light LT. Then, the charge generated in the read photoconductive layer 33 is read by the charge amplifier 39. The pitch for moving the readout line light source 36 (hereinafter referred to as scanning pitch) is a pattern period in the x direction of the projected image of the absorption grating 21 projected onto the detection surface of the FPD 20 (the surface 30a of the first electrode layer 30). Short enough than. The arrangement pitch in the y direction of the transparent linear electrodes 34a and the light shielding linear electrodes 34b is not particularly limited, but is preferably about the same as the pattern period.

  Next, the photographing operation of the FPD 20 at each scanning position will be described. As shown in FIG. 3A, the X-rays irradiated to the FPD 20 are transmitted to the recording photoconductive layer 31 through the first electrode layer 30, and the recording photoconductive layer 31 is irradiated with X-rays. Charge pairs are generated. At this time, a negative voltage is applied to the first electrode layer 30 by the power supply circuit 38, and a negative charge is charged. Of the charge pair generated in the recording photoconductive layer 31, the positive charge is combined with the negative charge of the first electrode layer 30 and disappears. One negative charge is accumulated in the power storage unit 37 as a latent image charge (electrostatic latent image) as shown in FIG.

  Next, as shown in FIG. 4, reading light LT is emitted from the reading line light source 36 in a state where the first electrode layer 30 is grounded by the power supply circuit 38. The reading light LT passes through the transparent linear electrode 34 a and is applied to the reading photoconductive layer 33. Among the charge pairs generated in the reading photoconductive layer 33 by the reading light LT, the positive charge is combined with the latent image charge of the power storage unit 37, and the negative charge is passed through the charge amplifier 39 through the light shielding linear electrode 34b. It binds to the positive charge that is charged.

  A current flows through the charge amplifier 39 due to the combination of the negative charge generated in the reading photoconductive layer 33 and the positive charge charged in the light shielding linear electrode 34b, and this current is integrated and output as a pixel signal. The The pixel signal output from the charge amplifier 39 is digitized by an A / D converter (not shown), subjected to offset correction, gain correction, linearity correction, and the like by a correction circuit (not shown), and sequentially input to the memory 13. Is done.

In FIG. 5, the X-ray shielding portions 21 a of the absorption type grating 21 are arranged at a predetermined pitch p ₁ in the x direction with a predetermined distance d ₁ therebetween. The X-ray shielding part 21a is disposed on an X-ray transmissive substrate (not shown) (for example, a glass substrate). The absorption type grating 21 does not give a phase difference to incident X-rays but gives an intensity difference, and is also called an amplitude type grating. Incidentally, (a region of the interval d _1, d ₂₎ slit portion may not be a void, X-rays low absorption material such as a polymer or light metal may be filled.

The absorptive grating 21 is configured to linearly project X-rays that have passed through the slit portion regardless of the presence or absence of the Talbot interference effect. Specifically, by setting the interval d ₁ to a value sufficiently larger than the peak wavelength of the X-rays emitted from the X-ray source 11, most of the X-rays included in the irradiated X-rays are not diffracted by the slit portion. In addition, it is configured to pass 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, if the distance d1 is about ₁ μm to 10 μm, most of the X-rays are projected linearly without being diffracted by the slit portion. In this case, the grating pitch p ₁ is about 2 μm to 20 μm.

The X-ray irradiated from the X-ray source 11 is not a parallel beam but a cone beam with the X-ray focal point 11a as a light emission point, and thus a projected image projected through the absorption grating 21 onto the detection surface of the FPD 20 (Hereinafter, this projection image is referred to as a G1 image or a fringe image) expands in proportion to the distance from the X-ray focal point 11a. The pitch p ₂ of the periodic pattern of the G1 image is geometric when the distance from the X-ray focal point 11a to the absorption grating 21 is L ₁ and the distance from the absorption grating 21 to the detection surface of the FPD ₂₀ is L _2. Based on this relationship, it is expressed by the following formula (1).

Further, the width d _{2 in} the x direction of the bright portion of the G1 image corresponding to the interval d ₁ of the absorption type grating 21 is expressed by the following equation (2).

In the case of a Talbot interferometer, the distance L ₂ from the absorption grating 21 to the detection surface of the FPD ₂₀ is limited to the Talbot interference distance determined by the grating pitch of the diffraction grating and the X-ray wavelength. in part 12, the absorption type grating 21 be one to project without diffracting the incident X-ray, an image having passed through the absorption grating 21, to similarly enlarged, the distance L _2, the Talbot distance Can be set independently of

As described above, the imaging unit 12 of the present embodiment does not constitute a Talbot interferometer, but Talbot interference is assumed when X-ray diffraction occurs in the absorption grating 21 and a Talbot interference effect is generated. The distance Z is expressed by the following formula (3) using the grating pitch p ₁ of the absorption grating 21, the X-ray wavelength (peak wavelength) λ, and a positive integer m.

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

  The X-ray shielding part 21a preferably shields (absorbs) X-rays completely in order to generate a periodic pattern image with high contrast, but is made of a material having excellent X-ray absorption (such as gold or lead). Even if is used, there are not a few X-rays that are transmitted without being absorbed. For this reason, in order to improve the X-ray shielding property, it is preferable to increase the thickness of the X-ray shielding part 21a (the thickness in the z direction) as much as possible (that is, to increase the aspect ratio). For example, when the tube voltage of the X-ray tube is 50 kV, it is preferable to shield 90% or more of the irradiated X-rays. In this case, the thickness of the X-ray shielding part 21a is 30 μm or more in terms of gold (Au). It is preferable that

Read line light source 36 of the FPD 20, as shown in FIG. 6, with respect to G1 image projected on the detection surface of the FPD 20, it is scanned in a short scan pitch than the pattern period _{p 2} of the G1 image. Therefore, the above photographing operation is performed at each scanning position, and a pixel signal is acquired. Scanning pitch, for example, a quarter of the pattern period p _2, 1 4 single scanning position with respect to the pattern period p ₂ is allocated, the pixel signal is obtained at each scanning position. These pixel signals are subjected to the above-described digitization and correction processes, and are input to the memory 13 as image data.

Multiple sampling processing unit 22 of the image processing unit 14, as shown in FIG. 7, the image data stored in the memory 13, in the same sampling cycle and pattern pitch p ₂ of the G1 image, while shifting the phase in the x direction By performing the sampling once, four images (image group) intensity-modulated at a plurality of relative positions (k = 0, 1, 2, 3) having different phases with respect to the G1 image are generated.

Specifically, the sampling processing unit 22 generates the image group by performing the following first to fourth sampling processes. Here, a pixel signal corresponding to the pattern period _{p 2} of the G1 image I1,12, expressed as I3,14. In the first sampling process, the sampling processing unit 22 samples the pixel signals I1 and I2 and adds them one by one to generate pixel data D0, thereby generating an image of k = 0. In the second sampling process, pixel signals I2 and I3 are sampled by shifting the sampling position by ¼ period (phase difference “0.5π”), and these sets are added one by one to obtain pixel data D1. An image with k = 1 is generated. Further, in the third sampling process, the pixel signals I3 and I4 are sampled by shifting the sampling position by ½ period (phase difference “π”), and these sets are added one by one to obtain pixel data D2. An image with k = 2 is generated. In the fourth sampling process, the pixel signals I4 and I1 are sampled by shifting the sampling position by 3/4 period (phase difference “1.5π”), and these sets are added one by one to obtain pixel data D3. As a result, an image of k = 3 is generated.

By the above sampling process, an intensity modulation signal composed of each combination of the pixel data D0 to D3 is obtained. The intensity-modulated signal, place the absorptive grating as in the prior art have the same pattern period and the pattern period p ₂ of the G1 image immediately before the detection surface of the FPD 20, x-direction to a relative to the G1 image This is equivalent to the intensity modulation signal obtained by scanning. The pattern period p ₂ in G1 image, due to manufacturing errors or arrangement errors, so the above sampling cycle and some differences may occur, when a difference occurs, moire fringes occur in the intensity-modulated image. It is preferred that ideally no moire fringes occur, the moire fringes can be utilized to verify the deviation between the pattern period p ₂ and the sampling period.

  When the subject H is disposed between the X-ray source 11 and the absorption grating 21, the image data obtained by the FPD 20 is modulated by the subject H. This modulation amount is proportional to the angle of the X-ray deflected by the refraction effect by the subject H, and the phase contrast image of the subject H is generated by analyzing the image group obtained by the sampling process. Can do.

  Next, a method for analyzing the image group will be described. FIG. 5 illustrates one X-ray refracted according to the phase shift distribution Φ (x) in the x direction of the subject H. Reference numeral 40 indicates an X-ray path that goes straight when the subject H does not exist. Reference numeral 41 denotes an X-ray path that is refracted and deflected by the subject H when the subject H exists.

  The phase shift distribution Φ (x) of the subject H is expressed by the following equation (5), where n (x, z) is the refractive index distribution of the subject H and z is the direction in which the X-ray travels.

  The G1 image projected from the first absorption type grating 21 onto the detection surface of the FPD 20 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 X-ray refraction angle φ 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) of the subject H.

  Thus, the displacement amount Δx of the G1 image due to X-ray refraction at the subject H is related to the phase shift distribution Φ (x) of the subject H. The amount of displacement Δx is the amount of phase shift ψ of each intensity modulation signal generated by the sampling processing unit 22 (the amount of phase shift of the intensity modulation signal with and without the subject H), and the following: It is related as shown in Equation (8).

  Therefore, by obtaining the phase shift amount ψ of the intensity modulation signal, the refraction angle φ is obtained from the equation (8), and the differential amount of the phase shift distribution Φ (x) is obtained using the equation (7). Is integrated, a phase shift distribution Φ (x) of the subject H, that is, a phase contrast image of the subject H can be generated. In the present embodiment, the phase shift amount ψ is calculated using the fringe scanning method shown below.

In the fringe scanning method, in general, the G1 image is obtained by a lattice pattern equivalent to the G1 image while shifting the phase by a pitch (p ₂ / M) obtained by dividing the pattern period p ₂ of the G1 image into M (integers of 2 or more). Apply intensity modulation to the image. In the present embodiment, as shown in FIG. 7, M = 4, and the sampling processing unit 22 performs periodic sampling processing on the high resolution image data obtained as the G1 image, thereby virtually Intensity modulation is performed.

  In the case of M = 4, at the position of k = 0, mainly the X-ray component that is not refracted by the subject H in the G1 image is sampled. In contrast, at the position of k = 1, the X-ray component that is not refracted by the subject H decreases, while the X-ray component refracted by the subject H increases.

  Next, at the position of k = 2, only the part corresponding to the X-ray refracted by the subject H is mainly sampled in the G1 image. In contrast, at the position of k = 3, the X-ray component refracted by the subject H decreases while the X-ray component not refracted by the subject H increases.

Hereinafter, in general, the pattern period p ₂ of the G1 image illustrating a method of calculating the phase shift amount ψ of the intensity modulation signal in the case of dividing into M. When the intensity modulation signal at the sampling position k is denoted as I _k (x), I _k (x) is expressed by the following equation (9).

Here, x is the coordinate in the x-direction, A ₀ corresponds to the intensity of the incident X-ray, A _n is a value corresponding to the contrast of the intensity-modulated signal (here, n represents is a positive integer ). Φ (x) represents the refraction angle φ as a function of the coordinate x.

  Next, using the relational expression of the following expression (10), the refraction angle φ (x) is expressed as the expression (11).

Here, arg [] means extraction of the declination and corresponds to the phase shift amount ψ. Therefore, the refraction angle φ (x) is obtained by calculating the phase shift amount ψ from the actual data of the intensity modulation signal I _k (x) generated by the sampling processing unit 22 based on the equation (11), and the phase is obtained. A differential amount of the shift distribution Φ (x) is obtained.

Specifically, the intensity modulation signal I _k generated by the sampling processing unit 22 _(x), as shown in FIG. 8, the sampling positions k, varies with a period of the pattern cycle p ₂ of the G1 image. The broken line in the figure shows the change of the pixel data when the subject H does not exist, and the solid line in the figure shows the change of the pixel data when the subject H exists. The phase difference between the two waveforms corresponds to the phase shift amount ψ.

  In the above description, the y coordinate in the y direction is not taken into consideration, but by performing the same calculation on the y coordinate, a two-dimensional phase shift distribution ψ (x, y) in the x direction and the y direction is obtained. It is done. This phase shift distribution ψ (x, y) corresponds to the phase differential image.

  The phase differential image generation process is performed by the phase differential image generation unit 23, and the generated phase differential image is input to the phase contrast image generation unit 24. The phase contrast image generation unit 24 integrates the input phase differential image along the x axis to generate a phase shift distribution Φ (x, y) of the subject H, and records this as a phase contrast image. Input to unit 15.

  In the X-ray imaging system 10 configured as described above, when the subject H is placed between the X-ray source 11 and the imaging unit 12, an X-ray source is input from the console 17 by the operator. The X-ray irradiation is performed once from 11 to the subject H, and the fringe image (G1 image) generated by passing through the subject H and passing through the absorption grating 21 is detected by the FPD 20. A pixel signal is read from the FPD 20 in accordance with the scanning of the read line light source 36, and each read pixel signal is stored in the memory 13 as image data.

Image data stored in the memory 13, the sampling processor 22, the pattern period p ₂ of the G1 image multiple times sampling is performed while shifting the phase in the x direction, is intensity modulated at a plurality of relative positions with different phases A plurality of images (image group) are generated. Based on this image group, a phase differential image is generated by the phase differential image generation unit 23, and a phase contrast image is generated by the phase contrast image generation unit 24. The phase contrast image is recorded on the image recording unit 15 and then displayed on the monitor.

  In the above embodiment, when the distance from the X-ray source 11 to the FPD 20 is increased, the blur of the G1 image due to the focus size of the X-ray focal point 11a (generally about 0.1 mm to 1 mm) affects the phase. Since there is a risk of degrading the image quality of the contrast image, a multi slit (ray source grid) may be arranged immediately after the X-ray focal point 11a.

  The multi-slit is an absorption type grating having a configuration similar to that of the absorption type grating 21, and a plurality of X-ray shielding parts extending in one direction (in this embodiment, the y direction) are X-ray shieldings of the absorption type grating 21. They are periodically arranged in the same direction as the portion 21a (in this embodiment, the x direction). This multi-slit partially shields X-rays from the X-ray source 11 to reduce the effective focal size in the x direction, and forms a large number of point light sources (dispersed light sources) in the x direction. Suppresses the blur of the G1 image.

Moreover, in the said embodiment, although the absorption-type grating | lattice 21 is comprised so that the X-ray which passed the slit part may be projected linearly, this invention is not limited to this structure, A slit part The X-ray diffraction may cause a so-called Talbot interference effect (configuration described in International Publication WO 2004/058070). However, in this case, it is necessary to set the distance L2 between the absorption grating 21 and the detection surface of the FPD ₂₀ as the Talbot interference distance. In this case, a phase-type grating (phase-type diffraction grating) can be used instead of the absorption-type grating 21, and the phase-type grating used instead of the absorption-type grating 21 has a Talbot interference effect. The resulting fringe image (self-image) is projected onto the FPD 20.

  The difference between the phase type grating and the absorption type grating is only the thickness of the X-ray high absorption material (X-ray shielding part), and the thickness of the X-ray shielding part is about 30 μm in terms of Au in the case of the absorption type grating. On the other hand, in the case of a phase type grating, it is set to about 1 μm to 5 μm. In the phase-type grating, the X-ray high-absorbing material applies a predetermined amount (π or π / 2) of phase modulation to the X-rays irradiated from the X-ray source 11, thereby generating a Talbot interference effect and causing fringes. An image (self-image) is generated.

  In the above embodiment, the subject H is arranged between the X-ray source 11 and the absorption type grating 21, but the subject H can also be arranged between the absorption type grating 21 and the FPD 20. In this case, a phase contrast image can be generated in the same manner.

In the above embodiment, although generating an intensity modulated signal by combining the pixel data D0~D3 obtained by sampling one each, as shown in FIG. 9, than the pattern period p ₂ of the G1 image The phase differential image generation unit 23 is configured to generate an intensity modulation signal by setting a larger range as an integration interval and integrating the pixel data D0 to D3 included in the integration interval with respect to pixel data of the same image. Also good. This figure, the integrating period, an example in which the set to correspond to 3 times the pattern pitch p _2, differential phase image generator 23, by integrating the pixel data D0~D3 respectively in each integrating period, new Pixel data D0 ′ to D3 ′ are generated, and intensity modulation signals are generated from the pixel data D0 ′ to D3 ′.

Pattern period p ₂ is as small as about several [mu] m, the signal value of the pixel data D0~D3 is very small, by integrating pixel data as described above, although the resolution is reduced, since the signal amount is increased, A few phase contrast images can be obtained.

The pixel size in the y direction of the phase contrast image, because determined by the arrangement pitch in the y direction of the transparent linear electrodes 34a and the light blocking linear electrodes 34b, when the arrangement pitch is larger than the pattern pitch p _2, said By making the integration interval coincide with the arrangement pitch, the pixel sizes in the x direction and the y direction can be made equal.

  In the above embodiment, the image data output from the FPD 20 is temporarily stored in the memory 13, and then the image data is sampled by the sampling processing unit 22, thereby generating an image group of k = 0 to 3. Therefore, there is a problem that the processing time is long. Therefore, it is also preferable to use the sampling processing circuit 50 having the configuration shown in FIG.

  The sampling processing circuit 50 stores a selector 51 that distributes image data output from the FPD 20 for each pixel data, an addition circuit 52 that adds a predetermined number of pixel data distributed by the selector 51, and the added pixel data. Buffer 53. The selector 51 includes first to fourth switches 51a to 51d, and the adder circuit 52 is divided into first to fourth adder circuits 52a to 52d so as to correspond to the first to fourth switches 51a to 51d. It is divided into first to fourth buffers 53a to 53d. The selector 51 and the addition circuit 52 are configured by an FPGA or the like. The buffer 53 may be composed of one memory composed of a plurality of sectors, and each of the first to fourth buffers 53a to 53d may be composed of different memories.

  When the pixel data I1 is input from the FPD 20 to the selector 51, the first and fourth switches 51a and 51d are turned on, and the pixel data I1 is input to the first and fourth adder circuits 52a and 52d. When the pixel data I2 is input, the first and second switches 51a and 51b are turned on, and the pixel data I2 is input to the first and second adder circuits 52a and 52b. When the pixel data I3 is input, the second and third switches 51b and 51c are turned on, and the pixel data I3 is input to the second and third adder circuits 52b and 52c. When the pixel data I4 is input, the third and fourth switches 51c and 51d are turned on, and the pixel data I4 is input to the third and fourth adder circuits 52c and 52d.

  The first addition circuit 52a performs addition processing each time a set of pixel data I1 and I2 is input, generates pixel data D0, and inputs the pixel data D0 to the first buffer 53a. The second addition circuit 52b performs addition processing every time a set of pixel data I2 and I3 is input, generates pixel data D1, and inputs the pixel data D1 to the second buffer 53b. The third addition circuit 52c performs addition processing each time a set of pixel data I3 and I4 is input, generates pixel data D2, and inputs the pixel data D2 to the third buffer 53c. The fourth addition circuit 52d performs addition processing every time a set of pixel data I4 and I1 is input, generates pixel data D3, and inputs the pixel data D4 to the fourth buffer 53d.

  With the above processing, the above-described image groups of k = 0 to 3 are generated in the first to fourth buffers 53a to 53d, respectively, and these image groups are input to the phase differential image generation unit 23. As described above, the sampling processing circuit 50 performs the sampling process while reading out the pixel data from the FPD 20, so that the image group can be generated at high speed.

Further, in the above embodiment, are set scanning step of the read line light source 36 as divided into four pattern period p ₂ of the G1 image, the number of divisions is not limited thereto, it can be appropriately changed. Further, it is preferable that the sampling pattern (a combination of pixel data to be sampled and added) by the sampling processing units 22 and 50 corresponds to the bright part of the G1 image. In this case, the amplitude (contrast) of the intensity modulation signal is the highest. Although it increases, it is not always necessary to set the sampling pattern so as to correspond to the bright part of the G1 image.

  The present invention can be applied to other radiography systems for industrial use, in addition to radiography systems for medical diagnosis.

DESCRIPTION OF SYMBOLS 10 X-ray imaging system 11 X-ray source 11a X-ray focus 12 Imaging part 14 Image processing part 16 Imaging control part 18 System control part 20 Flat panel detector (FPD)
DESCRIPTION OF SYMBOLS 21 Absorption-type grating | lattice 21a Line shielding part 22 Sampling process part 23 Phase differential image generation part 24 Phase contrast image generation part 30 1st electrode layer 30a Surface 31 Recording photoconductive layer 32 Charge transport layer 33 Reading photoconductive layer 34 1st 2 electrode layer 34a transparent linear electrode 34b light shielding linear electrode 35 glass substrate 36 readout line light source 37 power storage unit 38 power supply circuit 39 charge amplifier 39a reset switch

Claims (6)

A radiation source that emits radiation; and
A grid that passes the radiation to generate a fringe image;
The fringe image is recorded as an electrostatic latent image, and the recorded electrostatic latent image is scanned with reading light at a pitch shorter than the period in the periodic pattern direction of the fringe image, whereby the electrostatic latent image is converted into image data. An optical reading radiation image detector that reads out as
Sampling processing for generating an image group composed of a plurality of intensity-modulated images by sampling the image data at a plurality of relative positions having the same period as the periodic pattern and having different phases with respect to the periodic pattern. Means,
Image processing means for generating a phase contrast image of a subject arranged between the radiation source and the radiation image detector based on the image group generated by the sampling processing means;
A radiation imaging system comprising: The image processing means, based on the image group generated by the sampling processing means, a phase differential image generating means for calculating a phase shift amount of an intensity modulation signal for each pixel and generating a phase differential image;
The radiation imaging system according to claim 1, comprising phase contrast image generation means for generating a phase contrast image by integrating the phase differential image in the periodic pattern direction.   The radiation imaging system according to claim 1, wherein the grating is an absorption grating, and projects the radiation from the radiation source onto the radiation image detector as a fringe image.   The radiographic imaging according to claim 1, wherein the grating is a phase type grating, and projects radiation from the radiation source as a fringe image onto a radiation image detector by a Talbot interference effect. system. The radiation image detector is generated in a first electrode layer that transmits radiation, a recording photoconductive layer that generates a charge pair by incidence of radiation transmitted through the first electrode layer 30, and a recording photoconductive layer. The charge pair that acts as an insulator for the charge of one polarity of the charge pair, and acts as a conductor for the charge of the other polarity, The reading photoconductive layer to be generated, the second electrode layer having a plurality of linear electrodes, and extending in a direction orthogonal to the linear electrodes and driven in the extending direction of the linear electrodes, the reading light A read line light source that irradiates the read photoconductive layer through the second electrode layer,
5. The radiography according to claim 1, wherein a power storage unit that accumulates the electrostatic latent image is formed at an interface between the recording photoconductive layer and the charge transport layer. system.   The linear electrode includes a transparent linear electrode that is connected to a charge amplifier for reading out charges and transmits the reading light, and a light-shielding linear electrode that is provided with a ground potential and shields the reading light. The radiographic system according to claim 5, wherein

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