WO2014126189A1

WO2014126189A1 - X-ray imaging device and x-ray imaging method

Info

Publication number: WO2014126189A1
Application number: PCT/JP2014/053436
Authority: WO
Inventors: 浩一尾川; 明敏勝又; 勉山河; 政廣辻田
Original assignee: 株式会社テレシステムズ
Priority date: 2013-02-14
Filing date: 2014-02-14
Publication date: 2014-08-21
Also published as: JPWO2014126189A1

Abstract

[Problem] To accurately and simply execute a CT reconstruction in order to identify substance type in an imaged jaw or the like of a patient. [Solution] This x-ray imaging device is equipped with an x-ray tube and a photon-counting detector. The detector is equipped with a pixel group obtained by two-dimensionally positioning pixels for outputting an electrical pulse according to the photons of an x-ray, and a counting circuit for capturing the number of x-ray photons per pixel. This counting circuit has N number (N≥3) of energy thresholds assigned thereto in order to discriminate the energy for each of the x-ray photons. As a result, the electrical pulse outputted from each pixel is discriminated by electrical thresholds corresponding to the N number of energy thresholds. Consequently, the number of x-ray photons is captured for each pixel according to each of the (N-1) number of energy bands. When CT imaging, imaging data based on the number of x-ray photons is captured for each of at least two energy bands among the (N-1) number of energy regions. On the basis of this imaging data, a CT image is reconstructed for each energy band.

Description

X-ray imaging apparatus and X-ray imaging method

The present invention relates to an X-ray imaging apparatus that irradiates a desired imaging region of an object, detects radiation transmitted through the object, and identifies a tissue constituting the imaging region based on the detection data, In particular, tissue identification of an imaging region based on X-ray absorption data corresponding to the number of X-ray photons (photons) in each of a plurality of X-ray energy bands detected by a photon counting (photon counting) type detector, The present invention relates to an X-ray imaging apparatus and an X-ray imaging method that provide functions such as data collection of shape data and imaging of tomographic images.

In recent years, there have been remarkable technological advances in devices that diagnose and image the internal structure and functions of objects using radiation such as X-rays and gamma rays. A detector for detecting radiation is indispensable for this type of apparatus, and the improvement of the performance of this radiation detector plays a part in the above-mentioned technological progress. In particular, so-called digitization, in which detection signals are output in a digital format, is becoming increasingly finer and the detection surface is becoming larger.

As for the radiation detection method of this radiation detector, in addition to the conventional integration method (integration mode), a detection method called a photon counting method (photon counting method) is also attracting attention. This photon counting method has been conventionally used in gamma ray detectors in the field of nuclear medicine (see, for example, Patent Document 1). On the other hand, in recent years, there have been reports of cases in which this photon counting method is applied to an X-ray detector in order to obtain effects such as enhancement of image enhancement capability, reduction of metal artifacts, and reduction of the effects of beam hardening.

The one proposed in Patent Document 2 is known as one example of this type. In other words, “the photon counting type detector that outputs the electric signal corresponding to the energy of the particle by regarding the radiation incident on each of the plurality of pixels as a photon is provided. Based on this, the count value of the number of photons of the radiation classified into a plurality of energy regions on the energy spectrum of the radiation is calculated, and the calculated count value of each of the plurality of energy regions for each pixel is given for each energy region. The radiation detection apparatus that performs weighting of the weighting factors, adds the count values of the plurality of weighted energy regions for each pixel, and outputs the added data as radiation image generation data for each pixel.

Thus, in the case of a photon counting X-ray detector, one or more (preferably a plurality of) thresholds for discriminating the energy of each incident X-ray photon are prepared. Since the energy range is defined by this threshold value, it is possible to determine to which energy range the energy of each X-ray photon belongs. As a result of this determination, the number of X-ray photons discriminated into the respective energy ranges is counted. This count value information is reflected as the pixel value of the image.

On the other hand, in recent years, attempts have been made to identify (or identify or identify) the types of substances contained in the measurement object using an X-ray apparatus. As an attempt of this identification, as described above, there is one using a photon counting X-ray flat panel detector capable of outputting energy information divided into energy bands (for example, Non-Patent Document 1 and Non-Patent Document 2 related to the paper). reference). Specifically, this photon counting X-ray flat panel detector is mounted on a CT scanner, and the substance is identified (identified) using the detected energy information. In the case of this identification method, interesting results have been obtained at the research level. However, this technology remains at the research level and has not yet been put into practical use in terms of poor performance balance and productivity.

Also known is a method of identifying a substance using a CT scanner. In this case, an X-ray tube for supplying separate tube voltages and two detectors each paired with these X-ray tubes are mounted. Therefore, the substance is identified using the difference in X-ray energy detected by the two detectors. The number of X-ray tube / detector pairs is not necessarily plural, and a pair of X-ray tubes and detectors may be used. In this configuration, for example, the tube voltage supplied to the X-ray tube is switched between two at high speed at a speed faster than the acquisition speed, and two types of images collected under the two tube voltages are separated and extracted. Substance identification can also be performed.

In the background of substance identification using such a CT scanner, there is a case where the CT scanner is a device for obtaining a so-called CT value. In other words, the CT value differs if the type of substance is different, and even if the substance is the same, the CT value will differ if the X-ray energy (tube voltage) applied is different. It is to try.

JP-A-11-109040 JP 2006-101926 A

However, in the case of a method in which two X-ray tubes that apply two different tube voltages are mounted, it has already been commercialized, but there is a limit to accurate substance identification. For this reason, it is applied to identification with a large difference in CT values that is relatively easy to identify, and the accuracy is insufficient to identify subtle differences in substances. The reason is that, since the X-ray has a continuous spectrum, there is an overlap between the two X-ray energy bands, and the energy band information cannot be correctly extracted. This is because the identification accuracy is limited.

If you want to switch three or more tube voltages, you need to add two or more tubes and use three or more X-ray tubes, or switch the tube voltage at a higher speed. On the other hand, if the S / N ratio is to be maintained, the X-ray dose increases, which is not preferable in terms of X-ray exposure.

The present invention has been made in view of the above circumstances, and is detected by a photon counting type detector in an X-ray imaging apparatus having a dental CT scanner function for imaging an imaging target such as a subject's jaw. Provided is an X-ray imaging apparatus capable of accurately and simply executing CT reconstruction processing for identifying (or identifying and identifying) the type of substance existing in an imaging target using the projection data for each energy region. That is the purpose.

In order to achieve the above object, an X-ray imaging apparatus according to one aspect of the present invention includes two X-ray tubes that emit white X-rays and two pixels that output an electric pulse corresponding to incident X-ray photons. A detection circuit having a dimensionally arranged pixel group and N or more energy thresholds (N ≧ 4) are given to the continuous energy of the X-ray, and the N electric pulses output from each pixel are given. And the number of photons of the X-rays is divided into the N energy thresholds divided by the N energy thresholds and between the N energy thresholds. And a photon counting detector having a collecting means for collecting each pixel, the X-ray tube and the detector face each other, and at the time of CT imaging, the X-ray tube and the detector A support that is rotatably supported around Means, scanning means for rotating the X-ray tube and the detector around the object to be imaged based on the instructed scanning method at the time of the CT imaging, and at least one of the “N−1” energy bands For each of the two energy bands, the projection data based on the number of photons of the X-rays collected by the collecting means is subjected to reconstruction processing for CT (computed tomography), and each of the at least two energy bands. Reconstructing means for reconstructing a CT image.

According to another aspect, a detection circuit having an X-ray tube that emits white X-rays, and a pixel group in which pixels that output electric pulses corresponding to incident X-ray photons are two-dimensionally arranged; A photon counting detector having a collecting means for collecting the number of photons of the X-ray for each pixel, the X-ray tube and the detector facing each other, and at the time of CT imaging, the X-ray tube and the detection A support unit that rotatably supports the imaging device around the imaging target, and a scanning unit that rotates the X-ray tube and the detector around the imaging target based on the instructed scanning method during the CT imaging, An X-ray imaging method is provided in an X-ray imaging apparatus comprising: In this X-ray imaging method, the CT imaging is performed in a state where N energy thresholds (N ≧ 3) or more are given to the acquisition means with respect to the continuous energy of the X-ray, and the CT image output from each pixel is output. An electric pulse is discriminated by the N energy thresholds by the collecting means, and the number of photons of the X-ray is divided by the N energy thresholds and “N−” between the N energy thresholds. The X-rays collected for each pixel in accordance with each of the “1” energy bands, and collected by the collecting means for each of at least two energy bands of the “N−1” energy bands. The projection data based on the number of photons is subjected to reconstruction processing for CT (computed tomography) to reconstruct CT images of the at least two energy bands.

In the present invention, “identifying (identifying, identifying) a substance” means identifying the type of substance present in an imaging target such as a subject's jaw from a CT image obtained by reconstructing projection data. It is a concept that includes knowing the thickness (relative and absolute thickness) of the target substance. This material identification is called X-ray transmission information based on a count value obtained by capturing X-rays as particles, that is, photons, and counting the X-ray photons for each energy region, and beam hardening exhibited by the X-ray particles. And information on curing of wire quality. Beam hardening is a phenomenon in which the high energy component of X-rays is relatively larger than the low energy component, and generally uses the expression that the quality of the wire becomes harder (wire quality hardening phenomenon). This phenomenon of linear hardening usually becomes stronger (larger) as the thickness of the substance is larger and the density of the substance is higher. Moreover, although there is no difference in transmission information depending on the substance, there are many cases in which there is a difference in the radiation hardening phenomenon. Conversely, there is a case where there is no difference in the line hardening phenomenon but there is a difference in the transmission information.

Preferably, the apparatus includes a substance identification unit that identifies the type of substance that forms the imaging target based on pixel values of the image of each of the at least two energy bands reconstructed by the reconstruction unit. It is.

For example, the substance identifying means is a quantity based on a sum of pixel values of the image of all the at least two energy bands, and a relative attenuation index for each pixel representing a relative attenuation with respect to a reference substance, and the at least 2 Two or more dimensions including a pixel quality change index for each pixel based on a division between a pixel value of the image in the higher energy band of the energy bands and a pixel value of the image in the lower energy band It is desirable to provide a scatter diagram creating means for creating a scatter diagram having the following dimensions.

Furthermore, it is desirable that the highest threshold value among the N energy threshold values is set to a desired upper limit value of the energy value of the X-ray photon.

According to the present invention, when performing CT imaging of an imaging target such as a subject's jaw, the number of X-ray photons is “N−1” (N ≧ 3) using a photon counting detector. It counts for every pixel according to each energy band. Reconstruction processing for CT (computed tomography) is performed on projection data based on the counted number of photons of X-rays for each of at least two energy bands among the “N−1” energy bands. . Thereby, the image of each of the at least two energy bands is reconstructed. For this reason, CT image data reflecting the number of X-ray photons obtained in this energy band can be used to create a scatter diagram having two or more dimensions including, for example, a relative attenuation index and a quality change index. Become. Since this scatter diagram shows the scatter characteristics according to various substances included in the imaging target, the substance can be identified using this scatter diagram. Therefore, the type of substance to be imaged can be identified (identified and discriminated) more accurately and easily than various conventional substance identification methods.

In the accompanying drawings,
FIG. 1 is a perspective view of an X-ray hybrid machine having a CT imaging function according to an embodiment. FIG. 2 is a diagram for explaining the arrangement of the X-ray tube and the detector of the X-ray hybrid machine, FIG. 3 is another view for explaining the arrangement of the X-ray tube and the detector of the X-ray hybrid machine together with the rotatable directions of the X-ray tube and the detector and their facing states. FIG. 4 is a diagram of an arrangement example illustrating the outline of the configuration of the detector, FIG. 5 is a block diagram illustrating the electrical configuration of the detector. FIG. 6 is a graph for explaining the relationship between an electrical pulse detected in response to the incidence of X-ray photons and a threshold value, which is given to a photon counting detector. FIG. 7 is a graph for explaining an example of the relationship between the energy spectrum with respect to the incidence frequency (count value) of X-ray photons and the energy region given by the discrimination circuit; FIG. 8 is a block diagram showing an outline of the electrical configuration of the entire X-ray hybrid machine, FIG. 9 is a flowchart for explaining the flow of substance identification processing executed in the embodiment; FIG. 10 is a schematic diagram illustrating the relationship between the energy region and the reconstruction process for each region. FIG. 11 is a graph showing dimensions of a scatter diagram adopted in the embodiment. FIG. 12 is a graph schematically showing an example of a scatter diagram created by the substance identification process of the embodiment; FIG. 13 is a graph schematically showing how the types of substances in the jaw are identified on the scatter diagram. FIG. 14 is a graph illustrating the advantage of discriminating two substances using a two-dimensional scatter diagram. FIG. 15 is another graph illustrating the advantages of discriminating two substances using a two-dimensional scatter diagram. FIG. 16 is a circuit diagram showing a modification of the detector.

Hereinafter, embodiments of the present invention will be described with reference to the accompanying drawings.

(First embodiment)
A first embodiment of an X-ray CT apparatus according to the present invention will be described with reference to FIGS.

This X-ray CT apparatus is configured as a dental X-ray hybrid machine that also has an X-ray panoramic imaging function that captures a panoramic image of a subject's jaw (including a dentition) as an imaging target. In other words, this single dental X-ray hybrid machine has one piece of hardware, but the operator selects the type of software from “panoramic photography” and “CT photography” so that the same hardware can be used. Both imaging can be selectively performed using the wear.

When “Panorama Imaging” is selected, a pseudo three-dimensional cross-sectional image of the jaw of the subject (the image itself is a two-dimensional image, but is displayed in a three-dimensional manner according to the shape of the imaging part such as a dentition) Can be taken). When “CT imaging” is selected, a CT image of the jaw of the subject can be obtained. Both functions use an X-ray tube and an X-ray detector (hereinafter simply referred to as a detector) that go around the jaw of the subject. X-rays emitted from the X-ray tube pass through the jaw and are detected by a detector. As will be described later, this detector captures X-rays (actually, X-ray beams formed into a fan shape or a cone shape by a slit (aperture)) as a bundle of photons (particles), and the photons It is a photon counting (photon counting) type detector that obtains counting information according to the number of the photons. For this reason, the X-ray transmission data transmitted through the jaw reflects the number of photons. The fan-shaped X-ray beam is used for panoramic imaging, and the cone-shaped X-ray beam is used for CT imaging.

Therefore, this dental X-ray hybrid machine is configured to operate to generate image data of a desired or specified cross section of the jaw under the tomosynthesis method when obtaining a “panoramic image”. Yes. Thereby, panoramic image data of the cross section is created. Since this panoramic photography itself is known, the description thereof is simplified or omitted in the present embodiment.

In addition, although the X-ray hybrid machine which concerns on this embodiment is demonstrated as a dental imaging device, the use of this X-ray hybrid machine is not necessarily restricted to dentistry. As other imaging applications, it can be applied to imaging various parts such as mammography, otolaryngology, bones and joints of limbs. It can also be applied to uses such as corpse identification for identity identification and nondestructive inspection.

FIG. 1 shows an appearance of a dental X-ray hybrid machine 1 according to this embodiment.

The dental X-ray hybrid machine 1 includes a pedestal 12 on which a caster 11 is mounted, an elevating unit 13 and a power supply box 14 mounted on the pedestal 12, and a console 17. The elevating unit 13 includes an elevating mechanism (not shown) therein, and the upper elevating unit of the unit is configured to be movable up and down within a predetermined range electrically with respect to the base 12 (that is, the floor surface). Yes. If the vertical movement direction of the elevating unit 13 is the Z axis, XYZ orthogonal coordinates as shown in the figure can be assumed. The power supply box 14 supplies necessary power to each part of the system.

In addition, the dental X-ray hybrid machine 1 includes two

arms

15 and 16 extending in the X-axis direction (that is, laterally) from the lifting unit of the lifting unit 13. The two

arms

15 and 16 are in the Y-axis direction. Are both formed in a substantially L shape, are superposed so that one end of each of the

arms

15 and 16 overlaps each other, and is attached to the side surface of the elevating unit. Is equipped with a rotation mechanism 13D that can rotate the two

arms

15 and 16 independently of each other, that is, at different speeds. The X-ray tube 21 and the detector 22 are each equipped with a slit (aperture) 23 for forming X-rays into a fan shape or a cone shape on the front surface of the X-ray tube 21 on the X-ray irradiation side. The rotation mechanism 13D and

arms

15 and 16, for the X-ray tube 21 and the detector 22, the support means for driving rotatably supported independently of one another are constructed.

The X-ray tube 21 is configured as a rotating anode type X-ray tube using, for example, tungsten as an anode material. The X-ray tube 21 has a dotted X-ray focal point (for example, a diameter of 0.1 mm to 0.5 mm) FP. The X-ray tube 21 emits pulsed X-rays in response to driving power supplied from a high-voltage generator described later. X-rays exposed from the X-ray focal point FP of the X-ray tube 21 are narrowed by the slit 23 and formed into a fan-shaped X-ray beam. Thereafter, the X-ray beam passes through the jaw JW of the subject P and is attenuated, and the transmitted X-ray beam reflecting the attenuated state is incident on the detector 22.

At the time of imaging, as shown in FIG. 2, the jaw portion JW of the subject P is positioned at a predetermined position in the three-dimensional imaging space IS defined between the X-ray tube 21 and the detector 22. For this reason, the X-ray tube 21 and the detector 22 face each other (face to face) across the jaw. The irradiated X-ray beam passes through the jaw JW (such as a dentition) and is then detected by the detector 22. Since the two

arms

15 and 16 are rotationally driven by the rotation mechanism 13D at the time of imaging, the X-ray tube 21 and the detector 22 rotate around the jaw part along a predetermined circular orbit around the rotation center O. During the rotation, irradiation and detection of the X-ray beam are executed at predetermined intervals.

When viewed along the X-axis direction facing the YZ plane, the X-ray tube 21 and the detector 22 are respectively driven to rotate along circular trajectories Tx and Td centered on the rotation center O determined in advance on the system side. The The radii Dx and Dd from the rotation center O to the circular trajectories Tx and Td are set to different values in consideration of X-ray exposure, detection accuracy, downsizing of the apparatus, mechanical interference with the patient, and the like. (See FIG. 2). In the present embodiment, Dx ≠ Dd, and particularly Dx> Dd. The reason why the distance (radius Dd) from the rotation center O to the detector 22 is smaller than that (radius Dx) from the rotation center O to the X-ray tube 21 is that the position of the detector 22 is set as much as possible. This is to reduce the attenuation of the incident intensity of X-rays. The distance (radius Dx) from the rotation center O to the X-ray tube 21 is set to a value that can ensure the distance between the X-ray tube and the skin defined by the standard.

Therefore, the X-ray tube 21 and the detector 22 are always opposed to each other (facing to each other), and irradiation and detection of X-rays along a plurality of predetermined desired X-ray paths with respect to the jaw JW (dentition) are performed. In order to perform the above, the X-ray tube 21 and the detector 22 are independently driven at different angular velocities.

Note that “opposing each other” described above is a cone-like shape that is irradiated from the dotted X-ray focal point FP of the X-ray tube 21 when viewed along the X-axis direction as shown in FIG. The state in which the irradiation range of the shaped X-ray beam and the X-ray detection surface 22A (described later) of the detector 22 coincide with each other. In particular, a state in which the X-ray beam includes an axis T that intersects the center position C of a module Bm (described later) at 90 ° in the width direction of the X-ray detection surface is referred to as a “facing state”. (See FIG. 3).

For this reason, in order to realize the above-mentioned “always opposite (or facing each other)”, the opposing

arm portions

15A and 16A including the X-ray tube 21 and the detector 22 out of the

arms

15 and 16 have the axis It can be independently rotated (rotated) around AXs and AXd (see FIGS. 1 to 3). For this purpose,

rotation driving mechanisms

15B and 16B such as motors are provided on the

arms

15 and 16, respectively. The drive control of the

rotation drive mechanisms

15B and 16B is executed by a controller of the console 17 described later.

In the present embodiment, the positions of the intersection position C and the axis AXd are matched in the Y-axis direction. Further, the circular trajectories Tx and Td shown in FIG. 3 follow the positions of the axes AXs and AXd, respectively, when viewed in the YZ plane.

As shown in FIG. 4, the detector 22 has an array (sensor circuit) of a plurality of detection modules B1 to Bm in which X-ray imaging elements are two-dimensionally arranged. The plurality of detection modules B1 to Bm are created as blocks independent of each other, and are mounted in a predetermined shape (for example, a rectangular shape) on a substrate (not shown) to form the entire detector 22.

A plurality of detection modules B1 to Bm are provided in two dimensions in the vertical (X-axis) and horizontal (Y-axis) directions (15 in the vertical direction, horizontal direction) while providing a certain gap between the individual modules. 8, together with arranging by further placing the 5 upper and lower ends), they are arranged obliquely inclined by an angle θ with respect to the scanning direction O _Y individual modules. This angle θ is set to about 14 °, for example. The surface of a rectangular shape (in the case of CT imaging) having a small length ratio (in the case of CT imaging) or a large ratio of length to width, that is, an elongated rectangular shape (in the case of panoramic imaging), is created by the plurality of detection modules B1 to Bm. An X-ray detection surface 22A is formed. Since the detection modules B1 to Bm are arranged obliquely, the X-ray detection surface 22A is formed so as to follow (inscribe) the inside of each detection surface of the plurality of modules B1 to Bm.

The structure of the detector 22 having this detection module arranged obliquely and the processing of the detection signal by the sub-pixel method are known, for example, from WO2012 / 0886648A1.

It should be noted that a plurality of modules arranged in a column arranged in a line on the left side in FIG. 4 function as modules for panoramic imaging. The opening area for panoramic photography is indicated by reference numeral 22B. Further, a module group of a two-dimensional array formed by all the modules except the two modules at the upper and lower ends of the left side row and the remaining modules functions as a module for CT imaging. The opening area for CT imaging is indicated by reference numeral 22A. The module group for panoramic imaging or CT imaging is selected by controlling the opening area of the slit 23. Reference numeral AXd in FIG. 4 is a central axis when the detector 22 itself rotates (rotates).

Individual detection modules B1 (˜Bm) are made of a semiconductor material that directly converts X-rays into electrical pulse signals. For this reason, the detector 22 is a photon counting X-ray detector of a direct conversion method using a semiconductor.

As described above, the detector 22 is formed as an array of a plurality of detection modules B1 to Bm. Each detection module Bm includes a detection circuit Cp (see FIG. 5) for detecting X-rays and a data counting circuit 51 _n (see FIG. 5) stacked together with the detection circuit Cp, as is well known. The detection circuit Cp includes, for each detection module, a semiconductor layer that directly converts X-rays into an electrical signal, and a charging electrode and a collecting electrode that are respectively stacked on both sides (not shown). X-rays are incident on the charged electrode. The charged electrode is a common electrode, and a high bias voltage is applied between the charged electrodes. The semiconductor layer and the collecting electrode are divided into a grid pattern, and by this division, a plurality of small regions are formed that are arranged in a two-dimensional array at a certain distance from each other. As a result, a plurality of stacked bodies of semiconductor cells C (see FIGS. 4 and 5) and collecting electrodes arranged in a two-dimensional manner on the charged electrode are formed. The plurality of stacked bodies to form a plurality of pixels S _n arranged in a two dimensional grid pattern.

As a result, a plurality of pixels S _n (n = 1 to N) occupying a predetermined area necessary for the detector 22 by the whole of the plurality of detection modules B1 to Bm (however, depending on the opening area 22A at the time of CT imaging: see FIG. 4). ) Is formed. The plurality of pixels _{S n} constitutes a pixel group Cp (refer to FIG. 5).

The size of each pixel S _n is, for example, 200 [mu] m × 200 [mu] m, the pixel size is set to a detectable value X-rays incident as a set of multiple photons. Each pixel S _n is responsive to incident of each photon of X-ray, and outputs an electrical pulse of amplitude corresponding to the energy possessed by the photon. That is, each pixel S _n may convert the X-rays incident on that pixel directly, into electric signals.

Therefore, the detector 22, the photon constituting the cone beam-like X-rays incident, counts for each pixel S _n which constitute the detection surface of the detector 22, the quantity of electricity that reflects the count value Data is output at a high frame rate of, for example, 75 fps. This data is also called frame data.

As the semiconductor material of the semiconductor layer, that is, the semiconductor cell C, cadmium telluride semiconductor (CdTe semiconductor), cadmium zinc telluride semiconductor (CdZnTe semiconductor (CZT semiconductor)), silicon semiconductor (Si semiconductor), thallium bromide (T1Br) Mercury iodide or the like is used. Instead of this semiconductor cell, it is composed of a cell that combines a scintillator material that is subdivided into columns and optically shielded from each column, and a photoelectric converter composed of a combination of fine avalanche photodiodes. May be.

For this reason, when X-rays enter the semiconductor cell C, charges (electrons, holes) are generated inside the cell, and a pulse current corresponding to the amount of the charge flows. This pulse current is detected by the current collecting electrode. As a result, the amount of charge varies depending on the energy value of the X-ray photons. Therefore, the detector 22 outputs an electrical pulse signal corresponding to the energy value of the photons for respective pixels S _n.

The detector 22 further comprises respective semiconductor cell C, that the data counting circuit ₅₁ on the output side of each of the plurality of pixels _{S n} n the (n = 1 ~ N). Here, each pixel _{S n,} i.e., the path from each semiconductor cell C to the data counting circuit ₅₁ 1 (~ ₅₁ _N), optionally, referred to as acquisition channels _{CN n (n = 1 ~ N} ) (See FIG. 5).

The structure of this group of semiconductor cells C is also known from Japanese Patent Application Laid-Open Nos. 2000-69369, 2004-325183, and 2006-101926.

Incidentally, the size of each pixel S _n described above (200 [mu] m × 200 [mu] m) is adapted to a sufficiently small value that is capable of detecting X-rays as photons (particles). In the present embodiment, the size capable of detecting X-rays as the particles is “between electric pulses responding to each incident when a plurality of radiation (for example, X-ray) particles are successively incident at or near the same position. The occurrence of the superposition phenomenon (also called pile-up) is defined as “a size that can be substantially ignored or whose amount is predictable”.

However, even with such a pixel size, it is not possible to avoid all occurrences of the superposition phenomenon. Even when two or more electrical pulses are both observed in the same pixel, the superposition phenomenon does not occur if they are separated from each other in time. On the other hand, when two or more electric pulses are difficult to separate in time in the same pixel, a superposition phenomenon occurs, and the two electric pulses overlap to be observed as one electric pulse having a high peak value. Is done.

When this superposition phenomenon occurs, X-ray particle countdown (also called pile-up count loss) occurs in the characteristic of “number of incidents versus actual count value” of X-ray particles. Therefore, the size of the pixel S _n to form the X-ray detector 22, the magnitude of which can be regarded as the counting loss does not occur or does not substantially occur, or are set to an extent counting the drop amount can be estimated .

Subsequently, a circuit electrically connected to the detector 22 will be described with reference to FIG. Each of the plurality of data counting circuits 51 _n (n = 1 to N) includes a charge amplifier 52 that receives an electrical signal of an analog amount output from each semiconductor cell C. Circuit 53, multi-stage comparator 54 _i (here i = 1 to 4), energy region distribution circuit 55, multi-stage counter 56 _i (here i = 1 to 4), multi-stage D / A converter 57 _i ( Here, i = 1 to 4), a latch circuit 58, and a serial converter 59 are provided.

Each charge amplifier 52 is connected to each current collecting electrode of each semiconductor cell S, charges up the current collected in response to the incidence of X-ray particles, and outputs it as a pulse signal of electric quantity. The output terminal of the charge amplifier 52 is connected to a waveform shaping circuit 53 whose gain and offset can be adjusted. The waveform of the detected pulse signal is processed with the previously adjusted gain and offset to shape the waveform. The gain and offset of the waveform shaping circuit 53, in consideration of the variation in non-uniformity and the circuit characteristics for charge-charge characteristic for each pixel S _n of semiconductor cell C, is calibrated. As a result, it is possible to increase the output of the waveform shaping signal from which non-uniformity has been eliminated, and the relative threshold setting accuracy. As a result, corresponding to each pixel S _n, i.e., the characteristics reflecting the energy value of the X-ray particle pulse signal waveform formatted output from the waveform shaping circuit 53 for each collection channel CN _n is substantially incident Have. Therefore, the variation between the collection channels CN _n is greatly improved.

The output terminal of the waveform shaping circuit 53 is connected to the comparison input terminals of the plurality of comparators 54 ₁ to 54 ₄ . As shown in FIG. 5, analog amount threshold values (voltage values) th _i (here, i = 1 to 4) having different values are applied to the reference input terminals of the plurality of comparators 54 ₁ to 54 _4, respectively. Yes. This makes it possible to compare one pulse signal with each of the different analog amount thresholds th ₁ to th ₄ . FIG. 6 shows the magnitude relationship (th ₁ <th ₂ <th ₃ <threshold) between the peak value (representing energy) of the pulse voltage generated in response to the input of one X-ray photon and the threshold values th ₁ to th _4. th ₄ ) schematically.

The reason for this comparison is to examine which region (discrimination) the energy value of the incident X-ray particle enters among the energy regions set in advance divided into a plurality. A determination is made as to which of the analog amount threshold values th ₁ to th ₄ exceeds the peak value of the pulse signal (that is, the energy value of the incident X-ray photon). Thereby, the energy area | region discriminated differs. Note that the lowest analog amount threshold th ₁ is usually used to prevent detection of disturbance, noise due to circuits such as the semiconductor cell S and the charge amplifier 52, or low-energy radiation that is not necessary for imaging. Is set as the threshold value. Further, the number of thresholds, i.e., the number of comparators is not necessarily limited to four, three, including the amount of the analog amount threshold th _1, or may be five or more.

Analog amount threshold th 1 _~ th ₄ described above, specifically, given from the calibration computing unit 38 of the console 17 for each pixel S _n in a digital value through the interface 31, i.e., for each acquisition channels. Therefore, the reference input terminals of the comparators 54 ₁ to 54 ₄ are connected to the output terminals of the _four D / A converters 57 ₁ to 574, respectively. The D / A converter ₅₇ 1-57 ₄ is connected to the threshold receiving end _T 1 via the latch circuits 58 (~ _{T N),} the interface 31 of the threshold receiving end _{_T} 1 (~ _T _N) console 17 It is connected.

The latch circuit 58 latches the thresholds th ₁ ′ to th ₄ ′ of digital quantities given from the threshold applier 41 via the interface 31 and the threshold receiving end T ₁ (to T _N ) at the time of imaging, and the corresponding D / are output to a converters ₅₇ 1-57 _4. Therefore, the D / A converters 57 ₁ to 57 ₄ can supply the commanded analog amount thresholds th ₁ to th ₄ to the comparators 54 ₁ to 54 ₄ as voltage amounts, respectively. Each collection channel CN _n is one from the D / A converter 57 _i (i = 1 to 4) to the counter 56 _i (i = 1 to 4) via the comparator 54 _i (i = 1 to 4). Or it is connected to a plurality of circuit systems. This circuit system is called “discrimination circuit” DS _i (i = 1 to 4).

FIG. 7 shows a setting example of the energy threshold TH _i (i = 1 to 4) corresponding to the analog amount threshold th _i (i = 1 to 4). This energy threshold TH _i (i = 1 to 4) is of course a discrete value that is set discretely and can be set to an arbitrary value by the user. FIG. 7 schematically shows an X-ray spectrum when an appropriate material is used for the anode material of the X-ray tube 21. The horizontal axis indicates X-ray energy, and the vertical axis indicates the incidence frequency of X-ray photons. This incidence frequency is a factor representative of the count value (count) or intensity of X-ray photons.

The analog amount threshold th _i is an analog voltage applied to the comparator 54 _i in each discrimination circuit DS _i , and the energy threshold TH _i is an analog value for discriminating the X-ray energy (keV) of the energy spectrum. The waveform shown in FIG. 7 shows a continuous spectrum of the energy of X-rays exposed from an X-ray tube that is normally used, for example, using tungsten as an anode material. The count value (count) on the vertical axis is an amount proportional to the photon generation frequency corresponding to the energy value on the horizontal axis, and the energy value on the horizontal axis is an amount depending on the tube voltage of the X-ray tube 21. With respect to this spectrum, the first analog quantity threshold th ₁ is set as the X-ray photon count non-counting area (the area where there is no meaningful X-ray information for counting and the circuit noise is mixed) and the lower first to set corresponding to the energy region ER ₁ and energy threshold value TH ₁ capable discrimination of. Further, the second and third analog amount threshold values th ₂ and th ₃ are set so as to sequentially provide the second and third energy threshold values TH ₂ and TH ₃ which are higher than the first energy threshold value TH ₁ . . Further, the fourth energy threshold TH ₄ is set to an energy value equal to the applied voltage to the X-ray tube, in which X photon count value = 0 if there is no superposition phenomenon in the energy spectrum. Here, the fourth energy threshold TH _4, each pixel S _n, it is one of the important features of the present application are in accordance with the energy value of the count value = 0.

As a result, appropriate discrimination points based on energy spectrum characteristics and design values are defined, and energy regions ER ₁ to ER ₄ are set.

These energy thresholds TH _i are determined so that one or more subjects as a reference are assumed and the count value for a predetermined time for each energy region is substantially constant.

Therefore, the output of the comparator ₅₄ 1-54 _3, as shown in FIG. 5, is connected to the input ends of the plurality of counters ₅₆ 1-56 _4.

Each of the counters ₅₆ 1-56 ₄ counts up every time the output of the comparator ₅₄ 1-54 ₃ (pulse) is turned on. As a result, the number of X-ray photons having energy equal to or higher than the energy value discriminated into the energy region ER ₁ (to ER ₄ ) that each counter 56 ₁ (to 56 ₄ ) takes charge of is accumulated value W ₁ ′ ( it can be counted for each pixel _{S n} as ~ _{W 4} ').

Specifically, this counting operation is determined by the relationship between the detection voltage V _dec (detected energy value of photons) input to the _four comparators 54 ₁ to 54 ₄ and the threshold values th ₁ to th ₄ . That is, when the detection voltage V _dec <th ₁ to th ₄ , the outputs of all the comparators 54 ₁ to 54 ₄ are turned off. In other words, the output = 0 of the pixel _{S n.} As a result, noise components smaller than the energy threshold TH ₁ defined as the input energy counting limit are not counted. This noise component corresponds to an energy value signal belonging to the non-countable region ERx in FIG.

However, if the detection voltage V _dec exceeds the minimum threshold th ₁ (V _dec ≧ th ₁ ), the number of photons is counted. If the relationship is V _dec ≧ th ₁ , the outputs of all the comparators 54 ₁ to 54 ₄ are turned on. That is, the count value _W 1 of all the counters _{_{_{56 1 ~ 56 4 '~ W}}} 4' is counted up.

When the relationship of V _dec ≧ th ₂ is established, the outputs of the three comparators 54 ₂ to 54 _{4 in} the second and subsequent stages are turned on. Thus, the three counters ₅₆ 2-56 ₄ counts _{_W} 2 '~ _W _4' is counted up. If the relationship of V _dec ≧ th ₃ is established, the outputs of the third-stage and fourth-stage comparators 54 ₃ and 54 ₄ are turned on. As a result, the count values W ₃ ′ and W ₄ ′ of the two counters 56 ₃ and 56 ₄ are counted up.

_Furthermore, if the relationship between _{V dec} ≧ th _4, the output of the comparator module 54 ₄ of the fourth stage is turned on, only the counter 56 ₄ count value _{W 4} 'of the fourth stage is counted up. In this case, the energy value of the photon related to the input is a noise component belonging to the region ER ₄ exceeding the third high energy region ER ₃ , disturbance, etc., which is not suitable for imaging or counting. On the other hand, the count value W ₄ ′ can be used as information for estimating or excluding photons that have caused a superposition phenomenon or simultaneously incident photons.

As described above, in the present embodiment, the counters 56 ₁ to 56 ₄ count the number of photons having energy exceeding the energy range ER ₁ (to ER ₄ ) to be counted by the counters 56 ₁ to 56 ₄ , respectively. Therefore, the number of X-ray photons having energy belonging to each of the first to fourth energy regions ER ₁ to ER ₄ , that is, the number of X-ray photons to be obtained for each energy region is expressed as W ₁ , W ₂ , W ₃ , W ₄ , the relationship between the count values W ₁ ′, W ₂ ′, W ₃ ′, W ₄ ′ of the counters 56 ₁ to 56 ₄ is
W ₁ = W ₁ '-W ₂ '
W ₂ = W ₂ '-W ₃ '
W ₃ = W ₃ '-W ₄ '
It becomes. Note that W ₄ = W ₄ ′ is not calculated because it is meaningless information (that is, the energy region of the X-ray photon cannot be specified) due to the superposition phenomenon.

Therefore, the count values W ₁ to W _{4 that} are to be truly obtained are obtained by subtraction processing based on the above equation by a data processor described later. Ideally, W ₄ = W ₄ ′ = 0.

Thus, in the present embodiment, X-ray photon number _W 1 to _{W 4} belonging to each of the first through fourth energy regions _ER 1 to ER _4, the actual count value _W 1 '~ _{W 4'} Is obtained by calculation (subtraction). For this reason, a circuit for deciphering which energy region ER1 to ER4 the current event, that is, the incidence of X-ray photons belongs, becomes unnecessary from the combination of turning on and off the outputs of the comparators 54 ₁ to 54 ₄ . This simplifies the circuit configuration mounted on the data counting circuit 51 _n of the detector 22.

In addition, the meaning of “collection” for each energy region of the number of X-ray photons according to the present application is the meaning of “obtaining by calculation” from the actual count value as described above, and for each energy region as in a modification example described later. Both meanings of directly “counting” the number of X-ray photons are included.

The counter 56 _1-56 ₄ described above start and stop signals is supplied via a start-stop terminal T2 from below to the controller of the console 17. Counting for a fixed time is managed from the outside using a reset circuit included in the counter itself.

Thus, during a certain period of time until reset by a plurality of counters 56 _1-56 _4, the number of photons of X-rays incident on the detector 22 is counted for each pixel S _n. The number of photons counted value W _k of the X-ray _'(k = 1 ~ 4), after being outputted from each of the counters 56 _1-56 ₄ in parallel as the count value of the digital quantity, serial format by serial converter 59 Is converted to The serial converter 59 ₁ is connected to the serial converter 59 2 _~ 59 _N and serial all remaining acquisition channels. Therefore, the count of all digital content is output from the last channel of the serial converter 59 _N serially sent to the console 17 via the transmitting end T3.

In the console 17, the interface 31 receives these count values and stores them in a storage unit to be described later.

In the present embodiment, it is integrally constructed in CMOS by the semiconductor cell C and the data counting circuit 51 _n corresponding to N pixels S _n described above ASIC (Application Specific Integrated Circuit). Of course, the data counting circuit 51 _n may be configured as a circuit or device separate from the group of semiconductor cells C.

In the above-described embodiment, the plurality of detection modules B1 to Bm are connected to a scintillator array in which a plurality of scintillators processed into columnar shapes are bundled, and receives light incident from the scintillator. A plurality of avalanche photodiodes are mounted on the light receiving surface, and the avalanche photodiodes belonging to the region are electrically connected by a quenching element for each rectangular region having a predetermined size corresponding to the cell on the light receiving surface. And a silicon photomultiplier.

The material of the scintillator is LFS (lutetium silicate), GAGG: Ce (gadolinium aluminum gallium garnet), LuAG: Pr (praseodymium-added lutetium aluminum garnet), or the same decay time and light emission amount as the LuAG: Pr. It may be a material having a specific gravity.

As shown in FIG. 8, the console 17 includes an interface (I / F) 31 that performs input and output of signals, a controller 33 that is communicably connected to the interface 31 via a bus 32, and a first storage unit 34, a data processor 35, a display unit 36, an input unit 37, a calibration calculator 38, a second storage unit 39, ROMs 40A to 40D, and a threshold value assigner 41.

The controller 33 controls the drive of the X-ray hybrid machine 1 according to a program given in advance to the ROM 40A. This control includes sending a command value to the high voltage generator 42 that supplies a high voltage to the X-ray tube 21 and a drive command to the calibration calculator 38. The first storage unit 34 stores frame data and image data that are count values sent from the detector 22 via the interface 31.

The data processor 35 operates based on a program given in advance to the ROM 40B under the control of the controller 33. During CT imaging, the data processor 35 performs the CT image reconstruction process by processing the frame data stored in the first storage unit 34 by a desired CT reconstruction method. On the other hand, at the time of panoramic photography, the data processor 35 operates tomosynthesis based on a known arithmetic method called “shift and add” to the frame data stored in the first storage unit 34. To implement. Thereby, a CT image or a panoramic image of the oral cavity of the subject P is obtained. The display unit 36 is responsible for displaying an image to be created, information indicating the operation status of the apparatus, and operator operation information given via the input unit 37. The input device 37 is used by an operator to give information necessary for imaging to the apparatus.

Further, the calibration computing unit 38, under the control of the controller 33, operating under program built in advance in ROM40C, giving for each energy discriminator circuit for each pixel S _n in the data counting circuit, X-rays energy Calibrate the digital quantity threshold for discrimination.

Under the control of the controller 33, the threshold value applicator 41 calls the threshold value of the digital quantity stored in the second storage unit 39 for each pixel and for each discrimination circuit at the time of imaging, and uses the threshold value as a command value as an interface. 31 to the detector 22. In order to execute this process, the threshold value assigner 41 executes a program stored in advance in the ROM 40D.

The controller 33, the data processor 35, the calibration calculator 38, and the threshold value assigner 41 are all provided with a CPU (central processing unit) that operates according to a given program. Those programs are stored in advance in each of the ROMs 40A to 40D.

In this embodiment, the data processor 35 reads the count value stored in the first storage unit 34 in response to an operator command from the input device 37, and uses this count value for image processing and substance identification. The commanded process such as the above process and the measurement process is executed. The image processing includes, for example, generation of a panoramic image of a cross section of a dentition based on a tomosynthesis method in “panoramic imaging” and generation of a tomographic image based on a desired reconstruction method in “CT imaging”. .

In addition, substance identification includes identification (specification) of types and states of a plurality of substances constituting the jaw using beam hardening information. This CT image reconstruction and substance identification process is one of the features of the present application. Although the process will be described below representing the features noteworthy were discriminated into a plurality of first to third energy regions ER 1 _to ER _3, respectively, the count value of the X-ray photons from each pixel S _n W ₁ This means that a set of (˜W ₃ ) (which is also collected data) is obtained. For this reason, as in the prior art, the data processor 35 applies moderately high weights to the set of the count values W ₁ to W ₃ and applies the data obtained by adding them to each other to the tomosynthesis method or the CT reconstruction method. Can be hung. Of course, processing can be performed without weighting. On the other hand, are discriminated in the respective first to third energy regions _ER 1 to ER _3, set to CT reconstruction for each area the count value _W 1 of each pixel _{S n} (~ _{W 3),} this result It is also possible to generate a scatter diagram from the obtained reconstruction data and to identify the substance from the scatter diagram.

Next, in the present embodiment, a series of processes from substance identification to three-dimensional rendering executed in cooperation with the controller 33 and the data processor 35 will be described with reference to FIG.

In the driving state of the X-ray hybrid machine 1, the controller 33 starts the process of FIG. 9 in response to an instruction from the operator. First, in step S1, the controller 33 determines whether to perform CT imaging or panoramic imaging interactively with the operator. When it is determined in this step that “perform panoramic imaging”, the process proceeds to step S2 where panoramic imaging of the jaw of the subject P is performed by a known method to obtain, for example, a panoramic image of the dentition.

On the other hand, when the determination in step S1 is “execution of CT imaging”, the controller 33 cooperates with the data processor 35 to sequentially execute the steps after step S3.

First, the controller 33 determines the scanning method interactively with the operator or according to the default setting (step S3). As this scanning method, a full scan and a half scan are prepared. The full scan is a method of collecting data while a pair of the X-ray tube 21 and the detector 22 makes a round (360 degrees) around the jaw. On the other hand, the half scan is a method of collecting data while the pair of the X-ray tube 21 and the detector 22 rotates around the jaw part by a half turn (180 degrees).

Next, the controller 33 instructs the X-ray scan to the jaw of the subject P positioned in the imaging space (step S4). In response to this command, the pair of the X-ray tube 21 and the detector 22 starts to rotate around the jaw, and data collection based on the commanded scanning method is started (step S5). During this rotation, cone beam-shaped X-rays are continuously or pulsedly irradiated from the X-ray tube 21, and the X-rays pass through the jaw. The transmitted X-ray is detected by the detector 22. At this time, the opening of the slit 23 is controlled so that the size of the projection surface of the transmitted X-ray with respect to the detector 22 matches the size and shape of the detection surface 22A.

As a result, are collected in as described above, a constant count of the photons of X-rays that reflect the degree of attenuation for each and discriminator circuit for each pixel S _n in the jaw period (e.g. 75fps). For this reason, for example, in the case of a full scan of 360 °, projection data including the count value of X-ray photons obtained by dividing 360 ° into 720 steps is collected. The projection data is sequentially stored in the first storage unit 34 of the console 17. When this data collection is completed, the controller 33 terminates scanning, that is, X-ray irradiation, and returns each of the tube 21 and the detector 22 to a predetermined home position (step S6). As a result, for example, projection data for “720 steps × 3 energy regions” is collected.

Data processor 35 determines the X-ray photon number with energy values belonging to each of the first to third energy regions _ER 1 to ER ₃ performs the subtraction for each pixel _{S n} (step S7).

Next, the controller 33 interactively determines with the operator whether or not to reconstruct the CT image using the data collected as described above (step S8). When it is determined to reconstruct the CT image (step S8, YES), the data processor 35 is instructed to perform the reconstruction process.

Then, the data processor 35 have been discriminated to each of the first to third energy regions _ER 1 to ER _3, an appropriate weighting to projection data composed of the count value _W 1 (- _{W 3)} of each pixel _{S n} in and (step S9), and the weighted projection data is added to each other for each pixel S _n to form a projection data composed of a set count value, a predetermined reconstruction method the projection data (e.g., FBP method) Reconfiguration is performed (step S10). For example, the high (or low) weighted by the projection data consisting of the count value W ₁ of each pixel S _n which is discriminated in the first energy area ER _1, each pixel is discriminated in the third energy regions ER ₃ S a low (or high) weighted by the projection data consisting of the count value W ₃ of _n. As a result, a CT image in which lower-energy transmitted X-rays or higher-energy transmitted X-rays are emphasized is generated. This weighting can be implemented in various ways, and may be reconfigured without weighting. The generated CT image is displayed on the display 36 (step S11).

On the other hand, if it is determined NO in step S8, the controller 33 further interactively communicates with the operator using the data collected as described above to identify a plurality of types of substances constituting the jaw. Is determined (step S12). Here, the substance identification is a process for specifying the type of each substance. When it is determined NO in step S12, the series of processing ends without performing substance identification.

On the other hand, when the controller 33 determines to perform substance identification (step S12, YES), it instructs the data processor 35 to perform subsequent reconstruction processing. In response, the data processor 35, which is discriminated in each of the first to third energy regions _ER 1 to ER _3, CT re for each region using the projection data consisting count _{W k} of each pixel _{S n} Configure (step S13: see FIG. 10). Thereby, three sets of CT images IM ₁ to IM _{3 in the} 3D (volume) space including the jaws are reconstructed for the _three energy regions ER ₁ to ER _{3 from the} _first to the _third . The data of the CT images IM ₁ to IM ₃ are stored in the first storage unit 34 so as to be readable.

Based on the command from the operator, the controller 33 determines whether or not to display the three sets of CT images IM ₁ to IM ₃ as they are (step S14). If this display is desired (step S14, YES), one or more sets of the CT images IM ₁ to IM ₃ are displayed on the display 36 in an appropriate manner (step S15).

When the controller 33 determines not to display the above (step S14, NO), it is further determined whether to execute the substance identification process interactively (step S16). When this determination is made, the scatter diagram creation process is left to the data processor 35. The data processor 35 reads the data of the _three CT images IM ₁ to IM ₃ from the first storage unit 34, and creates a two-dimensional scatter diagram shown in FIG. 11 (step S17). This two-dimensional scatter diagram includes a relative attenuation index RAI (Relative Attenuation Index) shown on the vertical axis forming one dimension and a quality change index SDI (Spectrum Deformation Index) shown on the horizontal axis forming another dimension. Of course, this scatter diagram only needs to include the relative attenuation index RAI and the quality change index SDI, and may be a three-dimensional or four-dimensional or more scatter diagram combining other indexes.

Among these, the relative attenuation index RAI is a index corresponding to _(pixel distinguished from the S _n of the detector) CT values each pixel S _B forming a three-dimensional volume space, in the present embodiment, the pixel S _B Each relative attenuation index RAI = W1 + W2 + W3 is defined. In contrast, the radiation quality variation index SDI, a index to a substance-specific linear absorption coefficient, in this embodiment, is defined as a radiation quality variation index SDI = W 3 _/ W ₁ of each pixel S _B. More specifically, the relative attenuation index RAI represents the relative attenuation with respect to the reference material an amount based on a sum of at least two energy bands all the pixel values for each pixel S _B. The line quality change index is a line for each pixel S _B based on a division between a pixel value of a higher energy band of at least two energy bands and a pixel value of the same pixel of a lower energy band. Indicates a change in quality. The quality change index SDI may be defined as W ₂ / W ₁ or W ₃ / W ₂ .

Data of the thus pixel S relative attenuation index RAI and radiation quality variation index SDI is computed for each _B is stored in the first storage unit 34.

Next, the data processor 35 reads the data of the relative attenuation index RAI and the radiation quality change index SDI from the first storage unit 34, creates a two-dimensional scatter diagram from these data, and displays the scatter diagram on the display 36. (Step S18). An example of this display is shown in FIG. The drawing the two-dimensional scatter, for each pixel S _B, the corresponding point of the relative attenuation index RAI and radiation quality variation index SDI is displayed by dots.

Next, the data processor 35 analyzes the dispersion state of the dots shown on the scatter diagram, and various substances (soft tissue, gingiva, alveolar bone, cortical bone, tooth, metal, etc.) forming the jaw. Is identified (step S19). Specifically, the scatter diagram of the dot points on the scatter diagram is the scatter diagram previously obtained by CT imaging for each substance in advance (that is, for each device or for each system regardless of the device). A dispersion position with noise added is analyzed from a dispersion table indicating a distribution area, and grouping is performed according to the distribution characteristics. Further, the position of the group of dot groups is compared with reference data in the first storage unit 34 in advance to determine which group of dot groups corresponds to which substance.

Furthermore, display data to the user is generated, for example, by coloring each determined substance with a different color, adding an annotation AN with characters, or indicating a group of dot groups grouped for each substance by ROI. (Step S20). Of course, since the identification result includes fluctuation components such as noise, it is desirable to perform noise reduction processing such as smoothing.

Thereafter, the data processor 35 displays the prepared display data on the display 36 (step S21). An example of this display is shown in FIG. In this example, each substance is indicated by a different symbol, an irregular ROI, and an annotation indicating the substance name.

Note that this display data is preferably generated and displayed in different display modes for each type of identified substance. If expressed in a generalized manner, it is preferable to generate and display one or a plurality of elements of hue, saturation, brightness, and pattern that are different from each other in combination. Among them, displaying with different patterns means, for example, displaying a two-dimensional region identified with a specific substance with diagonal lines, coloring only the boundary surface, and displaying the inside with black dots.

Thus, by creating a scatter plot, or is composed of each pixel S _B is any substance constituting the three-dimensional space becomes an imaging target comprising jaws have been identified.

Therefore, the data processor 35 creates three-dimensional shape data representing the actual jaw using the identification information (step S22), and displays the created three-dimensional shape data on the display 36 (step S23). The three-dimensional shape data is also preferably generated and displayed by combining one or more elements of hue, saturation, brightness, and pattern, as described above. The three-dimensional shape data may be data indicating a physical shape of only a designated substance. This three-dimensional shape data is stored in the first storage unit 34.

Also, the three-dimensional shape data is transmitted to an external dental CAD / CAM system for an impression under the control of the controller 33 in response to an instruction from the operator (step S24).

In the embodiment described above, the processing in steps S1 to S6 in FIG. 9 executed by the controller 33 constitutes scanning means, and the processing in steps S12 and S13 in FIG. 9 executed by the data processor 35 constitutes reconstruction means. is doing. Further, the processing of steps S16 to S19 in FIG. 9 corresponds to the substance identifying means, and among these, the processing of step S17 corresponds to the scatter diagram creating means. The processing in steps S20 to S23 in FIG. 9 corresponds to the image generation means, among which the processing in step S22 constitutes a three-dimensional shape data creation means, and the processing in step S23 constitutes a three-dimensional shape display means. Further, step S24 corresponds to transmission means. Further, a collecting means is formed by the data counting circuit 51n and the step S7 of FIG.

As described above, according to the present embodiment, a scatter diagram is created from the reconstructed CT image for each of the _three energy regions ER ₁ to ER ₃ given to the X-ray photons, and based on this scatter diagram. The type of substance that forms the jaw can be identified. Based on the identification result, three-dimensional shape data of each substance in the jaw can be created.

There are various substances in the jaw from soft tissue to hard tissue with different X-ray absorption characteristics. However, three CT image data reconstructed for each of the three energy regions ER1 to ER3, that is, the relative attenuation index RAI based on the X-ray absorption information obtained by discriminating into the low energy region, the medium energy region, and the high energy region. By comparing the (Relative Attenuation Index) and the quality change index SDI (Spectrum Deformation Index), it is possible to discriminate and identify substances in the jaw more accurately.

This point will be described in detail. The count value of each energy region (W1, W2, W3 in order from the lower energy region) is obtained for each pixel, and the sum of W1 + W2 + W3 (this sum is expressed as RAI relative attenuation index) on the vertical axis. : Relative Attenuation Index, which is an example of normalization with tough bones in panoramic images, but in the case of CT images, it is the same as the definition of the CT value by using water as a reference). W3 / W1 on the horizontal axis (SDI quality change index:
When the scatter diagram is drawn, it is gathered in a certain spot range for each substance in a unique place according to each substance. By classifying this and defining a density value in that class, or assigning a color, it is possible to generate an image classified or colored according to the nature of the substance.

In this method of classifying substances, the SDI elements are added one-dimensionally rather than simply judging from RAI alone. Will improve. Of course, the RAI may be W3 / W2 or W2 / W1 as well as W3 / W1.

From the information with a very wide dynamic range thus obtained, it is possible to optimally display from soft tissue to hard tissue at the same time. It is possible to display them independently according to each substance, or to perform fusion display by changing the color display. In addition, since the type of the substance can be identified with high accuracy, the three-dimensional shape data of soft tissue and hard tissue can be stably obtained with higher accuracy.

This three-dimensional shape data is extremely effective scan data for a dental CAD / CAM system. The CAD / CAM system is a system in which a CAD (Computer-Aided (Assisted) Design) system and a CAM (Computer-Aided (Assisted) Manufacturing) are integrated. That is, a system for designing and manufacturing dental artifacts using computer graphics technology. In the case of a dental CAD / CAM system, measurement of the dentition model and the patient's teeth is necessary, and a measurement model and a three-dimensional measurement model (three-dimensional shape data) of the patient's teeth are constructed from the measurement results. .

However, since it is essential for dental prosthesis to adapt to the abutment tooth and chin on the oral cavity, it is necessary to measure not only the morphological information of the target biological information but also the quantitative information on the hardness of the living body as the premise of CAD. is there. When conventional X-ray CT is used, the dynamic range of the detector is narrow, and it is difficult to image hard tissue and soft tissue simultaneously. In recent years, so-called optical impressions have been frequently used in which an impression is obtained by optically scanning the oral cavity. The optical impression is easily affected by the hand shake of the surgeon, and there is a limit on the properties of light in measuring the complicated hollow form of the measurement object, the overlapping shape of the dentition, and the like. In the creation of a crown, it is necessary to obtain boundary information with a depth of hard tissue and soft tissue in order to determine the height of the crown, which is also limited by the properties of light.

On the other hand, according to the present embodiment, the type of substance constituting the jaw is directly identified from the three-dimensional volume data reconstructed based on the projection data collected by the photon counting detector 22. Based on the identification result, three-dimensional shape data indicating the structure of the jaw is created. As described above, the reconstructed three-dimensional volume data is created based on pixel data having a wide dynamic range from soft tissue to hard tissue. For this reason, the drawing ability of the three-dimensional shape data indicating the soft tissue and hard tissue of the oral cavity is also high. That is, an accurate impression can be made.

In this way, the jaw data necessary for the impression of creating the tooth mold can be obtained directly from the 3D data obtained by CT, and the data necessary for the “CT impression” corresponding to the conventional optical impression. Can be conveniently provided.

In addition, surgical guides for implant surgery that were conventionally designed by CAD / CAM represented by optical impressions can be created from CT images, and fusion with surgical guides can be easily performed with high accuracy from CT images of hard tissues. This is possible only with a CT scanner. This eliminates the need to purchase expensive optical equipment for optical impressions, and can provide a surgical guide in a dental clinic at a low cost in a short time only by purchasing a CT scanner. In addition, when an image that makes the soft tissue appear stable in this way, it can also be used for a diagnosis in which a soft tissue lesion or a positional relationship between the soft tissue and the hard tissue is important, which cannot be obtained by conventional X-ray medical care. Can do. Furthermore, it is possible to automate the preparation of abutment teeth.

In the present embodiment, for example, the highest energy threshold of the four energy thresholds is set to the X-ray tube voltage, and the other three thresholds divide the range of energy information contributing to imaging into three equal parts. Installed and information on the uppermost energy band (This information is information on the number of photons due to the superposition phenomenon because the energy band is a band equal to or higher than the value corresponding to the X-ray tube voltage, and is not worthy of imaging. It is also possible to obtain a highly accurate reconstructed image optimized for soft tissue and hard tissue by reconstructing the image by dividing it into three energy bands without using only (some).

Furthermore, in the present embodiment, a scatter diagram composed of a two-dimensional “relative attenuation index RAI−radiation quality change index SDI” is used to identify the type of substance present in the jaw. Compared to the case where only the relative attenuation index RAI is used, the scatter diagram is used to increase the noise resistance of the substance identification and improve the identification accuracy.

The reason for this will be described with reference to FIGS. In both figures, it is assumed that the coordinates of two substances A and B are now represented by A (SDI _A , RAI _A ) and B (SDI _B , RAI _B ). That is, the relative attenuation index RAI = RAI _{A and the} quality change index SDI = SDI _A of the substance _A , and the relative attenuation index RAI = RAI _{B and the} quality change index SDI = SDI _B of the substance B. However, the coordinates A and B are representative value positions representative of the scattering regions of the substances A and B. The representative values are, for example, the mode value position, the median value position, the arithmetic mean value position, or the center of gravity. Suppose that it is a position.

In the example shown in FIG. 14, the coordinates A (SDI _A , RAI _A ) and B (SDI _B , RAI _B ) are relatively distant from each other with respect to the line quality change index SDI (distance L _SDI ), and can be easily distinguished. The relative attenuation index RAI is very close (distance L _RAI ). in this case,
Distance L _RAI = RAI _A -RAI _B
However, the distance between opposite sides created by the distance L _RAI and the distance L _SDI , that is, the distance L _AB between the coordinates A and B on the scatter diagram is
Distance L _AB = {(RAI _A -RAI _B ) ² + (SDI _A -SDI _B ) ² } ^1/2
And always distance L _AB > distance L _RAI .

The scattering characteristics of the substances A and B are theoretically represented by one point on the scatter diagram. However, in an actual measurement system, there is an influence of noise that includes factors such as photon noise, detector sensitivity unevenness, mechanism operation accuracy, and reconstruction error. For this reason, each scattering characteristic of the substances A and B is not actually a point but is dispersed in a certain range (region). For this reason, when identifying the types of substances A and B, it is better to discriminate them on the scatter diagram including the quality change index SDI than the relative attenuation index RAI, that is, the CT value alone. Since the distance L _AB between B becomes larger, the noise resistance is improved, and therefore the identification (discrimination) accuracy is remarkably improved.

Further, in the example shown in FIG. 15, the relative attenuation index RAI is substantially equal, and RAIA≈RAIB, which is almost indistinguishable on the axis of the relative attenuation index RAI. Even if it is attempted to discriminate between the two substances A and B using only the relative attenuation index RAI, it is difficult to be buried in noise. However, as in this embodiment, by adding the axis of the quality change index SDI, both substances A and B can be distinguished in consideration of the information of the distance LSDI. That is, even if the X-ray absorption coefficients of the substances A and B are very similar, they can be distinguished from each other, and the identification accuracy is improved.

(Modification)
A modification of the detector that detects X-rays will be described with reference to FIG. The detector 122 according to this modification includes a data counting circuit 151 _n for each pixel _Sn (n = 1 to N), that is, for each acquisition channel CN _n . The data counting circuit 151 _n alone constitutes a collecting means.

The data counting circuit 151 _n, as shown in FIG. 16, the energy region distribution circuit 55 is disposed between the four-stage comparator ₅₁ 1-51 ₄ and 4-stage counter ₅₆ 1-56 _4. That is, as in the embodiment described above, comparator ₅₁ 1-51 output ₄ is unlike intact counter ₅₆ 1-56 ₄ respectively input to the configuration, the energy region distribution circuit 55 comparator ₅₁ 1-51 ₄ The energy range ER ₁ to ER ₄ to which the energy associated with the incidence of each X-ray photon belongs is decoded, and the energy range ER ₁ (to ER ₄ to which the energy value belongs. ) Is sent to the counter 56 ₁ (˜56 ₄ ) that is in charge of counting. Thus, only the counter 56 ₁ (˜56 ₄ ) that has received the pulse signal counts up, thereby counting X-ray photons having energy values belonging to the energy region ER ₁ (˜ER ₄ ). The data counting circuit 51 _n of the above-described embodiment calculates (collects) the count value for each energy region of the final X-ray photon by calculation, but the data counting circuit 151 _{n according} to this modification directly calculates it. Can be counted.

A specific configuration will be described. As shown in FIG. 16, the output terminals of the comparators 54 ₁ to 54 ₃ are connected to the energy region distribution circuit 55. This energy region distribution circuit 55 compares the output of the plurality of comparators 54 ₁ to 54 ₄ , that is, the pulse voltage corresponding to the detected energy value of the X-ray particles and the analog amount threshold th ₁ (to th ₄ ). And the energy range ER ₁ to ER ₄ is classified.

Specifically, this distribution is determined by the relationship between the detection voltage V _dec (detected energy value of photons) input to the _four comparators 54 ₁ to 54 ₄ and the threshold values th ₁ to th ₄ . That is, when the detection voltage V _dec <th ₁ to th ₄ , the outputs of all the comparators 54 ₁ to 54 ₄ are turned off. In other words, the output = 0 of the pixel _{S n.} As a result, noise components smaller than the energy threshold TH ₁ defined as the input energy counting limit are not counted. This noise component corresponds to an energy value signal belonging to the non-countable region ERx in FIG.

However, if the detection voltage V _dec exceeds the minimum threshold th ₁ (V _dec ≧ th ₁ ), the number of photons is counted. If the relationship is th ₁ <V _dec ≦ th ₂ (˜th ₄ ), only the output of the first-stage comparator 54 ₁ is turned on, and the energy value of the photon related to the input is in the lower energy region. Decoded as ER ₁ discriminated. If the relationship of th ₂ <V _dec ≦ th ₃ (˜th ₄ ) is established, the outputs of only the first-stage and second-stage comparators 54 ₁ and 54 ₂ are turned on, and the energy value of the photon related to the input is What is discriminated into the medium second energy region ER ₂ is decoded. If th ₃ <V _dec ≦ th ₄ , the outputs of only the first-stage, second-stage, and third-stage comparators 54 ₁ , 54 ₂ , 54 ₃ are turned on, and the photon energy associated with the input value is decrypted to be discriminated in the high third energy region ER _3. Furthermore, if the relationship of th ₄ <V _dec is satisfied, the outputs of all the comparators 54 ₁ to 54 ₄ are turned on, and the energy value of the photon related to the input is not suitable for imaging or counting, the third high It is deciphered that the noise component belongs to the region ER ₄ exceeding the energy region ER ₃ , disturbance, and the like.

The energy region distribution circuit 55 sends a pulse signal indicating the discrimination result to any of the counters 56 ₁ to 56 ₄ in accordance with the discrimination result by the comparators 54 ₁ to 54 ₄ described above. For example, if there is an event to be discriminated in the energy region ER _1, and sends a pulse signal to the counter 56 ₁ in the first stage. If there is an event to be discriminated in the energy region ER _2, it sends a pulse signal to the second-stage counter 56 _2. If events are distinguished in the energy region ER _3, and sends a pulse signal to the counter 56 ₃ of the third stage. If events are distinguished in the energy region ER _4, it sends a pulse signal to the counter 56 ₄ of the fourth stage.

Note that when the photon count value discriminated in the energy region ER ₄ corresponding to the noise is unnecessary information for image processing, substance identification processing, various measurement processing described later, the value can be ignored. . This counter 56 ₄ of the fourth stage may not be provided in consideration.

Other configurations of the detector 122 except for the energy region distribution circuit 55 described above are the same as the configuration of FIG. 5 described above.

Thus, according to this detector 122, by providing the energy region distribution circuit 55, events corresponding to the energy regions ER ₁ to ER ₄ (incident X-ray photons) can be directly counted. For this reason, the above-mentioned subtraction process can be made unnecessary. That is, step S7 of FIG. 9 executed by the data processor 35 described in the above embodiment is not necessary.

Furthermore, it is possible to provide a modification related to the imaging target of the X-ray CT apparatus according to the present invention. In the above-described embodiment and modifications thereof, the imaging target of the X-ray CT apparatus is a living body, that is, a subject (patient's jaw). That is, the patient's head is positioned in the imaging space and the jaw is CT scanned. On the other hand, in this modified example, an impression (material) taken in the oral cavity or a dental model made by pouring a model material such as gypsum into the impression is positioned at a predetermined position in the imaging space, and the impression or CT scan of the dental model. Thereby, the three-dimensional data of a tooth | gear or a dentition can be obtained similarly to CT impression of a biological body. In this case, an X-ray contrast agent is added to a material such as a mold preparation (impression material) or gypsum for creating a dental model. This X-ray contrast agent is a component that enhances the contrast difference with air with respect to at least one of the above-described RAI (relative attenuation index) and SDI (line quality change index) for X-rays. Examples of components of such an X-ray contrast agent include powdered iodine, barium, strontium, titanium oxide, and zinc oxide. Among these components, one or more types are added in advance to the material for making the impression or the tooth model. Note that the CT scan slice width is set as thin as possible. As a result, the enhancement of the contrast by the X-ray contrast agent and the improvement of the resolution by the thin scan can accurately detect the edge of the inner surface shape of the tooth mold and the contour shape of the outer surface of the tooth mold, and more accurate three-dimensional data for the tooth mold Can provide.

Note that the present invention is not limited to the configurations shown in the above-described embodiments and modifications, and can be implemented with various modifications without departing from the spirit described in the claims.

1 X-ray hybrid machine (X-ray imaging device for dental CT imaging and dental panoramic imaging)
13D rotating mechanism (supporting means)
17 Console
15,16 arm (support means)
21 X-ray tube 22, 122 Detector 23 Slit 33 Controller (one of the elements that functionally realizes various data processing and control means)
34 1st memory | storage part 35 Data processor (One of the elements which implement | achieves the means concerning various data processing functionally)
36 Display 37 Input device 40A to 40D ROM
51 _n , 151 _n Data counting circuit 54 Comparator 55 Energy domain allocating circuit 56 Counter 57 D / A converter 58 Latch circuit 59 Serial converter C Semiconductor cell Cp detection circuit S _n pixel DS _i discrimination circuit CN _n data collection channel

Claims

An X-ray tube that emits white X-rays;
A detection circuit having a pixel group in which pixels that output electric pulses corresponding to incident X-ray photons are two-dimensionally arranged, and N or more energy thresholds (N ≧ 3) with respect to the continuous energy of the X-rays And discriminating the electric pulse output from each pixel by the N energy thresholds, and the number of photons of the X-ray is divided by the N energy thresholds and the N energy thresholds A photon counting detector having collection means for collecting each pixel in accordance with each of "N-1" energy bands between each other;
A support means for making the X-ray tube and the detector face each other, and supporting the X-ray tube and the detector around an imaging target so as to be rotatable;
Scanning means for rotating the X-ray tube and the detector around the imaging object based on an instructed scanning method during the CT imaging;
Reconstruction for CT (computed tomography) into projection data based on the number of photons of the X-rays collected by the collecting means for each of at least two energy bands out of the “N−1” energy bands Reconstruction means for performing processing to reconstruct a CT image of each of the at least two energy bands;
An X-ray imaging apparatus comprising:
The apparatus further comprises: a substance identifying means for identifying a kind of a substance forming the imaging target based on a pixel value of the image in each of the at least two energy bands reconstructed by the reconstruction means. Item 2. The X-ray imaging apparatus according to Item 1.
The image generation means which produces | generates the image which shows the structure of the said imaging target based on the pixel value of the said image of each of the said at least 2 energy band reconfigure | reconstructed by the said reconstruction means. 2. The X-ray imaging apparatus according to 2.
The substance identification means includes
A relative attenuation index per pixel that is an amount based on the sum of pixel values of the image in all of the at least two energy bands and represents a relative attenuation with respect to a reference material, and a higher of the at least two energy bands Create a scatter diagram with two or more dimensions including a pixel quality change index for each pixel based on the division between the pixel value of the image in the lower energy band and the pixel value of the image in the lower energy band The X-ray imaging apparatus according to claim 2, further comprising: a scatter diagram creating means.
The X-ray imaging apparatus according to claim 4, wherein the substance identification unit is configured to determine a type name of the substance from the scatter diagram.
The image generating means includes
Three-dimensional shape data creation means for creating three-dimensional shape data classified into substances that are qualitatively similar or of the same quality in the imaging target from the data of the two-dimensional scatter chart created by the scatter diagram creation means The X-ray imaging apparatus according to claim 4, comprising:
The imaging target is a subject's jaw that is a living body,
The image generating means includes
A three-dimensional shape for displaying the three-dimensional shape data created by the three-dimensional shape data creating means in one or more combinations of hue, saturation, brightness, and pattern, which are different from each other for each type of the substance. A display means,
The type of the substance includes at least one of gingiva of the jaw, muscle, tongue, soft tissue including pus in the maxillary sinus, tooth, alveolar bone, cortical bone, and hard tissue which is a metal and / or prosthesis. The X-ray imaging apparatus according to claim 6, further comprising:
The X-ray imaging apparatus according to claim 6, further comprising a transmission unit that transmits the three-dimensional shape data created by the three-dimensional shape data creation unit to an external system.
The detection circuit includes a semiconductor layer formed of a semiconductor material that directly converts the X-rays into the electric pulse and divided for each pixel, a charged electrode stacked on one surface of the semiconductor layer, and the semiconductor A current collecting electrode stacked on the other surface of the layer and divided for each pixel,
The collecting means is built as an ASIC (Application Specific Integrated Circuit) layer as a layer integral with the detection circuit,
The semiconductor material is CdTe, CZT, or TlBr.
The X-ray imaging apparatus according to any one of claims 1 to 8, wherein
The detection circuit includes:
A scintillator array in which a plurality of scintillators processed into columnar shapes are bundled;
A plurality of avalanche photodiodes are mounted on a light receiving surface that is optically connected to the scintillator array and receives light incident from the scintillator, and each rectangular region having a predetermined size corresponding to a cell on the light receiving surface belongs to the region. A silicon photomultiplier having a configuration in which the avalanche photodiode is electrically connected by a quenching element;
The collecting means is built as an ASIC (Application Specific Integrated Circuit) layer as a layer integral with the detection circuit,
The material of the scintillator is LFS (lutetium silicate), LuAG: Pr (praseodymium-added lutetium / aluminum / garnet), or a material having decay time, light emission amount and specific gravity equivalent to the LuAG: Pr,
The X-ray imaging apparatus according to any one of claims 1 to 8, wherein
The X value according to any one of claims 1 to 10, wherein the highest threshold value among the N energy threshold values is set to a desired upper limit value of the energy value of the X-ray photon. Line imaging device.
12. The X-ray imaging apparatus according to claim 11, wherein the desired upper limit value is set to a value belonging to a desired energy range centered on a value corresponding to a tube voltage of the X-ray tube.
Desired energy range for the desired upper limit, when said lowest threshold value of the N energy threshold was set to TH 1, wherein, characterized in that there -TH 1/2 <a desired upper limit <+ TH 1/2 Item 13. The X-ray imaging apparatus according to Item 12.
The detection circuit and the collecting means are formed as a module array in which a plurality of modules of a predetermined size in which the pixels are two-dimensionally arranged are arranged two-dimensionally with a gap between each other. 14. The X-ray imaging apparatus according to any one of 1 to 13.
The support means is configured such that a jaw portion of a subject sitting or lying on a dental chair separate from the X-ray imaging apparatus can be positioned in an imaging space between the X-ray tube and the detector. The X-ray imaging apparatus according to claim 6, wherein the X-ray tube and the detector can be supported.
The part of the plurality of modules of the detector can be used for a panoramic image of a subject's jaw in addition to the CT imaging. X-ray imaging device.
The imaging object is an impression taken in the oral cavity of a subject who is a living body, or a dental model made from the impression, and X is previously applied to a material forming the impression or the dental model. A line contrast agent is added,
The X-ray imaging apparatus according to claim 1, wherein the X-ray contrast medium includes at least a component of iodine, barium, strontium, titanium oxide, or zinc oxide.
An X-ray tube that emits white X-rays;
Photon counting having a detection circuit having a pixel group in which pixels that output electric pulses corresponding to incident X-ray photons are two-dimensionally arranged, and collecting means for collecting the number of X-ray photons for each pixel A mold detector;
The X-ray tube and the detector are opposed to each other, and support means for rotatably supporting the X-ray tube and the detector around the imaging target during CT imaging,
Scanning means for rotating the X-ray tube and the detector around the imaging object based on an instructed scanning method during the CT imaging;
In an X-ray imaging method in an X-ray imaging apparatus comprising:
The CT imaging is performed in a state where N energy thresholds (N ≧ 3) or more are given to the collecting means with respect to the continuous energy of the X-rays, and the electric pulses output from the pixels are collected by the collecting means. Discriminating by the N energy thresholds, the number of photons of the X-ray is divided by the N energy thresholds and “N−1” energy bands between the N energy thresholds. Collect for each pixel according to each,
Reconstruction for CT (computed tomography) into projection data based on the number of photons of the X-rays collected by the collecting means for each of at least two energy bands out of the “N−1” energy bands An X-ray imaging method, comprising: performing processing to reconstruct a CT image of each of the at least two energy bands.