JP6218392B2 - Panorama imaging apparatus and panoramic image reconstruction method - Google Patents

Description

  The present invention detects a transmission X-ray of an imaging part while scanning a desired imaging part of a subject with X-rays, and processes the detected X-ray data by a tomosynthesis method. The present invention relates to a panoramic imaging apparatus and a panoramic image reconstruction method for reconstructing a panoramic image of the cross section.

  In recent years, tomography of a subject by the tomosynthesis method has been actively performed. Although the principle of this tomosynthesis method has been known for a long time (see, for example, Patent Document 1), in recent years, a tomographic method has also been proposed in which it is desired to enjoy the simplicity of image reconstruction based on the tomosynthesis method. (For example, see Patent Document 2 and Patent Document 3). In addition, many examples are seen for dental use (see, for example, Patent Document 4 and Patent Document 5).

  As one application of the tomosynthesis method to dentistry, a panoramic imaging apparatus that obtains a panoramic image in which a curved dentition is developed in a two-dimensional plane is generally put into practical use. This panoramic imaging apparatus normally draws a pair of an X-ray tube and a detector having pixels over a longitudinal width around the oral cavity of a subject to form a constant trajectory along a dentition whose rotation center is assumed. Thus, a mechanism for rotating the rotation center in a complicated manner is provided. A constant value is maintained between the X-ray tube and the detector. The above-described constant trajectory is a trajectory for focusing on a reference tomographic plane (a tomographic plane existing three-dimensionally) set in advance along a dentition regarded as a standard shape and size. During this rotation, X-rays emitted from the X-ray tube are transmitted through the subject at regular intervals and detected as digital frame data by the detector. For this reason, frame data focused on the reference tomographic plane is collected at regular intervals. This frame data is reconstructed by the tomosynthesis method to obtain a panoramic image of the reference tomographic plane.

  Patent Document 6 discloses an example of a panoramic imaging apparatus having an imaging system in which both the X-ray tube and the detector can draw a circular orbit and can be rotated independently of each other. The jaw is positioned in the circular orbit. X-rays irradiated from the X-ray tube are controlled so as to always face the detection surface of the detector.

JP-A-57-203430 JP-A-6-88790 JP-A-10-295680 US Patent Publication US2006 / 0203959 A1 JP2007-136163A International Publication WO2012 / 008492

  In the case of the panoramic imaging apparatus described in Patent Document 6 described above, the distance between the X-ray tube and the detector is not fixed, and the distance changes for each X-ray path irradiated by scanning. Therefore, it is easy to design a desired X-ray path that avoids the cervical vertebrae and the left and right upper jaws as much as possible, but the control of the angular velocity on the circular orbit for rotating the X-ray tube and the detector becomes complicated. Furthermore, since the distance between the detector and the X-ray tube changes for each path of the X-ray beam, the magnification of the subject changes greatly. It is very difficult to control the rotation of the X-ray tube and the detector, which can avoid the complicated control of the angular velocity and suppress the change of the enlargement ratio. For this reason, the reconstruction is tolerated to some extent the change in the enlargement ratio.

  On the other hand, it is possible to easily design an angular velocity that realizes an X-ray path that avoids as much as possible a structure such as a cervical vertebra that is obstructive to dentition imaging. Shift and add processing is complicated. Moreover, in order to perform this reconstruction process more accurately, various calibrations are performed using a phantom in order to grasp the relationship between the distance in the tomographic direction and the amount of shift-and-add, that is, the structure of the imaging space. There is a need. That is, the calibration process is complicated and takes time. In addition, only the collection scenes corresponding to the number of calibrations can be dealt with, and the X-ray tube and the detector can be freely rotated independently of each other, that is, the distance between the X-ray tube and the detector is In order to effectively use the advantages of the folding angle of the variable imaging system, it is necessary to consider that the combination pattern of the X-ray beam path angle and the rotation angle (posture) of the X-ray tube and detector is infinite. Don't be. For this reason, the difficulty of constructing the theory is also a major factor that has hindered its realization.

  The present invention has been made in view of the above circumstances, and makes it easier to obtain panoramic images while taking advantage of an imaging system in which both the X-ray tube and the detector can draw a circular orbit and can be rotated independently of each other. It is an object of the present invention to provide a panorama imaging apparatus and a panorama image reconstruction method that can perform the reconstruction of the above.

  In order to achieve the above object, an X-ray imaging apparatus according to an aspect of the present invention includes an X-ray tube that irradiates X-rays having continuous energy, and an opening that shapes the X-rays into a fan-shaped X-ray beam. A detector including a slit, a detection circuit having a pixel group in which pixels that output electric pulses corresponding to photons of the X-ray beam are two-dimensionally arranged, and the X-ray tube and the detector are always opposed to each other It is possible to support the rotation means independently of each other around the imaging region of the subject, the X-ray tube and the detector always facing each other, and predetermined along the imaging region. Scanning means for controlling the rotation of the X-ray tube and the detector independently of each other so that the X-ray beam is focused on a reference tomographic plane, and scanning the imaging region with the X-ray beam; and Rotation system The fixed distance determined in advance within the range of change in the distance from the X-ray tube to the detector, or the distance at which the detector is actually located within the range of change. In response, the slit control means for controlling the size of the opening, and the electric pulse reflecting the transmission state of the X-ray beam output from the detector during the scanning by the scanning means through the imaging region is frame data. And a data collecting means for collecting at a predetermined cycle.

  As a preferred example, the support means includes a rotation mechanism that mechanically rotates the X-ray tube and the detector along two circular orbits having different diameters. Preferably, the rotation mechanism is configured to rotate the X-ray tube and the detector around the same rotation center.

  More preferably, a storage unit having a memory for storing frame data on the assumption that the detector is virtually arranged at the position of the farthest distance between the X-ray tube and the detector, Mapping means for mapping the frame data collected at each predetermined period by the data collecting means to the memory and obtaining frame data in a state where the detector is virtually arranged at the most distant distance. And reconstructing means for reconstructing a panoramic image of the tomographic section of the jaw based on the frame data mapped by the mapping means.

  According to the present invention, panoramic images can be reconstructed more easily while taking advantage of an imaging system in which both the X-ray tube and the detector can draw a circular orbit and can be rotated independently of each other. A panoramic imaging apparatus and a panoramic image reconstruction method can be provided.

In the accompanying drawings,
FIG. 1 is a perspective view illustrating an overview of a panoramic imaging device according to an embodiment; FIG. 2 is a diagram for explaining an arrangement configuration of an X-ray tube and a detector of the panorama imaging apparatus, FIG. 3 is another diagram illustrating the arrangement configuration of the X-ray tube and the detector of the panoramic imaging apparatus 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 panorama imaging apparatus, FIG. 9 is a diagram for explaining the relationship between the position of the imaging region of the subject, the actual trajectories of the X-ray tube and the detector, and the trajectory for virtually moving the detector at the time of scanning executed in the embodiment; FIG. 10 is a diagram for explaining a known path for each rotation angle of an X-ray beam passing through a dentition while focusing on a reference tomographic plane set in the dentition; FIG. 11 shows a slit control mode 1 in which the distance between the X-ray tube and the detector is changed for each rotation angle, the position on the orbit where the detector is virtually positioned, and the detection for controlling the opening of the slit. The figure explaining the position of the device and the irradiation field of the X-ray beam, FIG. 12 shows a slit control mode 2 in which the distance between the X-ray tube and the detector is changed for each rotation angle, the position on the orbit where the detector is virtually positioned, and the detection for controlling the opening of the slit. The figure explaining the position of the device and the irradiation field of the X-ray beam, FIG. 13 shows a slit control mode 3 in which the distance between the X-ray tube and the detector is changed for each rotation angle, the position on the orbit where the detector is virtually positioned, and the detection for controlling the opening of the slit. The figure explaining the position of the device and the irradiation field of the X-ray beam, FIG. 14 shows a slit control mode 4 in which the distance between the X-ray tube and the detector is changed for each rotation angle, the position on the orbit where the detector is virtually positioned, and the detection for controlling the opening of the slit. The figure explaining the position of the device and the irradiation field of the X-ray beam, FIG. 15 is a flowchart for explaining an outline of processing executed in cooperation between the controller and the data processor; FIG. 16 illustrates the relationship between the focused X-ray beam irradiation field and the size of the detection surface of the detector, and the relationship between the slit opening control and the size of the storage area of the memory mapping the detector frame data. Figure to FIG. 17 is an image diagram schematically illustrating an example of a reconstructed panoramic image; FIG. 18 is a diagram for explaining an enlargement ratio corrected in the present embodiment.

  Embodiments of the present invention will be described below with reference to the accompanying drawings.

  With reference to FIGS. 1-18, one Embodiment of the panoramic imaging device based on this invention is described.

  This panoramic imaging apparatus is configured as a dental apparatus that captures a panoramic image of a subject's jaw (including a dentition). According to this apparatus, a pseudo three-dimensional cross-sectional image of the jaw of the subject (the image itself is a two-dimensional image, depending on the shape of the imaging region such as a dentition, etc., depending on the configuration and functions described below. A cross-sectional image displayed two-dimensionally) can be taken. Although it is configured as a panoramic imaging apparatus according to the present embodiment, it is not necessarily limited to the field of dentistry, and can be applied to various parts such as mammography, otolaryngology imaging, 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 panoramic imaging apparatus 1 according to the present embodiment.

  The panoramic imaging apparatus 1 includes a pedestal 12 on which casters 11 are 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 unit 14 supplies necessary power to each part of the system.

  The panoramic imaging apparatus 1 also 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 extend along the Y-axis direction. When viewed, both are formed in a substantially L shape, and one end of each of the arms 15 and 16 is superposed so as to overlap each other and attached to the side surface of the elevating unit. A rotation mechanism 13D is provided that can rotate the two arms 15 and 16 independently of each other, that is, at different speeds, and an X-ray tube 21 is provided at the distal end of each of the two arms 15 and 16. And a detector 22. A slit (diaphragm) 23 for forming X-rays into a fan shape is disposed on the front surface of the X-ray tube 21 on the X-ray irradiation side. The opening area is controlled by an opening drive unit 23D (see FIG. 8) such as a motor, which will be described later, and the X-ray tube 21 and the arms 15 and 16 by the rotation mechanism 13D and the arms 15 and 16. Support means for supporting the detector 22 so as to be driven independently of each other is configured.

  The X-ray tube 21 is configured as a rotary anode type X-ray tube used for an appropriate anode material such as tungsten. 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 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. The X-ray beam is then transmitted through the jaw JW of the subject P and attenuated, and the transmitted X-ray beam reflecting the attenuated state enters the detector 22.

  At the time of imaging, as shown in FIG. 2, the jaw JW of the subject P is positioned at a predetermined position in a 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 slit 23, then passes through the jaw portion JW (dentition, etc.) and is detected by the detector 22. Since the two arms 15 and 16 are rotationally driven by the rotating mechanism 13D at the time of imaging, the X-ray tube 21 and the detector 22 around the jaw part, respectively, around the jaw part along a predetermined circular orbit. Rotate. 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, the center line in the direction along the YZ plane of the X-ray beam includes an axis T that intersects the center position C in the width direction of the X-ray detection plane (the width in the direction along the YZ plane) at 90 °. The state is referred to as a “facing state” (see FIG. 3). In FIG. 3, a linear position extending in the Z-axis direction from the mechanical rotation center O is defined as a rotation angle θ = 0, and ± rotation directions are set clockwise and counterclockwise from this rotation position.

  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 Rotation (spinning, that is, posture) can be independently performed 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 made to coincide with each other 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 B <b> 1 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 from each other, and are mounted in a predetermined rectangular shape on a substrate (not shown) to form the entire detector 22.

A plurality of detection modules B1 to Bm are arranged in the vertical (X axis) direction (17 in the vertical direction) while providing a constant gap between the individual modules, and the individual modules are arranged in the scanning direction O _Y. The angle θ is set to, for example, about 14 °, and the ratio of the vertical and horizontal lengths created by the plurality of detection modules B1 to Bm is large. The elongated rectangular surface forms the X-ray detection surface 22 A. Since the detection modules B1 to Bm are arranged obliquely, the X-ray detection surface 22A is located inside the individual detection surfaces of the plurality of modules B1 to Bm. Of course, the angle θ may be set to 0 °.

  The structure of the detector 22 having this obliquely arranged detection module and the processing of the detection signal by the subpixel method are known, for example, from International Patent Publication WO 2012/0866648 A1.

  4 is a central axis when the detector 22 itself rotates (rotates).

  Each detection module B1 (~ Bm) is made of a semiconductor material that converts X-rays directly 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 are formed by the entirety of the plurality of detection modules B1 to Bm. The plurality of pixels _{S n} constitutes a pixel group Cp (refer to FIG. 5).

The number of pixels detecting module B1~Bm each is 40 × 40 pixels, the size of each pixel _{S n} is 200 [mu] m × 200 [mu] m, for example. This pixel size is set to a value that allows detection of incident X-rays as a collection of many 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 For example, data is output at a high frame rate of 300 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 a (n = 1~N). Here, each pixel _{S n,} i.e., a route to each of the data counting circuit ₅₁ 1 from each (to 51 _N) of the semiconductor cell C, and 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 12, 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, and the waveform shaping is performed at the subsequent stage of the charge amplifier 52. Circuit 53, multi-stage comparator 54 _i (here i = 1 to 4), multi-stage counter 56 _i (here i = 1 to 4), multi-stage D / A converter 57 _i (here i = 1 to 4) 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 . Analog reference thresholds (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 as shown in FIG. 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. It is determined which value of the analog quantity threshold 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 normally set so as not to detect disturbances, noise caused by circuits such as the semiconductor cell S and the charge amplifier 42, 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 to 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 32, i.e., for each acquisition channels. For this reason, the reference input ends of the comparators 54 _{1 to} 54 ₄ are connected to the output ends of the _four D / A converters 57 _{1 to} 574, respectively. The D / A converter ₅₇ 1-57 ₄ is connected to a threshold receiving end _T 1 via the latch circuit 58 (~T _N), the threshold receiving end _T 1 (~T _N) is the interface 32 of the console 17 It is connected.

The latch circuit 58 latches the threshold values th ₁ ′ to th ₄ ′ of digital quantities given from the threshold value applicator 40 via the interface 31 and the threshold value receiving end T ₁ (to T _N ) at the time of imaging, and corresponding D / are output to a converters ₅₇ 1 to 57 _4. Thus, D / A converter ₅₇ 1-57 ₄ may be provided to each comparator ₅₄ 1-54 ₄ threshold _th 1 to TH ₄ analog amounts commanded as a voltage amount. Each acquisition 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 discriminating 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 the X-ray energy depending on the tube voltage of the X-ray tube 21, and the vertical axis indicates the incidence frequency of the 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. For the continuous spectrum shown in FIG. 7, the first analog amount threshold th ₁ is lower than the region where counting the number of X-ray photons is unnecessary (the region where there is no meaningful X-ray information for counting and circuit noise is mixed). The first energy region ER ₁ of the eye is set corresponding to an energy threshold TH ₁ that can be distinguished. 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 important feature that is matched to the energy value as a count value = 0.

As a result, appropriate discrimination points based on the characteristics and design values of the energy spectrum are defined, and the 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 are respectively connected to a plurality of counters ₅₆ 1 to 56 ₄ of the input terminal.

Each of the counters ₅₆ 1 to 56 ₄ counts up every time the output of the comparator ₅₄ 1-54 ₃ (pulse) is turned on. Thus, the counters ₅₆ 1 (to 56 ₄₎ the accumulated value _W 1 of the X-ray photon number every fixed time with the energy of more than the energy value which is discriminated in the energy region _ER 1 (to Er ₄₎ in charge of the '( to _W-4 ') as it can be counted for each pixel _{S n.}

Specifically, this counting operation is determined by the relationship between the detection voltage V _dec (photon detection energy value) 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 _{~th 4} is a all comparators ₅₄ 1 to 54 ₄ output = 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 their relationship is _{V dec} ≧ th _1, the outputs of all the comparators ₅₄ 1 to 54 ₄ is turned on. That is, all of the counters ₅₆ 1 to 56 ₄ of the count value _{W 1'~W 4} 'is counted up.

If the relationship of V _dec ≧ th _2, the output of the second and subsequent stages of the three comparators ₅₄ 2-54 ₄ is turned on. Thus, three counters ₅₆ 2-56 ₄ count value _{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 region ER ₁ (to ER ₄ ) to be counted by the counters 56 _{1 to} 564. For this reason, 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 ′, and 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 truly desired 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 the first to fourth energy regions _ER 1 to Er ₄ respectively, the actual count value _{W 1'~W 4} ' Is obtained by calculation (subtraction). Therefore, output on the comparator 54 _{1-54 4,} a combination of off now events, namely circuit to decipher whether the incident X-ray photons belongs to which energy region ER1~ER4 becomes unnecessary. 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 to 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 output from the respective counters ₅₆ 1 to 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 and the remaining serial converter ₅₉ 2-59 all acquisition channels _N. 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 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 embodiment, the plurality of detection modules B1 to Bm receive a light incident from the scintillator, which is optically connected to the scintillator array in which a plurality of scintillators processed into columnar shapes are bundled, and the scintillator array. 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 device 36, an input device 37, a calibration calculator 38, a second storage unit 39, first to fourth ROMs 40A to 40D, and a threshold value assigner 41.

  The controller 33 controls the driving of the panoramic imaging device 1 in accordance with a program given in advance to the first ROM 40A. For this control, the command value is sent to the high voltage generator 42 that supplies a high voltage to the X-ray tube 21, the command value is sent to the opening drive unit 23D to change the opening area of the slit 23, and A drive command to the calibration calculator 38 is also included. The first storage unit 34 stores the frame data that is the count value sent from the detector 22 via the interface 31 and the created image data.

  The data processor 35 operates based on a program given in advance to the second ROM 40B under the control of the controller 33. During panoramic photography, the data processor 35 applies a tomosynthesis method based on a known calculation method called shift and add to the frame data stored in the first storage unit 34 by the operation. carry out. Thereby, the panoramic image of the tomographic plane with the oral cavity of the subject P is obtained.

  In addition, as will be described later, the data processor 35 also performs preprocessing applied to the frame data output from the detector 22 and stored in the first storage unit 34.

  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 embedded in advance in the third ROM40C, giving for each energy discriminator circuit for each pixel S _n in the data counting circuit, Calibrate digital quantity threshold for X-ray energy 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 fourth 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 first to fourth ROMs 40A to 40D.

In this embodiment, as known from International Publication No. WO2011 / 142343 (International Application No. PCT / JP2011 / 060731), the structure of the imaging space IS is analyzed using a phantom, and the collection channel CN _{n of the} detector 22 is analyzed. Is calibrated. This calibration is executed at an appropriate timing such as before imaging or during maintenance inspection.

  Specifically, a phantom (not shown) having a marker positioned on a predetermined standard tomographic plane and capable of imaging known position information with X-rays is arranged in the imaging space IS of the panoramic imaging apparatus 1. . X-ray transmission data from the X-ray officer 21 is collected by the detector 22 to create a panoramic image. From the known position information of the marker and the marker position information on the panoramic image, the distance information between the X-ray tube 21 and the detector 22 and the height information of the X-ray tube with respect to the detector are calculated. From this calculation result and collected data, various parameters that define the positional relationship of the X-ray tube, the detector, and the tomographic plane, with the amount of change in the position of the line connecting the X-ray tube and the detector, are calculated. Thereby, parameters necessary for 3D image reconstruction are calibrated. For this reason, the projection direction can be expressed three-dimensionally by grasping the structure of the imaging space IS three-dimensionally. Therefore, as long as the panoramic image is in focus, an accurate panoramic image can be constructed with no distortion or little distortion in the three-dimensionally expressed image.

  Next, an imaging system scan characteristic of the panorama imaging apparatus 1 and processing of frame data collected by the scan will be described.

  In this panorama imaging apparatus 1, pre-processing for passing the frame data output from the detector 22 to reconstruction processing is performed before reconstructing the panoramic image by applying the tomosynthesis method to the frame data. The This preprocessing assumes that the detector 22 rotates along a virtual trajectory preset in the imaging space IS, and the frame data collected by the detector 22 is transferred from the focal point FP of the X-ray officer 21 to the assumed virtual trajectory. The projected frame data is mapped to the memory area M (see FIG. 8). Thereby, it can be assumed that the detector 22 is rotating while facing the X-ray tube 21 along the virtual trajectory.

This virtual trajectory _Tmax is shown in FIG. In the figure, a horseshoe-shaped line SS indicates a locus projected onto the YZ plane of the 3D reference tomographic plane SS set in advance along the dentition TR in the jaw JW of the subject P. Although the surface of the 3D reference tomographic plane SS itself is two-dimensional, it is a tomographic plane (cross section) along the dentition TR that is statistically standard, and therefore exists in a three-dimensional manner. Similarly, a circular line CS schematically shows a locus projected onto the YZ plane of the cervical vertebra in the back of the jaw. The reference tomographic plane SS has a similar horseshoe shape but a plurality of different sizes, considering that there are differences in the size of the jaws of subjects such as adults and children. It is desirable that the tomographic plane is selected.

  In FIG. 9, a trajectory Tx having a rotation angle θ range of slightly over 180 ° is a mechanical real trajectory of the X-ray tube 21 drawn with a radius Dx around the mechanical rotation center O. Similarly, another trajectory Td in which the range of the rotation angle θ is slightly over 180 ° is a mechanical real trajectory of the detector 22 drawn concentrically with the radius Dd (<Dx) around the same mechanical rotation center O. It is. Both the trajectories Tx and Td are the same as those in FIG. 3 described above, and as described above, the trajectories are drawn by the rotation axes AXs and AXd of both the devices 21 and 22. The X-ray tube 21 and the detector 22 irradiate the X-ray tube 21 while moving at different rotational speeds on the orbits Tx and Td facing each other across the mechanical rotation center O. The detector 22 detects the fan-shaped X-ray beam that has passed through the jaw.

In the present embodiment, the virtual trajectory _Tmax described above is created using paths of X-ray beams XB having various angles that pass through the 3D reference tomographic plane SS described above. This will be described with reference to FIG. In the figure, the 3D reference tomographic plane SS set in the dentition TR and the path of the X-ray beam XB determined so as to avoid the reflection of the cervical vertebra CS as an obstacle shadow are shown for each predetermined value of the rotation angle θ. It is shown. The X-ray tube 21 and the detector 22 are rotated along the trajectories Tx and Td independently of each other so as to be focused on the cross section SS.

The distance between each of the trajectories Tx and Td and the mechanical rotation center O, that is, the length “Dx + Dd” obtained by adding the diameters Dx and Dd is the distance R _max where the X-ray tube 21 and the detector 22 are farthest from each other. become. Hereinafter, this is referred to as “maximum separation distance R _max ”. As can be seen from FIG. 10, the path of the X-ray beam XB does not necessarily pass through the mechanical rotation center O. Rather, most paths do not pass through the mechanical rotation center O in order to avoid the cervical vertebra CS as much as possible. Only the path with the rotation angle θ = 0 ° passes through the mechanical rotation center O. The upper limit of the rotation angle θ is ± α ° (for example, α = 115 °), and the path approaches the mechanical rotation center O in the vicinity of the upper limit value ± α °.

Furthermore, in the case of a path other than the path passing through the mechanical rotation center O, the distance between the X-ray tube 21 and the detector 22 (hereinafter, “tube-detector distance” is smaller than the maximum separation distance R _max . The distance between the X-ray tube 21 and the detector 22 is the smallest distance R _min (hereinafter referred to as the “closest distance R _min ”) except for the farthest distance R _max. , And a distance in which the distance between them is appropriately determined between the furthest distance R _max and the closest distance R _min (hereinafter, the “intermediate distance R _int ” is a typical index.

That is, the distance between the tube and the detector is three kinds of distances, that is, the most separated distance R _max ,
It can be selected from the closest distance R _min and the intermediate distance R _int .

Among these, when selecting the furthest distance R _max , as shown in FIG. 10, the distance “Dx + Dd” is set from the actual mechanical trajectory Tx of the X-ray tube 21 in the path of the X-ray beam XB at each rotation angle θ. . As a result, as shown in FIG. 9, the imaging space IS has a bulge outside the actual mechanical trajectory Td of the detector 22 and bulges the trajectory Td outward, and at a rotation angle θ = 0 °. A virtual most spaced trajectory T _max is set that matches the 22 mechanical real trajectories Td.

In the present embodiment, assuming that the actual detector 22 moves along the virtual farthest trajectory _Tmax , the frame data output from the detector 22 is used as the position on the farthest trajectory _Tmax , that is, the X-ray tube. The memory area M is mapped according to the distance from 21 to a position on the same trajectory _Tmax . This mapping can be performed in various ways.
The area of the memory area M is secured, for example, in the first storage unit 34 (see FIG. 8).

The data processor 35 temporarily stores the frame data collected at predetermined intervals from the detector 22 through the interface 31 once in a predetermined storage area of the first storage unit 34. Next, when the data processor 35 receives a mapping command, for example, after collecting all data by scanning or during scanning, the data processor 35 sends frame data stored in a predetermined storage area of the first storage unit 34 from the X-ray tube 21. Mapping is performed in the memory area M in accordance with the distance to each position of the furthest separation trajectory _Tmax . Therefore, this mapping has the meaning of projecting once collected frame data to the position of the distance. For this reason, the size of the memory area M also depends on the size relationship between the size of the detection surface 22A of the detector 22 and the size of the irradiation field of the X-ray beam XB incident thereon. This mapping is executed every time frame data is collected (for example, 300 fps), that is, every rotation angle of the X-ray beam XB corresponding to the timing.

For this reason, the frame data mapped to the memory area M can be regarded as a transmission X-ray detected in a state where the detector 22 is virtually located on the most distant trajectory _Tmax .

  The mapping of the frame data to the memory area M may be performed after scanning (after collecting frame data) and before reconstructing the panoramic image, as will be described later. Although not particularly shown, the mapping may be executed in parallel with scanning. This leads to a reduction in the overall shooting time.

The relationship between the control of the opening area of the slit 23 and the irradiation field of the X-ray beam XB will be described with reference to FIGS. Incidentally, shown in these figures, reference numeral 22min indicated by the phantom line, 22int, 22max the detector 22 closest distance _{R min,} intermediate distance _{R int,} and, that are located farthest distance _{R max.}

In order to control the opening area of the slit 23, there are two control methods, that is, whether the opening area is set to a fixed value or variably controlled. The former is schematically shown in FIGS. The latter is schematically shown in FIG. As can be seen from FIGS. 11 to 14, the position on the trajectory Td where the detector 22 is actually arranged is the distance between the furthest distance R _max and the closest distance R _min as the rotation angle θ changes. Vary between. Taken the position of the closest distance R _min when farthest from the X-ray tube 21, taking the position of the closest distance R _min when closest to the X-ray tube 21.

In the case of the slit control mode 1 shown in FIG. 11, when the detection surface 22A of the detector 22 is virtually located at the maximum separation distance _Rmax , the horizontal width of the rectangular opening of the slit 23 is the virtual maximum width. It is set to a fixed value that matches the horizontal width of the detection surface 22A at the separation distance _Rmax . The lateral direction is a short direction (Y-axis direction in FIG. 4) perpendicular to the longitudinal direction of the detection surface 22A. In this case, the width of the fan-shaped X-ray beam XB passing through the jaw JW of the subject P is narrower than the actual width of the detection surface 22A at any position in the beam propagation direction. For this reason, there is no excessive X-ray exposure to the subject P. Further, the size of the storage area required when virtually mapping to the memory area M matches the pixel size of the detection surface 22A of the detector 22. Of course, the length in the vertical direction of the opening of the slit 23 may be controlled by a fixed value as well as the width.

In the case of the slit control mode 2 shown in FIG. 12, when the detection surface 22A of the detector 22 is virtually located at the closest distance _Rmin , the horizontal width of the opening of the slit 23 is the virtual closest distance R. It is set to a fixed value that matches the horizontal width of the detection surface 22A at _min . In this case, the width of the fan-shaped X-ray beam XB passing through the jaw JW of the subject P matches the actual lateral width of the detection surface 22A at the closest distance _Rmin . However, although the irradiation field of the X-ray beam XB is larger than the actual width of the detection surface 22A at other positions in the beam propagation direction, the tomographic effect is increased and the sensitivity is increased as the irradiation field expands. Of course, the length in the vertical direction of the opening of the slit 23 may be controlled by a fixed value as well as the width.

In the case of the slit control mode 3 shown in FIG. 13, when the detection surface 22A of the detector 22 is virtually located at the intermediate distance _Rint , the lateral width of the opening of the slit 23 becomes the virtual intermediate distance _Rint . It is set to a fixed value that matches the horizontal width of the existing detection surface 22A. In this case, the width of the fan-shaped X-ray beam XB passing through the jaws JW of the subject P is consistent with the width of the actual detection surface 22A at a position intermediate the distance R _int. However, X-rays beam XB radiation field is greater than the width of the actual detection surface 22A at the position of the beam propagation direction the other (intermediate distance R _int large distance position than) or smaller (than the intermediate distance R _int Large distance position). In this case, the advantages of the control modes 1 and 2 described above, that is, the degree of X-ray exposure and the improvement in sensitivity can be balanced. Of course, the length in the vertical direction of the opening of the slit 23 may be controlled by a fixed value as well as the width.

  Furthermore, in the case of the slit control mode 4 shown in FIG. 14, regardless of the rotation angle θ of the X-ray beam XB, the horizontal width and the vertical length of the opening of the slit 23, that is, the area of the opening is determined by the detector. The detection surfaces 22A are controlled so as to coincide with the horizontal and vertical widths. For this purpose, information on the tube-detector distance along the path of the X-ray beam XB at each rotation angle θ is stored in advance in, for example, the first ROM 40A. The controller 33 controls the opening driving unit 23D based on the distance information, and controls the opening area AR (see FIG. 14).

  In the case of the fourth slit control mode, even if the distance between the tube and the detector changes, the irradiation field of the X beam XB incident on the detector 22 is always substantially the same as the edge of the detection surface 22A. For this reason, X-rays are not irradiated outside the detection surface 22A. Therefore, excessive X-ray exposure to the patient is reduced as much as possible. The frame data output from the X-ray device 22 is always data detected by the entire detection surface 22A. For this reason, the detection performance of the detector 22 is always constant and maximized, and detection with good sensitivity is performed.

  Next, image reconstruction processing executed in cooperation with the controller 33 and the data processor 35 in the present embodiment will be described with reference to the flowchart of FIG.

  When the controller 33 determines in step S1 in FIG. 15 that panoramic imaging (imaging) is performed by imaging the jaw portion JW of the subject P, for example, is the panoramic imaging performed with the opening size of the slit 23 fixed? It is determined whether or not (step S2).

  This determination may be performed interactively with the operator, or may be set as a default. When panoramic imaging is performed with the slit opening size fixed (YES in step S2), a fixed value of the opening size is selected (step S3). This selection may also be made interactively with the operator.

As an option for this selection, there is any one of the slit control modes 1, 2, and 3 described above. In the case of the slit control mode 1, irrespective of the rotation angle theta, and the area of the detection surface 22A of the X-ray beam XB detector 22 _max which always assumed to be located farthest distance R _max and irradiation field The opening of the slit 23 is narrowed so as to match. This aperture value is a fixed value and is either the lateral width of the aperture or the lateral width and the length in the vertical direction.

In the case of the slit control mode 2, regardless of the rotation angle θ, the irradiation field of the X-ray beam XB and the area of the detection surface 22A of the detector 22 _min assumed to be always located at the closest distance R _min are obtained. The opening of the slit 23 is narrowed so as to match. This aperture value is a fixed value and is either the lateral width of the aperture or the lateral width and the length in the vertical direction.

Further, in the case of the slit control mode 3, the area of the detection surface 22A of the detector 22 _int that is assumed to be always located at an intermediate distance R _int with the irradiation field of the X-ray beam XB regardless of the rotation angle θ. The apertures of the slits 23 are narrowed so as to match. This aperture value is a fixed value and is either the lateral width of the aperture or the lateral width and the length in the vertical direction.

  On the other hand, if NO is determined in step S2, that is, if it is determined that panoramic imaging is to be performed while changing the opening area of the slit 23 in accordance with the rotation angle θ, the controller 33 stores the X in the first ROM 40A in advance. Data on the tube-detector distance is read for each pass (rotation angle θ) of the line beam XB (step S4). Based on the read data, the movement information of the plate material of the slit 23 for achieving the opening area for each pass is set (step S5). In this opening variable control, the vertical and horizontal sizes of the opening of the slit 23 are changed, but only the horizontal width may be controlled.

  When the information for controlling the opening of the slit 23 is obtained in this way, the controller 33 determines whether to start scanning (step S6). When a scan command is issued (step S6, YES), as described above, the arm 15A, 16B is rotated and rotated, and fixed or variable for each X-ray beam path corresponding to the frame data collection cycle. In step S7A, the jaw JW is scanned by the X-ray beam (step S7). As a result, the frame data detected by the detector 22 is collected in the collection period. The collected frame data is temporarily stored in the first storage unit 34.

In the present embodiment, the structure of the imaging space IS is analyzed using a phantom as is known from International Publication No. WO2011 / 142343 (International Application No. PCT / JP2011 / 060731) at an appropriate timing before scanning. . Thereby, the gain for image reconstruction and the collection channel CN _{n of the} detector 22 are calibrated.

  Next, the controller interactively determines whether to reconstruct the panoramic image, for example, with the operator (step S8). When not reconfiguring, the process waits until a reconfiguration command is issued, but the process may be terminated without waiting. If it is determined to reconfigure, the process is passed to the data processor 35.

As described above, the data processor 35 temporarily stores data in the first storage unit 34 on the assumption that the detector 22 is virtually placed on the virtual trajectory (= the farthest separated trajectory) T _max as pre-processing before reconstruction. The stored frame data is mapped to the specific memory area M (step S8). This mapping is executed for all the frame data collected at a predetermined collection period (for example, 300 fps).

This state is schematically shown in FIG. As shown in the figure, the size of the irradiation field on the detection surface 22A of the detector 22 and the size of the detection surface 22A of the X-ray beam XB focused by the slit 23 in accordance with the slit control modes 1 to 4 Is different. That is, if the case where all the pixels S _n of the detection surface 22A detects X-rays, sometimes only the pixels S _n belonging to a small rectangular area than the maximum area of the detection surface 22A detects X-rays . For this reason, the size of the memory area M can be changed as appropriate.

  In this mapping, the pixel size and the number of pixels may be changed by a subpixel method.

When this preprocessing is completed, the data processor 35 reconstructs a panoramic image along the reference tomographic plane of the dentition based on the preprocessed frame data and displays it on the display 36 (step S10). The tomosynthesis method, that is, shift-and-add processing associated with this reconstruction is executed in accordance with International Publication WO2012 / 008492. Since the calibrated gain of the imaging space IS is held in advance, the actual position of the dentition is reflected by shifting the frame data by the shift amount along this gain and adding the pixel values to each other. A pseudo three-dimensional panoramic image PI _focus is automatically generated (see FIG. 17). This panoramic image PI _focus is called a 3D autofocus image.

The 3D autofocus image PI _focus may use frame data collected in a specific energy region ER1 (˜ER3) or may use those in all energy regions ER1 to ER3.

  As described above, according to the panoramic imaging apparatus 1 according to the present embodiment, an imaging system in which the X-ray tube 21 and the detector 22 can rotate independently of each other so as to draw concentric circles and different circular orbits is employed. ing. When this imaging system is driven, the slit 23 is arranged so that the X-ray irradiation field hits a position of a predetermined fixed distance or variable distance between the maximum value and the minimum value of the tube-detector distance that changes during scanning. The opening area is narrowed down. In addition, the detector 22 exists as if the frame data output from the detector 22 traveling around the actual trajectory is located on the virtual trajectory set corresponding to the maximum value of the tube-detector distance. Thus, the memory M is mapped in an appropriate manner. The mapped frame data is used to reconstruct the panoramic image.

As a result, since the distance between the virtual detector 22 and the X-ray tube 21 that travels around the virtual trajectory (= the most distant trajectory) T _max is always constant, the tomosynthesis used in the conventional panoramic reconstruction. The law can be used as is. That is, as described in International Publication WO2012 / 008492, panoramic image reconstruction is performed even in the case of an imaging system in which the tube-detector distance always changes each time the angle of the X-ray path for each acquisition period changes. The conventional tomosynthesis method based on the premise that the distance between the tube and the detector is constant can be used as it is. Thereby, even in an imaging system in which the distance between the tube and the detector constantly changes, the reconstruction process becomes very simple, and an increase in calculation load on the data processor 35 can be suppressed. In addition, an increase in processing time associated with image reconstruction can be suppressed.

Of course, even if each part of the dentition deviates from the reference tomographic plane, the position where each part actually exists in three dimensions is identified, and a panoramic image I _focus focused on the position is automatically provided.

  By the way, as shown in FIG. 18, the enlargement ratio (D2 / D1) of the tooth that is the subject is generated when the X-ray beam spreads from the X-ray focal point FP in a fan shape and is irradiated. In addition, the magnification varies depending on the angle θ of the X-ray beam XB because the positions of the dentition and the X-ray tube are different for each path. In this regard, as described above, in the present embodiment, the tube-detector distance is always a pseudo constant value regardless of the angle θ of the X-ray beam XB by the mapping process to the memory region M. . Image reconstruction is executed under this “distance = constant”. For this reason, the change in the enlargement ratio can be corrected more easily than the image reconstruction when “distance = change”.

  In the embodiment of the present application, the detector 22 is a photon counting type detector, but the scintillator and the photoelectric element are combined to accumulate an electrical signal for a predetermined time and output frame data. A so-called integral type detector may be used.

  In addition, the panoramic imaging device 1 described above is a device that captures an image in a state where the patient is lying on his / her back on the dental chair (the lying position). However, the panoramic imaging apparatus according to the present invention has an imaging system that rotates with the jaw of the subject positioned between the X-ray tube and the detector, regardless of the imaging in such a posture. For example, such an imaging system may be a device configuration in which such an imaging system is fixedly installed on the back surface of a chair or on a support and a patient is photographed while sitting or standing. Moreover, such an imaging system may be attached to a fixed structure such as a house or a vehicle wall or ceiling. Furthermore, such an imaging system may be configured as a portable unit, and may be configured to perform imaging by placing it on a patient's shoulder or installing behind a general chair.

  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 gist of the claims.

1 Dental panoramic imaging device 13D Rotating mechanism (supporting means)
17 Console
15,16 arm (support means)
21 X-ray tube 22 Detector 23 Slit 33 Controller (one of the elements for functionally realizing various means)
34 1st memory | storage part 35 Data processor (one of the elements which implement | achieves various means functionally)
36 Display 37 Input device 40A-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 (15)

An X-ray tube that emits X-rays having continuous energy;
A slit having an opening for shaping the X-ray into a fan-shaped X-ray beam;
A detector including a detection circuit having a pixel group in which pixels that output electric pulses corresponding to photons of the X-ray beam are two-dimensionally arranged;
Support means for always supporting the X-ray tube and the detector mutually and rotatably supporting the X-ray tube and the detector around the imaging region of the subject independently of each other;
The X-ray tube and the detector are rotated so that the X-ray tube and the detector always face each other and the X-ray beam is focused on a predetermined reference tomographic plane along the imaging region. Scanning means that controls the imaging region with the X-ray beam controlled independently of each other;
The drives the detector, the X-ray tube, a predetermined fixed distance between the X-ray tube that changes while the scan by said scanning means of the maximum value and the minimum value of the distance to the detector or a slit control means for controlling the size of the opening of the slit so that the X-ray irradiation field strikes the position of the variable distance,
Data collecting means for collecting the electrical pulses reflecting the transmission state of the imaging region of the X-ray beam output from the detector during the scan by the scanning means as frame data at a predetermined period. A panoramic imaging device characterized by the above.   The panoramic imaging apparatus according to claim 1, wherein the support means includes a rotation mechanism that mechanically rotates the X-ray tube and the detector along two circular orbits having different diameters.   The panoramic imaging apparatus according to claim 2, wherein the rotation mechanism is configured to rotate the X-ray tube and the detector around the same rotation center. A storage unit having a memory for storing frame data on the assumption that the detector is virtually arranged at the position of the farthest distance between the X-ray tube and the detector;
Mapping in which the frame data collected at the predetermined period by the data collecting unit is mapped to the memory to obtain frame data in a state where the detector is virtually arranged at the farthest distance. Means,
Reconstructing means for reconstructing a panoramic image of a tomographic image of the imaging region based on the frame data mapped by the mapping means;
The panorama imaging apparatus according to claim 1, further comprising:   The slit control means is configured to fix the size of the opening according to the predetermined fixed distance within a change width of the distance from the X-ray tube to the detector. The panoramic imaging device according to any one of the above. Said fixed distance is a distance that belongs to change the width of the distance, the closest distance at which the X-ray tube and the detector closest to each other, when the X-ray tube and the detector are farthest from each other the farthest distance, and, according to claim 5, wherein said either of the predetermined fixed distance between the closest distance and the detector, the X-ray tube, except for the farthest distance Panoramic imaging device. It said slit control means, assuming at each angle of the path of the X-ray beam irradiated from the X-ray tube, the shortest distance to the detector, the farthest distance, or is positioned at the predetermined fixed distance 7. The panorama according to claim 6, wherein the size of the aperture is controlled so that an irradiation field of the X-ray beam coincides with a detection surface of the detector located at the assumed distance when the X-ray beam is irradiated. Imaging device. A storage unit having a memory for storing frame data on the assumption that the detector is virtually arranged at the position of the farthest distance between the X-ray tube and the detector;
Mapping in which the frame data collected at the predetermined period by the data collecting unit is mapped to the memory to obtain frame data in a state where the detector is virtually arranged at the farthest distance. Means, and
The size of the storage area of the memory, the closest distance to the detector, the depending on whether assuming is positioned in any of the farthest distance, and said predetermined fixed distance, and being different from each other The panoramic imaging device according to claim 7. The slit control means is configured to variably control the size of the opening in accordance with a distance where the detector is actually located within a change width of a distance from the X-ray tube to the detector. The panoramic imaging device according to any one of claims 1 to 4. The slit control means controls the size of the opening so that the irradiation field of the X-ray beam coincides with the detection surface of the detector for each angle of the path of the X-ray beam irradiated from the X-ray tube. The panoramic imaging device according to claim 9, which is configured to do so. A storage unit having a memory for storing frame data on the assumption that the detector is virtually arranged at a position of the farthest distance between the X-ray tube and the detector;
The panorama imaging apparatus according to claim 10, wherein a size of a storage area of the memory is set to be the same as a pixel size of a detection surface of the detector. The imaging region is the subject's jaw,
The reconstruction means automatically focuses on the position of each tooth of the dentition existing in the jaw based on the frame data mapped in the memory, and reflects its actual position and shape 2 The panorama imaging apparatus according to claim 4 , configured to create a pseudo three-dimensional panoramic image obtained by three-dimensionally developing a three-dimensional panoramic image. The detector is a photon counting type detector,
N or more energy thresholds (N ≧ 2) are given to the continuous energy of the X-ray beam, and the electric pulse output from each pixel is discriminated by the N energy thresholds, and the X-ray beam 5. The collecting circuit according to claim 1, further comprising: a collecting circuit that collects the number of photons for each pixel according to each of a plurality of energy bands divided by the N energy thresholds and divided by the N energy thresholds. A panoramic imaging apparatus according to claim 1. The support means is configured such that a jaw portion, which is the imaging portion of the subject, who sits or lies on a dental chair separate from the panoramic imaging apparatus is positioned in an imaging space between the X-ray tube and the detector The panoramic imaging apparatus according to any one of claims 1 to 13 , wherein the X-ray tube and the detector can be supported as much as possible. An X-ray tube that emits X-rays having continuous energy;
A slit having an opening for shaping the X-ray into a fan-shaped X-ray beam;
A detector including a detection circuit having a pixel group in which pixels that output electric pulses corresponding to photons of the X-ray beam are two-dimensionally arranged;
Support means for always supporting the X-ray tube and the detector mutually and rotatably supporting the X-ray tube and the detector around the imaging region of the subject independently of each other;
The X-ray tube and the detector are rotated so that the X-ray tube and the detector always face each other and the X-ray beam is focused on a predetermined reference tomographic plane along the imaging region. Scanning means that controls the imaging region with the X-ray beam controlled independently of each other;
The drives the detector, the X-ray tube, a predetermined fixed distance between the X-ray tube that changes while the scan by said scanning means of the maximum value and the minimum value of the distance to the detector Or a slit control means for controlling the size of the opening of the slit so that the X-ray irradiation field hits a variable distance position ;
Panorama including data collection means for collecting the electrical pulses reflecting the transmission state of the imaging region of the X-ray beam output from the detector during the scan by the scanning means as frame data at a predetermined period. In a method for reconstructing a panoramic image in an imaging apparatus,
The supporting means is driven to control the rotational speed of the X-ray tube and the detector independently of each other and to a predetermined reference tomographic plane along the imaging region. Rotate the X-ray beam so that it is in focus,
While the X-ray tube and the detector are rotated around the imaging region, the electric pulse reflecting the transmission state of the X-ray beam output from the detector through the imaging region is used as frame data. collected by the predetermined cycle,
Assuming that the detector is virtually disposed at the position of the farthest distance between the X-ray tube and the detector, the frame data collected every predetermined period is mapped to a memory, respectively. ,
Reconstructing a panoramic image of the tomographic image of the imaging region based on the mapped frame data;
A panoramic image reconstruction method characterized by the above.

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