CN105899137B

CN105899137B - Radiation detector and computed tomography apparatus using the same

Info

Publication number: CN105899137B
Application number: CN201480072945.2A
Authority: CN
Inventors: 赵敏局
Original assignee: Samsung Electronics Co Ltd
Current assignee: Samsung Electronics Co Ltd
Priority date: 2013-11-12
Filing date: 2014-11-12
Publication date: 2020-01-07
Anticipated expiration: 2034-11-12
Also published as: CN105899137A; KR101812658B1; KR20150054666A

Abstract

A radiation detector and a tomographic imaging apparatus using the same are disclosed. The radiation detector includes: a plurality of image pixels each including at least one count pixel and restoring an image. The at least one counting pixel includes: a radiation absorbing layer that converts incident photons into electrical signals; an optical processor to count a number of photons based on the electrical signal transmitted from the radiation absorbing layer. The number of image pixels is less than the number of count pixels.

Description

Radiation detector and computed tomography apparatus using the same

Technical Field

One or more embodiments of the present invention relate to a radiation detector, a tomographic apparatus using the radiation detector, and an X-ray imaging apparatus using the radiation detector, and more particularly, to a radiation detector that counts incident radiation photons to measure the amount of incident radiation, a tomographic apparatus using the radiation detector, and an X-ray imaging apparatus using the radiation detector.

One or more embodiments of the present invention generally relate to a radiation detector and a radiographic imaging system using the radiation detector, such as a Computed Tomography (CT) apparatus, a Positron Emission Tomography (PET) apparatus, a mammography apparatus or a Single Photon Emission Computed Tomography (SPECT) apparatus or an X-ray system, and the like, and more particularly, to a radiation detector that counts incident radiation photons to measure an amount of incident radiation and a radiographic imaging system using the radiation detector.

Background

A medical image processing apparatus is an apparatus for obtaining an internal structure of a subject as an image. A medical image processing apparatus is a non-invasive examination apparatus that exhibits structural details, internal tissues, and fluid flow of the human body. A user such as a doctor diagnoses the health status and diseases of a patient by using medical images output from the medical image processing apparatus.

Representative examples of apparatuses for irradiating radiation onto a patient to photograph an object include a Computed Tomography (CT) apparatus and an X-ray apparatus. Other types of devices and systems also constitute representative examples of fields of use of the invention, such as Positron Emission Tomography (PET) devices, mammography devices, and/or Single Photon Emission Computed Tomography (SPECT).

Among medical image processing apparatuses, reference is made herein to a CT apparatus, by way of example only, because the CT apparatus provides a cross-sectional image of an object and expresses internal structures of the object (e.g., organs such as kidneys, lungs, etc.) so as not to overlap with each other, unlike a general X-ray apparatus. Therefore, the CT apparatus is widely used to accurately diagnose diseases.

The X-ray device is a medical imaging device that obtains an image of an internal structure of a human body by transmitting X-rays that pass through the human body. The X-ray apparatus can simply obtain a medical image of a target object in a shorter time than other medical imaging apparatuses including an MRI apparatus and a CT apparatus. Therefore, the X-ray system is widely used for simple chest photography, abdomen photography, bone frame photography, sinus photography, soft neck tissue photography, and chest photography.

A medical image processing apparatus for irradiating radiation to photograph a subject necessarily includes a radiation detector that detects radiation that has passed through the subject. Furthermore, when the radiation passing through the object is detected fast enough and accurately enough, an accurate medical image may be reconstructed in a subsequent process, e.g. an image processing, based on the radiation detected by the radiation detector. However, in order to achieve sufficient resolution of the reconstructed image, a sufficient amount of radiation that passes through the object impinges on the detector.

Therefore, it is desirable to provide a radiation detector and a medical image processing apparatus that not only detect radiation passing through an object quickly and accurately but also detect a sufficient amount of radiation to enable reconstruction of an image with an appropriate or higher resolution. Integrating these goals has long eluded those skilled in the art of the present invention.

Here, as prior art relating to the novel features proposed by the present disclosure, reference is made to US-2010/282972 for pixel-based sub-pixel division, wherein the sub-pixels belong to a single pixel having different cross-sectional areas to provide a detectable dynamic range of flux density, wherein the disclosure of this publication is limited to detectors based on indirect detection methods. Other prior art publications referred to herein are US-7829860, US-7473902, WO-2008/020379 and US-2005/285043.

Disclosure of Invention

Technical problem

Technical scheme

One or more embodiments of the present invention include a radiation detector that quickly and accurately detects radiation that passes through an object, a tomographic apparatus using the radiation detector, and an X-ray imaging apparatus using the radiation detector.

One or more embodiments of the present invention include a radiation detector that rapidly counts radiation photons that have passed through an object to accurately detect the amount of radiation, a tomographic apparatus using the radiation detector, and an X-ray imaging apparatus using the radiation detector.

Advantageous effects

Exemplary embodiments of the present invention may quickly and accurately detect radiation that passes through an object.

Drawings

These and/or other aspects will become apparent and more readily appreciated from the following description of the described embodiments, taken in conjunction with the accompanying drawings of which:

FIG. 1A is a schematic view of a generic CT system;

FIG. 1B is a diagram illustrating a structure of a CT system according to an embodiment of the present invention;

fig. 2 is a diagram showing a configuration of a communication unit;

fig. 3A is a diagram showing the configuration of an X-ray system;

fig. 3B is a diagram showing a stationary type X-ray apparatus;

FIG. 3C is a diagram showing a mobile X-ray device;

FIG. 4 is a diagram illustrating a radiation detector according to an embodiment of the present invention;

fig. 5 is a diagram for describing a plurality of pixels of fig. 4;

fig. 6 is a diagram for describing the count pixel of fig. 4;

fig. 7A is another diagram for describing the count pixel of fig. 4;

fig. 7B is another diagram for describing the count pixel of fig. 4;

fig. 8 is another diagram for describing the count pixel of fig. 4;

fig. 9 is a diagram illustrating a computed tomography apparatus according to an embodiment of the present invention.

Best mode for carrying out the invention

Additional aspects will be set forth in part in the description which follows, and in part will be obvious from the description, or may be learned by practice of the presented embodiments.

According to one or more embodiments of the present invention, a radiation detector for sensing radiation includes a plurality of image pixels each including at least one count pixel and restoring an image, wherein the at least one count pixel includes: a radiation absorbing layer that converts incident photons into electrical signals; a photon processor that counts a number of photons based on the electrical signal transmitted from the radiation absorbing layer, and the number of image pixels is less than the number of count pixels.

The at least one counting pixel may count a number of photons smaller than a number of photons incident on the corresponding image pixel.

The optical processor can count the number of photons based on the electrical signal according to a direct method of directly converting incident photons into electrical charge to detect photons.

The plurality of image pixels may each correspond to one pixel value constituting the image.

The plurality of image pixels may each include a plurality of the count pixels.

The plurality of image pixels may each be a pixel for calculating one pixel value included in the image based on the number of photons counted by the plurality of count pixels.

The optical processor may include a count memory that counts and stores a number of photons that is smaller than a number of photons incident on a corresponding image pixel within a specific time.

The optical processor may include: a comparator to compare the electrical signal to a reference value to determine whether the electrical signal exceeds the reference value; and a count memory for counting and storing the number of photons exceeding the reference value based on the comparison result of the comparator.

The at least one counting pixel may include: and a count memory for counting and storing the number of photons smaller than the number of photons incident on the corresponding image pixel within a specific time.

When the plurality of image pixels each correspond to a pixel of the radiation detector, the at least one counting pixel included in the pixel may be divided into at least one counting pixel group, and the number of photons counted by the at least one counting pixel group may correspond to one image pixel value in the image.

The number of counted pixel groups may be equal to or greater than the number of pixels.

The size of the at least one group of counted pixels may be equal to or less than the size of the pixels.

When the plurality of image pixels each correspond to a pixel of the radiation detector, a plurality of count pixels included in a plurality of adjacent pixels may be divided into at least one count pixel group, and the number of photons counted by each of the plurality of count pixel groups may correspond to one image pixel value in the image.

The radiation detector may be a radiation detector for generating a tomographic image.

The radiation detector can sense radiation emitted from an X-ray source attached to a gantry and rotating and passing through a subject.

The radiation detector may be a radiation detector for generating an X-ray image.

The radiation detector may sense radiation emitted from an X-ray source attached to the mobile device and adjusted in position and passing through the object.

The radiation absorbing layer can be formed from cadmium telluride (CdTe).

According to one or more embodiments of the present invention, a radiation detector includes a plurality of pixels that sense radiation. The plurality of pixels each include a plurality of count pixels that sense radiation used to restore an image, wherein the plurality of count pixels each include: a radiation absorbing layer that converts incident photons into electrical signals; an optical processor to count a number of photons based on the electrical signal.

The optical processor may include a count memory that stores a count value.

The plurality of counting pixels may each include counting a number of photons smaller than a number of photons incident on the corresponding pixel.

The radiation detector may be used to generate tomographic images.

The plurality of pixels may each absorb and count two hundred million or more photons per second.

The plurality of pixels may each include 24, 25, or 36 count pixels.

The number of photons counted by the count pixel may correspond to an image pixel value in the image.

The total number of photons counted by a counting pixel group may correspond to one image pixel value in the image, wherein the counting pixel group includes a plurality of counting pixels included in the pixels and arranged adjacent to each other.

The plurality of count pixels included in the plurality of adjacent pixels may be divided into a plurality of groups, and a total number of photons counted by each of the plurality of groups corresponds to one image pixel value in the image.

The optical processor may further include: a comparator to compare the electrical signal to a reference value to determine whether the electrical signal exceeds the reference value; and a count memory for counting and storing the number of photons exceeding the reference value.

According to one or more embodiments of the present invention, a tomographic imaging apparatus includes: a radiation detector including a plurality of image pixels each including at least one count pixel and restoring an image; an image processor to reconstruct a tomographic image based on the number of photons sensed by the radiation detector, wherein the at least one counting pixel includes: a radiation absorbing layer that converts incident photons into electrical signals; a photon processor that counts the number of photons based on an electrical signal transmitted from a radiation absorbing layer, and the number of image pixels is less than the number of count pixels.

According to one or more embodiments of the present invention, an X-ray imaging apparatus includes: a radiation detector including a plurality of pixels each including at least one count pixel and restoring an image; an image processor to reconstruct a tomographic image based on the number of photons sensed by the radiation detector, wherein the at least one counting pixel includes: a radiation absorbing layer that converts incident photons into electrical signals; a photon processor that counts the number of photons based on an electrical signal transmitted from a radiation absorbing layer, and the number of pixels is less than the number of counted pixels.

Detailed Description

The present application claims the benefit of korean patent application nos. 10-2013-.

Reference will now be made in detail to exemplary embodiments thereof as illustrated in the accompanying drawings, wherein like reference numerals refer to like elements throughout. In this regard, the present embodiments may take different forms and should not be construed as limited to the description set forth herein. Accordingly, the embodiments are described below by referring to the drawings only to explain aspects of the present description. When a statement such as at least one of the "follows a column of elements, that statement modifies the entire column of elements rather than a single element within the column.

Advantages and features of one or more embodiments of the present invention and methods of practicing the invention may be understood more readily by reference to the following detailed description of embodiments and the accompanying drawings. In this regard, the present embodiments may take different forms and should not be construed as limited to the description set forth herein. Rather, these embodiments are provided so that this disclosure will be thorough and complete, and will fully convey the concept of the embodiments to those skilled in the art, and the present invention will only be defined by the appended claims. Like reference numerals refer to like elements throughout the specification.

Hereinafter, terms used in the specification will be briefly defined, and embodiments will be described in detail.

All terms (including descriptive or technical terms) used herein should be understood to have meanings apparent to those of ordinary skill in the art. However, the terms may have different meanings according to intentions, precedent cases, or appearance of new technologies of those of ordinary skill in the art. In addition, some terms may be arbitrarily selected by the applicant, and in this case, the meaning of the selected terms will be specifically described in the detailed description of the present invention. Therefore, the terms used herein must be defined based on their meanings together with the description throughout the specification.

When a component "comprises" or "comprising" an element, the component may include other elements but not exclude the other elements unless there is a specific description to the contrary. Furthermore, the term "unit" in the embodiments of the present invention means a software component or a hardware component such as a Field Programmable Gate Array (FPGA) or an Application Specific Integrated Circuit (ASIC) and performs a specific function. However, the term "unit" is not limited to software or hardware. The "unit" may be formed to reside in an addressable storage medium or may be formed to operate one or more processors. Thus, for example, the term "unit" may refer to a component (such as a software component, an object-oriented software component, a class component, and a task component) and may include a process, a function, an attribute, a program, a subroutine, a program code segment, a driver, firmware, microcode, a circuit, data, a database, a data structure, a table, an array, or a variable. The functionality provided by the described components and "units" may be associated with a fewer number of components and "units" or may be divided into additional components and "units".

Reference will now be made in detail to exemplary embodiments thereof as illustrated in the accompanying drawings. In this regard, the present embodiments may take different forms and should not be construed as limited to the description set forth herein. In the following description, well-known functions or constructions are not described in detail since they would obscure the embodiments with unnecessary detail.

When a statement such as at least one of the "follows a column of elements, that statement modifies the entire column of elements rather than a single element within the column.

Throughout the specification, "image" may mean multi-dimensional data formed from discrete image elements (e.g., pixels in a two-dimensional (2D) image and voxels in a three-dimensional (3D) image). For example, the image may include a medical image of the object captured by a Computed Tomography (CT) image capture device.

Throughout the specification, a "CT image" may mean an image generated by synthesizing a plurality of X-ray images obtained by photographing an object while a CT image capturing apparatus is rotated about at least one axis relative to the object.

Throughout the specification, an "object" may include a human, an animal, or a part of a human or an animal. For example, the object may include an organ (such as liver, heart, uterus, brain, chest, abdomen, etc.) or a blood vessel. Further, the object may include a model. The model means a material having a volume very close to the density and effective atomic number of an organism, and may include a spherical model having characteristics similar to those of a human body.

Throughout the specification, a "user" may be, but is not limited to, a medical professional (including a doctor, a nurse, a medical laboratory technician, a medical image specialist) and a technician repairing a medical device.

Since the CT system can provide a cross-sectional image of an object, the CT system can express internal structures of the object (e.g., organs such as kidneys, lungs, etc.) without overlapping each other, compared to a general X-ray capturing apparatus.

The CT system may acquire a plurality of pieces of image data having a thickness of not more than 2mm several tens to several hundreds times per second and then may process the plurality of pieces of image data, so that the CT system may provide a relatively accurate cross-sectional image of the object. According to the prior art, only horizontal cross-sectional images of an object are available, but this problem has been overcome due to various image reconstruction methods. Examples of 3D image reconstruction methods are:

-a mask surface display (SSD) method: the SSD method is the original 3D imaging method, displaying only voxels with a predetermined Hounsfield Unit (HU) value.

Maximum Intensity Projection (MIP)/minimum intensity projection (MinIP) method: the MIP/MinIP method is a 3D imaging method that displays only voxels having the largest or smallest HU value among voxels constituting an image.

-Volume Rendering (VR) method: the VR method is an imaging method capable of adjusting the color and transparency of voxels constituting an image according to a region of interest.

Virtual endoscopic method: the method allows endoscopic observation in a 3D image reconstructed by using a VR method or an SSD method.

-multi-planar reconstruction (MPR) method: the MPR method is used to reconstruct images into different cross-sectional images. The user may reconstruct the image in each desired direction.

-an editing method: the method involves editing neighboring voxels to allow the user to easily view the region of interest while volume rendering.

Voxel of interest (VOI) method: the VOI method displays only selected regions at the time of volume rendering.

A CT system 20 according to an embodiment of the present invention will now be described with reference to fig. 1A. CT system 20 may include devices having various forms. As an alternative to the illustrated CT system 20, the present invention may be associated with other radiographic imaging systems. For example, the invention may relate to an X-ray system, a Positron Emission Tomography (PET) apparatus, a mammography device, or a Single Photon Emission Computed Tomography (SPECT). In which an image of the object can also be generated or reconstructed on the basis of the detection of the emitted and more or less transmitted radiation, which detection can be obtained by a radiation detector that is sensitive to the energy level and that can count photons.

Fig. 1A schematically illustrates a CT system 20 by way of example of an embodiment. Referring to FIG. 1A, CT system 20 may include a gantry 172, a table 175, an X-ray generation unit 176, and an X-ray detection unit 178.

The gantry 172 may include an X-ray generation unit 176 and an X-ray detection unit 178.

The object 10 may be located on the table 175.

The table 175 may be moved in a predetermined direction (e.g., at least one of up-down-left-right directions) during a CT imaging procedure. In addition, the table 175 may be tilted or rotated in a predetermined direction by a predetermined angle.

The stage 172 may also be tilted by a predetermined angle in a predetermined direction.

Fig. 1B is a diagram showing the structure of CT system 20.

CT system 20 may include a gantry 172, a table 175, a control unit 188, a storage unit 194, an image processing unit 196, an input unit 198, a display unit 191, and a communication unit 192 that may enable communication to server 134 and the like.

As described above, the object 10 may be positioned on the table 175. In the present embodiment, the table 175 may be moved in a predetermined direction (e.g., at least one of up, down, right, and left directions) by the control of the control unit 188.

The gantry 172 can include a rotating gantry 174, an X-ray generation unit 176, an X-ray detection unit 178, a rotational drive unit 180, a Data Acquisition System (DAS)186, and a data transmission unit 190.

The gantry 172 may include a rotating frame 174 having a ring shape that is rotatable with respect to a predetermined rotation axis RA. Further, the rotating frame 174 may have a disk shape.

The rotating frame 174 may include an X-ray generation unit 176 and an X-ray detection unit 178 that face each other to have a predetermined field of view FOV. Rotating frame 174 may also include an anti-scatter grid 184. An anti-scatter grid 184 may be located between the X-ray generation unit 176 and the X-ray detection unit 178.

In medical imaging systems, the X-ray radiation reaching the detector (or photosensitive film) includes not only attenuated primary radiation forming the useful image, but also scattered radiation that degrades the quality of the image. In order to transmit primary radiation and attenuate scattered radiation, an anti-scatter grid 184 may be positioned between the patient and the detector (or photosensitive film).

For example, anti-scatter grid 184 may be formed by alternately stacking strips of lead foil and a void material (such as a solid polymer material, a solid polymer, or a fiber composite). However, the formation of anti-scatter-grid 184 is not limited thereto.

The rotating frame 174 may receive a driving signal from the rotation driving unit 180 and may rotate the X-ray generating unit 176 and the X-ray detecting unit 178 at a predetermined rotation speed. The rotating frame 174 may receive a driving signal and power from the rotary driving unit 180 while the rotating frame 174 contacts the rotary driving unit 180 via a slip ring (not shown). In addition, the rotating frame 174 may receive a driving signal and power from the rotating driving unit 180 via wireless communication.

The X-ray generating unit 176 may receive voltage and current from a Power Distribution Unit (PDU) (not shown) via a slip ring (not shown) and a high voltage generating unit (not shown), and may then generate and emit X-rays. When the high voltage generating unit applies a predetermined voltage (hereinafter, referred to as "tube voltage") to the X-ray generating unit 176, the X-ray generating unit 176 may generate X-rays having a plurality of energy spectra corresponding to the tube voltage.

The X-rays generated by the X-ray generation unit 176 may have a predetermined form due to the collimator 182 and then may be emitted.

The X-ray detection unit 178 may be disposed to face the X-ray generation unit 176. The X-ray detection unit 178 may include a plurality of X-ray detection devices. The plurality of X-ray detection devices may each establish a channel, although one or more embodiments of the invention are not limited in this respect.

The X-ray detection unit 178 may detect the X-rays generated by the X-ray generation unit 176 and transmitted via the object 10, and may generate an electrical signal corresponding to the intensity of the detected X-rays.

The present invention relates to a direct type X-ray detection unit 108 for detecting radiation by directly converting the radiation into electric charges. The direct type X-ray detector may use a photon counting detector. The DAS 186 may be connected to the X-ray detection unit 178. The electrical signals generated by the X-ray detection unit may be collected by the DAS 186, either wired or wirelessly.

Depending on the layer thickness or the number of layers, only some of the pieces of data collected by the X-ray detection unit 178 may be provided to the image processing unit 196 via the data transmission unit 190, or the image processing unit 196 may select only some of the pieces of data.

The digital signal may be supplied to the image processing unit 196 via the data transmitting unit 190. The digital signal may be provided to the image processing unit 196 by wire or wirelessly.

Control unit 188 may control the operation of each of the described components, elements, assemblies, or modules within CT system 20. For example, the control unit 188 may control the operations of the table 175, the rotation driving unit 180, the collimator 182, the DAS 186, the storage unit 194, the image processing unit 196, the input unit 198, the display unit 191, the communication unit 192, and the like.

The image processing unit 196 may receive data (e.g., pure data prior to a processing operation) obtained from the DAS 186 via the data sending unit 190, and may perform preprocessing.

The preprocessing may include a process of correcting sensitivity irregularities between channels, a process of correcting signal loss due to a rapid drop in signal intensity or due to an X-ray absorbing material (such as metal, etc.).

The data output from the image processing unit 196 may be referred to as "raw data" or "projection data". The projection data and image capturing conditions (e.g., tube voltage, image capturing angle, etc.) during the acquisition of the data may be stored in the storage unit 194 in common.

The projection data may be a set of data values corresponding to the intensity of X-rays from the assembly of the X-ray source 106 and the collimator 182 that pass through the object 10 and are detected by the detection unit 108. For convenience of description, it is assumed that a group of pieces of projection data obtained from all channels at the same time from the same image capturing angle is referred to as a projection data set.

The storage unit 194 may include at least one storage medium selected from among a flash memory type storage medium, a hard disk type storage medium, a micro multimedia card storage medium, a card type memory (e.g., SD card, XD memory, etc.), a Random Access Memory (RAM), a Static Random Access Memory (SRAM), a Read Only Memory (ROM), an Electrically Erasable Programmable Read Only Memory (EEPROM), a Programmable Read Only Memory (PROM), a magnetic memory, a magnetic disk, and an optical disk.

The image processing unit 196 may reconstruct a cross-sectional image for the object 10 by using the projection data set. The cross-sectional image may be a 3D image. In other words, the image processing unit 196 may reconstruct a 3D image of the object 10 by using a cone beam reconstruction method or the like based on the projection data set.

The input unit 198 can receive external inputs for tomographic imaging conditions, image processing conditions, and the like. For example, the tomographic imaging conditions may include tube voltage, energy value settings for a plurality of X-rays, selection of an image capture protocol, selection of an image reconstruction method, setting of FOV area, number of layers, layer thickness, parameter settings for image post-processing, and the like. Further, the image processing conditions may include the resolution of the image, the attenuation coefficient setting for the image, the setting of the image combination ratio, and the like.

The input unit 198 may include a device for receiving a predetermined input from an external source. For example, the input unit 198 may include a microphone, a keyboard, a mouse, a joystick, a touch pad, a touch pen, a voice recognition device, a gesture recognition device, and the like.

The display unit 191 may display the tomographic image reconstructed by the image processing unit 196.

The exchange of data, power, etc. between the aforementioned elements may be performed by using at least one of wired communication, wireless communication, and/or optical communication.

The communication unit 192 may perform communication with an external device, an external medical apparatus, or the like via the server 134 or the like. The communication will now be described with reference to fig. 2.

Fig. 2 is a diagram showing the structure of the communication unit 192.

The communication unit 192 may be connected to the network 15 by wire or wirelessly and thus may perform communication with the server 134, the external medical device 136, or the external portable device 138. The communication unit 192 may exchange data with a hospital server or other medical equipment in a hospital connected via a Picture Archiving and Communication System (PACS). Further, the communication unit 192 may perform data communication with the portable device 138 or the like according to digital imaging and communications in medicine (DICOM) standard.

The communication unit 192 may transmit and receive data related to diagnosing the subject 10 via the network 15. Further, the communication unit 192 may transmit and receive medical images obtained from an external medical device 136, such as a Magnetic Resonance Imaging (MRI) device, an X-ray device, or the like.

In addition, the communication unit 192 may receive a diagnosis history or medical treatment schedule of the patient from the server 134 and may use the diagnosis history or medical treatment schedule for clinical diagnosis of the patient. Further, the communication unit 192 may perform data communication not only with the server 134 or the external medical device 136 in the hospital but also with the portable device 138 of, for example, a user or a patient. The communication unit 192 may also obtain control parameters as described above in relation to the input unit 198, and thus, in embodiments the input unit 198 and the communication unit 192 may be combined.

Further, the communication unit 192 may transmit information on a device error, information on a quality control state, and the like to a system manager or a service manager via the network 15, and may receive feedback corresponding to the information.

Fig. 3A is a diagram showing the configuration of the X-ray system 1000.

Referring to fig. 3A, an X-ray system 1000 includes an X-ray device 100 and a table 110. The X-ray device shown in fig. 3A may be a stationary type X-ray device or a mobile X-ray device. The X-ray apparatus 100 may include an X-ray irradiation unit 120, a high voltage generator 121, a detector 130, a manipulation unit 140, and a control unit 150. The control unit 150 may control the overall operation of the X-ray device 100.

The high voltage generator 121 generates a high voltage for generating X-rays and applies the high voltage to the X-ray source 122.

The X-ray irradiation unit 120 includes: an X-ray source 122 receiving a high voltage applied from the high voltage generator 121 to generate and irradiate X-rays; a collimator 123 for guiding a path of the X-ray irradiated from the X-ray source 122 to adjust an irradiation area of the X-ray.

The X-ray source 122 includes: an X-ray tube may be implemented as a diode comprising a cathode and an anode. The inside of the X-ray tube is set to a high vacuum state of about 10mmHg, and a filament of an anode is heated to a high temperature to generate thermal electrons. The filament may be a tungsten filament, and a voltage of about 10V and a current of about 3 to 5A may be applied to wires connected to the filament to heat the filament.

In addition, when a high voltage of about 10 to about 300kvp is applied between the cathode and the anode, thermal electrons are accelerated to impinge on a target material of the cathode, and then X-rays are generated. The X-rays are irradiated to the outside through a window, which may be formed of a beryllium thin film. Here, most of the energy of the electrons colliding with the target material is consumed as heat, and the remaining energy is converted into X-rays.

The cathode is formed primarily of copper, and the target material is disposed opposite the anode. The target material may be a high resistance material such as Cr, Fe, Co, Ni, W, or Mo. The target material may be rotated by a rotating field. When the target material is rotated, an electron impact area increases, and a heat accumulation rate per unit area may be increased to 10 times as much as that in the case where the target material is fixed.

The voltage applied between the cathode and the anode of the X-ray tube is referred to as "tube voltage", which is applied from the high voltage generator 121, and the magnitude of which can be expressed by a peak value (kvp). When the tube voltage increases, the velocity of thermal electrons increases, and therefore, the energy of X-rays (the energy of photons) generated when thermal electrons collide with the target material also increases. The current flowing in the X-ray tube is referred to as "tube current", wherein the tube current may be expressed as an average value (mA). When the tube current increases, the number of thermoelectrons emitted from the filament increases, and therefore, the dose of X-rays (the number of X-ray photons) generated when the thermoelectrons collide with the target material increases.

Therefore, the energy of the X-rays can be adjusted according to the tube voltage, and the intensity or dose of the X-rays can be adjusted according to the tube current and the X-ray exposure time.

The detector 130 detects the X-rays irradiated from the X-ray irradiation unit 120 and having passed through the object. The detector 130 may be a digital detector. The detector 130 may be implemented using a Thin Film Transistor (TFT) or a Charge Coupled Device (CCD). In fig. 3A, the detector 130 is illustrated as being included in the X-ray apparatus 100, but the detector 130 may be an X-ray detector as a separate apparatus detachably connected to the X-ray apparatus 100.

Furthermore, the X-ray device 100 may further comprise a manipulation unit 140 providing an interface for manipulating the X-ray device 100. The manipulation unit 140 may include an output unit 141 and an input unit 142. The input unit 142 may receive a command for manipulating the X-ray apparatus 100 and various information regarding X-ray photographing from a user. The control unit 150 may control or manipulate the X-ray device 100 based on information input to the input unit 142. The output unit 141 may output a sound indicating photographing related information (such as irradiation of X-rays) under the control of the control unit 150.

The table 110 and the X-ray device 100 may be connected to each other by wire or wirelessly. If the table 110 and the X-ray device 100 are wirelessly connected to each other, means (not shown) for synchronizing the clocks with each other may be further included. The table 110 may be arranged in a space physically separated from the X-ray device 100.

The table 110 may include an output unit 111, an input unit 112, and a control unit 113. The output unit 111 and the input unit 112 provide the user with an interface for manipulating the table 110 and the X-ray device 100. The control unit 113 may control the table 110 and the X-ray device 100.

The X-ray device 100 may be controlled by the table 110 and the X-ray device 100 may be controlled by a control unit 150 comprised in the X-ray device 100. Accordingly, the user may control the X-ray apparatus 100 through the table 110 or may control the X-ray apparatus 100 by using the manipulation unit 140 and the control unit 150 included in the X-ray apparatus 100. In other words, the user can remotely control the X-ray device 100 via the table 110 or can directly control the X-ray device 100.

In fig. 3A, the control unit 113 of the table 110 and the control unit 150 of the X-ray apparatus 100 are shown separately, but fig. 3A is only an example. As another example, the control unit 113 and the control unit 150 may be implemented as one integrated control unit, which may be included in only a selected one of the table 110 and the X-ray device 100. In the following, the control unit 113 and the control unit 150 represent the control unit 113 of the table 110 and/or the control unit 150 of the X-ray device 100.

The output unit 111 and the input unit 112 of the table 110 and the output unit 141 and the input unit 142 of the X-ray device 100 may each provide an interface for a user to manipulate the X-ray device 100. In fig. 3A, it is illustrated that the table 110 includes the output unit 111 and the input unit 112 and the X-ray radiation unit 100 includes the output unit 141 and the input unit 142, but the embodiment is not limited thereto. As another example, the output unit or the input unit may be included in only a selected one of the table 110 and the X-ray apparatus 100.

Hereinafter, the input unit 112 and the input unit 142 represent the input unit 112 of the table 110 and/or the input unit 142 of the X-ray device 100, and the output unit 111 and the output unit 141 represent the output unit 111 of the table 110 and/or the output unit 141 of the X-ray device 100.

The input unit 112 and the input unit 142 may each include, for example, a keyboard, a mouse, a touch screen, a voice recognizer, a fingerprint recognizer, and an iris recognizer, and may include input devices well known to those of ordinary skill in the art. The user may input a command for irradiating X-rays via the input unit 112 and the input unit 142, and for this, the input unit 112 and the input unit 142 may each include a switch for inputting the command. The switch may be configured such that an irradiation command for irradiating X-rays may be input only when the switch is pressed twice.

That is, when the user presses the switch, a preparation command for performing a preheating operation for X-ray irradiation may be input through the switch, and then, when the user presses the switch again, an irradiation command for irradiating X-rays may be substantially input through the switch. When the user manipulates the switch as described above, the control unit 113 and the control unit 150 each generate a signal (i.e., a preparation signal) corresponding to a command input through the switch manipulation and output the generated signal to the high voltage generator 121 generating a high voltage for generating X-rays.

The high voltage generator 121 starts a warm-up operation when the high voltage generator 121 receives the ready signal output from the control unit 113 and the control unit 150, and the high voltage generator 121 outputs the ready signal to the control unit 113 and the control unit 150 when the warm-up is completed. In addition, the detector 130 also needs to be prepared for detecting X-rays, and therefore, the control unit 113 and the control unit 150 transmit a preparation signal to the detector 130 at the same time as the preheating process of the high voltage generator 121 so that the detector 130 is prepared for detecting X-rays passing through the object. When receiving the preparation signal, the detector 130 prepares for detecting the X-rays, and when the preparation for detection is completed, the detector 130 transmits a detection ready signal to the control unit 113 and the control unit 150.

When the warm-up operation of the high voltage generator 121 is completed and the detector 130 is ready to detect X-rays, the control unit 113 and the control unit 150 transmit an irradiation signal to the high voltage generator 121. Accordingly, the high voltage generator 121 generates a high voltage to apply the high voltage to the X-ray source 122, and the X-ray source 122 irradiates X-rays.

In transmitting the irradiation signal, the control unit 113 and the control unit 150 may transmit a sound output signal to the output unit 111 and the output unit 141 so that the output unit 111 and the output unit 141 output a specific sound. Further, the output unit 111 and the output unit 141 may output a sound representing other information related to photographing, in addition to the X-ray irradiation. In fig. 3A, the output unit 141 is illustrated as being included in the manipulation unit 140, but the embodiment is not limited thereto. The output unit 141 or a portion of the output unit 141 may be located at a different position from the manipulation unit 140. For example, the output unit 141 may be located on a wall surface of an examination room in which X-ray photographing of the subject is performed.

The control unit 113 and the control unit 150 control the positions, the photographing timing, and the photographing conditions of the X-ray irradiation unit 120 and the detector 130 according to the photographing conditions set by the user.

Specifically, the control unit 113 and the control unit 150 control the high voltage generator 121 and the detector 130 in accordance with commands input via the input unit 112 and the input unit 142 so as to control the irradiation timing of the X-rays, the intensity of the X-rays, and the irradiation area of the X-rays. In addition, the control unit 113 and the control unit 150 adjust the position of the detector 130 according to the photographing condition and control the operation timing of the detector 130.

In addition, the control unit 113 and the control unit 150 generate a medical image of the subject by using the image data transmitted from the detector 130. Specifically, the control unit 113 and the control unit 150 receive the image data from the detector 130, and then generate a medical image of the object by removing noise in the image data and adjusting the dynamic range and interleaving of the image data.

The output unit 111 and the output unit 141 may output the medical image generated by the control unit 113 and the control unit 150. The output unit 111 and the output unit 141 may output information necessary for a user to manipulate the X-ray apparatus 100, for example, a User Interface (UI), user information, or object information. Examples of the output unit 111 and the output unit 141 may include a printer, a Cathode Ray Tube (CRT) display, a Liquid Crystal Display (LCD), a Plasma Display Panel (PDP), an Organic Light Emitting Diode (OLED) display, a Field Emission Display (FED), a Light Emitting Diode (LED) display, a Vacuum Fluorescent Display (VFD), a Digital Light Processing (DLP) display, a main flight display (PFD), a three-dimensional (3D) display, a transparent display, and other various output devices known in the art.

The workstation 110 shown in fig. 3A may further comprise a communication unit (not shown) connectable to the server 162, the medical device 164 and the portable terminal 166 via the network 15.

The communication unit may be connected to the network 15 via wire or wirelessly to communicate with the external server 162, the external medical device 164, or the external portable terminal 166. The communication unit may transmit or receive data related to diagnosis of the subject via the network 15, and may transmit or receive a medical image captured by another medical device 164 (e.g., a CT device, an MRI device, or an X-ray device). In addition, the communication unit may receive a medical history or a treatment schedule of a subject (e.g., a patient) from the server 162 to diagnose a disease of the subject. In addition, the communication unit may perform data communication with a portable terminal 166 such as a mobile phone of a doctor or a patient, a Personal Digital Assistant (PDA), or a laptop computer, and a server 162 or a medical device 164 in a hospital.

The communication unit may include one or more elements that enable communication with an external device, for example, a short-range communication module, a wired communication module, and a wireless communication module.

The short-range communication module is a module for communicating with devices located within a predetermined distance. The short-range communication technology may be wireless Local Area Network (LAN), Wi-Fi, Bluetooth, ZigBee, Wi-Fi direct (WFD), Ultra Wideband (UWD), Infrared data Association (IrDA), Bluetooth Low Energy (BLE), Near Field Communication (NFC), etc.; however, embodiments of the present invention are not limited thereto.

The wired communication module is a module for performing communication by using an electric signal or an optical signal, and the wired communication technology may be a wired communication technology using a twisted pair cable, a coaxial cable, or an optical fiber cable, and a wired communication technology known in the art.

The wireless communication module may transmit/receive a wireless signal to/from at least one of a base station, an external device, and a server in a mobile communication network. Here, the wireless signal may be a voice call signal, a video call signal, or various types of data transmitted according to a text/multimedia message.

The X-ray device 100 shown in fig. 3A may include a plurality of Digital Signal Processors (DSPs), a subminiature calculator, and processing circuitry for special purposes (e.g., high speed analog/digital (a/D) conversion, high speed fourier transform, array processing, etc.).

In addition, the communication between the table 110 and the X-ray generator 100 may use a high-speed digital interface such as Low Voltage Differential Signaling (LVDS), an asynchronous serial communication such as a Universal Asynchronous Receiver Transmitter (UART), a synchronous serial communication, or a low-delay network protocol such as a Controller Area Network (CAN), or any other various communication methods known in the art.

Fig. 3B is a diagram illustrating the stationary type X-ray apparatus 200.

Fig. 3B is a perspective view showing the stationary type X-ray apparatus 200. The X-ray device 200 of fig. 3B may be an embodiment of the X-ray device 100 of fig. 1. Among elements included in the X-ray device 200 of fig. 3B, the same elements as those of fig. 1 are denoted by the same reference numerals as those of fig. 1, and a repetitive description will not be provided.

As shown in fig. 3B, the X-ray apparatus 200 includes: a manipulation unit 140 for providing a user with an interface for manipulating the X-ray device 200; an X-ray irradiation unit 120 that irradiates X-rays to a subject; a detector 130 that detects X-rays that have passed through the object; a first motor 211, a second motor 212, and a third motor 213 providing driving force to convey the X-ray irradiation unit 120, the guide rail 220, the movable gantry 230, and the column gantry 240, wherein the guide rail 220, the movable gantry 230, and the column gantry 240 are formed to convey the X-ray irradiation unit 120 by using the driving force of the

motors

211, 212, and 213.

The guide rail 220 includes a first guide rail 221 and a second guide rail 222 provided to form a predetermined angle with respect to each other. The first guide rail 221 and the second guide rail 222 may extend in directions crossing each other, respectively.

The first guide rail 221 is provided on a ceiling of an examination room in which the X-ray apparatus 200 is arranged.

The second guide rail 222 is located below the first guide rail 221 and is mounted on the first guide rail 221 so as to slide along the first guide rail 221. A roller (not shown) movable along the first guide rail 221 may be provided on the first guide rail 221. The second guide rail 222 is connected to rollers (not shown) to move along the first guide rail 221.

The first direction D1 is defined as a direction in which the first guide rail 221 extends, and the second direction D2 is defined as a direction in which the second guide rail 222 extends. Thus, the first direction D1 and the second direction D2 intersect each other and may be parallel to the ceiling of the examination room.

The movable frame 230 is disposed below the second guide rail 222 so as to move along the second guide rail 222. A roller (not shown) moving along the second guide rail 222 may be provided on the movable frame 230.

Accordingly, the movable frame 230 may move together with the second guide rail 222 in the first direction D1 and may move along the second guide rail 222 in the second direction D2.

The column frame 240 is fixed to the movable frame 230 and is located below the movable frame 230. The column frame 240 may include a plurality of

columns

241, 242, 243, 244, and 245.

The plurality of

columns

241, 242, 243, 244, and 245 are connected to each other to be foldable, and thus, the column frame 240 may have a length adjustable in the up-down direction of the examination room in a state of being fixed to the movable frame 230.

The third direction D3 is defined as a direction in which the length of the column housing 240 increases or decreases. Accordingly, the third direction D3 may intersect the first direction D1 and the second direction D2.

The detector 130 detects X-rays that pass through the object and may be coupled to the table susceptor 290 or the vertical susceptor 280.

The rotary joint 250 is disposed between the X-ray irradiation unit 120 and the column gantry 240. The rotary joint 250 allows the X-ray irradiation unit 120 to be coupled to the column housing 240 and supports a load applied to the X-ray irradiation unit 120.

The X-ray irradiation unit 120 connected to the rotary joint 250 is rotatable on a plane perpendicular to the third direction D3. Here, the rotation direction of the X-ray irradiation unit 120 may be defined as a fourth direction D4.

Further, the X-ray irradiation unit 120 may be configured to be rotatable on a plane perpendicular to the ceiling of the examination room. Accordingly, the X-ray irradiation unit 120 may be rotated in a fifth direction D5, wherein the fifth direction D5 is a rotational direction based on an axis parallel to the first direction D1 or the second direction D2 with respect to the rotary joint 250.

The first motor 211, the second motor 212, and the third motor 213 may be provided to move the X-ray irradiation unit 120 in the first direction D1, the second direction D2, and the third direction D3. The first motor 211, the second motor 212, and the third motor 213 may be electrically driven, and the first motor 211, the second motor 212, and the third motor 213 may include encoders, respectively.

The first motor 211, the second motor 212, and the third motor 213 may be disposed at various positions by considering design convenience. For example, the first motor 211 moving the second guide rail 222 in the first direction D1 may be disposed around the first guide rail 221, the second motor 212 moving the movable frame 230 in the second direction D2 may be disposed around the second guide rail 222, and the third motor 213 increasing or decreasing the length of the column frame 240 in the third direction D3 may be disposed in the movable frame 230. In another example, the first motor 211, the second motor 212, and the third motor 213 may be connected to a driving force transmission unit (not shown) to linearly move the X-ray irradiation unit 120 in the first direction D1, the second direction D2, and the third direction D3. The driving force transmission unit (not shown) may be a general belt and pulley, a chain and sprocket, or a shaft.

As another example, motors may be disposed between the rotary joint 250 and the column gantry 240 and between the rotary joint 250 and the X-ray irradiation unit 120 to rotate the X-ray irradiation unit 120 in the fourth direction D4 and the fifth direction D5.

The manipulation unit 140 may be disposed at one side of the X-ray irradiation unit 120.

Although fig. 3B illustrates the fixed type X-ray apparatus 200 connected to the ceiling of the examination room, the X-ray apparatus 200 of fig. 3B is only an example for easy understanding. That is, the X-ray apparatus according to the embodiment of the present invention may be an X-ray apparatus having various structures, for example, a C-arm X-ray apparatus and an angiographic X-ray apparatus, without departing from the spirit and scope of the present invention, which will be apparent to those skilled in the art.

Fig. 3C is a diagram showing the mobile X-ray apparatus 300.

Fig. 3C is a diagram illustrating a configuration of a mobile X-ray device 300 capable of performing an X-ray photographing operation regardless of a place where the photographing operation is performed, according to an embodiment of the present invention. The X-ray device 300 of fig. 3C may be an embodiment of the X-ray device 100 of fig. 1. Among elements included in the X-ray device 300 of fig. 3C, the same elements as those of fig. 1 are denoted by the same reference numerals as those of fig. 1, and a repetitive description is not provided.

The mobile X-ray device 300 shown in fig. 3C comprises: a transport unit 370 comprising wheels for transporting the X-ray device 300; a main unit 305 (including a manipulation unit 140 providing an interface for manipulating the X-ray device 300; a high voltage generator 121 generating a high voltage applied to the X-ray source 122; a control unit 150 controlling the overall operation of the X-ray device 300); an X-ray irradiation unit 120 (including an X-ray source 122 generating X-rays; a collimator 123 for guiding a path of the X-rays generated and emitted from the X-ray source 122 to adjust an irradiation area of the X-rays); and a detector 130 for detecting the X-ray irradiated from the X-ray irradiation unit 120 and passing through the object.

In fig. 3C, the detector 130 is shown as a bench-top 390, but it is apparent that: the detector 130 may be implemented as a vertical type.

In fig. 3C, the manipulation unit 140 is shown as being included in the main unit 305, but the embodiment is not limited thereto. For example, as shown in fig. 2, the manipulation unit 140 of the X-ray device 300 may be provided at one side of the X-ray irradiation unit 120.

The radiation detector according to an exemplary embodiment is a device for sensing radiation and senses incident radiation photons in a direct type. Thus, the radiation detector according to an exemplary embodiment may be applied to all electronic devices that sense radiation photons.

Specifically, the radiation detector according to an exemplary embodiment may correspond to the X-ray detector 178 described above with reference to fig. 1A and 1B and may be included in the tomographic imaging system 20 described above with reference to fig. 1A and 1B. Specifically, the radiation detector according to the exemplary embodiment may be a radiation detector used to generate a tomographic image. In particular, the radiation detector according to an exemplary embodiment may be a radiation detector used for generating a CT image. Specifically, the radiation detector according to an exemplary embodiment may sense radiation emitted from the X-ray generation unit 176 and having passed through the object, wherein the X-ray generation unit 176 is an X-ray source that is attached to the gantry of fig. 1A and 1B and rotates.

Furthermore, the radiation detector according to an exemplary embodiment may correspond to the detector 130 described above with reference to fig. 3A and 3B and may be included in the X-ray system 1000 or the

X-ray device

100, 200, and 300 described above with reference to fig. 3A to 3C. In particular, the radiation detector according to an exemplary embodiment may be a radiation detector that is used for generating an X-ray image. Specifically, the radiation detector according to an exemplary embodiment may sense radiation emitted from the X-ray generation unit 176 and having passed through the object, wherein the X-ray generation unit 176 is an X-ray source attached to the mobile device and adjusted in position. Here, the moving apparatus to which the X-ray source is attached may include at least one selected from the guide rail 220, the movable gantry 230, and the column gantry 240, which have been described above with reference to fig. 3B. Further, the mobile device may include the transport unit 370 described above with reference to fig. 3C.

A radiation detector according to an exemplary embodiment will now be described in detail with reference to fig. 4 to 9.

Fig. 4 is a diagram illustrating a radiation detector 400. The radiation detector 400 is a counting detector that converts incident radiation into electric charges in a direct method to detect radiation that has passed through the object 10. Specifically, the radiation detector is a photon counting detector that converts incident radiation photons into electrical signals and counts the number of converted electrical signals corresponding to the photons.

Referring to fig. 4, a radiation detector 400 according to an embodiment of the present invention includes a plurality of pixels 410 and 430 that detect radiation. The plurality of pixels 410 and 430 each include a plurality of sub-pixels 411 and 415. Hereinafter, one pixel includes m sub-pixels, and the sub-pixel 411 of one of the sub-pixels of the pixel 410 will be described as an example.

In detail, the radiation detector 400 may be a radiation detector for generating a CT image and may correspond to the X-ray detection unit 108 of fig. 1 and 2.

As shown, the plurality of pixels 410 and 430 may have a tetrahedral structure arranged in a lattice form and the pixels may have the same shape and/or size.

Subpixel 411 includes a radiation absorbing layer 412 and an optical processor 413.

Here, the counting process of incident photons is performed in units of the sub-pixels 411, and thus the sub-pixels 411 may be referred to as counting pixels. Hereinafter, a sub-pixel which is a local pixel included in the pixel 410 is referred to as a count pixel. Further, one pixel value of an image restored based on the number of photons counted in the at least one count pixel may be determined, and thus, a count pixel group including the at least one count pixel may be referred to as an image pixel. For example, when one pixel value of an image is obtained based on the number of photons counted in the plurality of count pixels 411 and 415 included in the pixel 410, the image pixel becomes the pixel 410. As another example, when one pixel value of an image is obtained based on the number of photons counted in four adjacent count pixels, the image pixel may become a count pixel group including four count pixels.

Specifically, the radiation detector 400 includes at least one count pixel 411 and includes a plurality of image pixels for restoring an image. The count pixel 411 includes: a radiation absorbing layer 412 that converts incident photons into electrical signals; an optical processor 417 that counts the number of photons based on the electrical signal transmitted from the radiation absorbing layer 412. Here, the number of image pixels included in the radiation detector 400 is smaller than the number of count pixels 411. Further, the size of each of the image pixels included in the radiation detector 400 is larger than the size of the count pixel.

Specifically, the count pixel 411 counts the number of photons smaller than the number of photons incident on the image pixel.

Specifically, the image pixel corresponds to one pixel value constituting the image, and the one pixel value in the image is calculated based on the total number of photons counted in the one image pixel. Specifically, the image pixel may include a plurality of count pixels, and one pixel value in the image is calculated based on a total number of photons counted in a count pixel group including the plurality of count pixels. When a plurality of count pixels included in the pixels 410 constitute one count pixel group, one pixel 410 may become one image pixel. In addition, when a plurality of count pixels included in the pixel 410 constitute a plurality of count pixel groups, the pixel 410 corresponds to one image pixel of one count pixel group, and thus may include a plurality of image pixels.

In addition, in the radiation detector 400 including the plurality of pixels 410 sensing radiation, the pixels 410 include a plurality of count pixels 411 and 415 sensing radiation for restoring an image. Here, the count pixel 411 includes: a radiation absorbing layer 412 that converts incident photons into electrical signals; and an optical processor 413 for counting the number of photons based on the electrical signal.

The absorption layer may be arranged on any other surface than the surface facing the X-ray source, such as a side or a back surface, or any other surface. In the depicted embodiment, the radiation absorbing layer 412 may consist of the entire thickness of the layers depicted in fig. 4, or may be arranged on the sides of the sub-pixel 411.

Radiation absorbing layer 412 may convert incident X-ray photons into electrical signals. The radiation absorbing layer 412 may transmit electrical signals to an optical processor 413.

Further, a radiation absorbing layer 412 may be formed on at least a portion of the surface facing the X-ray source. Specifically, the radiation absorbing layer 412 may be formed on at least a part of a front surface of the radiation detector 400, which is a surface facing the X-ray source, a side of the surface facing the X-ray source, or a back surface of the radiation detector 400, on which X-rays from the X-ray source are likely to be incident due to scattering. In fig. 4, a case where the radiation absorbing layer 412 is formed on the front surface of the radiation detector 400 facing the X-ray source to have a uniform thickness is shown as an example.

Specifically, the radiation absorbing layer 412 can convert radiation photons directly into electrical signals and can be formed from cadmium telluride (CdTe). CdTe is a semiconductor material. Potentially, optical processors and memories may be formed in the underlying semiconductor layers as indicated at 520 in fig. 5, although the exemplary material CdTe itself is less suitable for integrating any semiconductor components therein. The underlying layer or back 520 may also be formed of other semiconductor materials for the same purpose. Regardless of whether the underlying layer or back is also made of CdTe or of an alternative semiconductor material, a very compact configuration can be obtained with a monolithic semiconductor assembly.

The optical processor 413 counts the absorbed photons. Specifically, the optical processor 413 counts the number of photons based on the electrical signal generated by the radiation absorbing layer 412 according to a direct method of directly converting incident photons into electrical charges to detect the photons.

Specifically, the optical processor 413 compares the energy of the absorbed photon with a reference value and counts the number of photons having an energy equal to or greater than the reference value.

The optical processor 413 may include a count memory (not shown) that stores the number of photons counted. Specifically, the photo-processor 413 included in the count pixel 411 may count and store the number of photons smaller than the number of photons incident on the image pixel at a specific time.

The memory (not shown) stores values obtained by the optical processor 413 through counting. The memory (not shown) has a storage capacity with a value of about n/m when the pixel 410 absorbs about n photons during a certain predetermined period of time, where m represents the number of sub-pixels.

Referring to fig. 4, the sub-pixel 411 includes an optical processor 413 and a memory (not shown), and the sub-pixel 415 includes an optical processor 417 and a memory (not shown).

As shown, a photon processor 413 is included in each count pixel 411, and allows a photon counting operation to be performed individually for each count pixel 411.

The structure of each pixel and the structure of each count pixel will be described in more detail with reference to fig. 5 to 7.

Fig. 5 is a diagram for describing a plurality of pixels of fig. 4.

Referring to fig. 5, the radiation detector 400 includes a plurality of pixels arranged in a lattice form. In fig. 5, in the lower front view, a case where 256 pixels (16 × 16 pixels) are included in the radiation detector unit is shown as an example. In the upper part of fig. 5, a perspective view of a part of the radiation detector is depicted.

Referring to fig. 5, the radiation absorbing layer 412 and the radiation absorbing layer 416 may be disposed in the front portion 510 of the radiation detector 400 or may be disposed in the front portion 510. Optical processor 413, optical processor 417, and memory (not shown) may be disposed on rear 520 of radiation detector 400, on rear 520, behind rear 520, or to one side of rear 520. In this embodiment, the front portion 510 may constitute or include an absorbent layer. The relative expressions "front" and "rear" are used herein with respect to the direction from which X-rays are incident on the radiation detector 400. As indicated above, the rear portion may be formed of a semiconductor material, and the optical processor and memory of each sub-pixel may be implemented therein to provide a very compact configuration, resembling or forming a sandwich.

In detail, radiation passing through the subject is incident on the front side 540 of the radiation detector 400, wherein the radiation absorbing layers 412 and 416 arranged in the front portion 510 or in the front portion 510 absorb the incident radiation and transmit the absorbed radiation to the respective optical processors 413 and 417 connected to the radiation absorbing layers 412 and 416, respectively. Specifically, the radiation absorbing layers 412 and 416 convert photons (incident radiation) into electrical signals and transmit the converted electrical signals to the optical processors 413 and 417.

The front of one pixel 541 may have a size of 1mm ^ 2. Specifically, a length of one side edge of one pixel 541, which defines a perimeter of the one pixel 541 in the lower front view of fig. 5, may be about 0.9mm to about 1.1 mm.

Fig. 6 is a diagram for describing the sub-pixel of fig. 4.

Referring to the upper diagram in fig. 6, one pixel 600 may include a plurality of sub-pixels. For example, one pixel 600 in the lower right front view 642 in fig. 6 may include 24 pixels (4 × 6 pixels), and one pixel 600 in the lower left front view 641 in fig. 6 may include 36 pixels (6 × 6 pixels). Optionally, pixel 600 may include (not shown) 16 pixels (4 x 4 pixels), 25 pixels (5 x 5 pixels), 36 pixels (6 x 6 pixels), 35 pixels (7 x 5 pixels), and as described and shown, the assembly of sub-pixels may be rectangular, or alternatively may be square, honeycomb, triangular, diamond-shaped, or cross-shaped (e.g., including 5 pixels, 20 pixels, 45 pixels, etc.), any combination of the shapes and/or other shapes, or have any other suitable front view shape, with one factor considered being that the sub-pixels should be closely adjacent or neighboring to cover the entire surface of the pixel constructed from the sub-pixels to which the sub-pixels belong.

A radiation detector comprised in a CT system may absorb a certain number of photons when the object is imaged under certain imaging conditions. The number of photons absorbed and counted by a pixel having a unit area of about 1mm 2 can be determined according to the following spectral modeling.

In the photographing condition for the photon counting detector included in the advanced CT system or the high-specification CT system, the tube voltage may be set to about 120kVp, the tube current may be set to a minimum value of 200mA or more, and the aluminum equivalent thickness corresponding to the filtering condition (filter condition) may be set to about 5.6 mm.

Under the photographing condition, the number of photons absorbed and counted by one pixel 600 may be calculated according to X-ray spectral modeling based on a tungsten anode spectral model using interpolation polynomial (tamip).

In detail, the number of photons absorbed by one pixel per second may be about two hundred million to about five hundred million. Here, one pixel may have a unit area of 1mm ^ 2.

For example, spectral modeling can be designed as follows.

In spectral modeling, the average photon energy was about 60.605keV (kilo-electron-volts) and the first half-value layer was 6.886mm Al. When the measurement was performed at a distance of about 1m, the exposure was about 7,730mR/mAs, and when the measurement was performed at a distance of about 1m, the air kerma was about 67.799 uGy/mAs. Under the modeling conditions, when the measurement is performed at a distance of about 1m, the unit area of 1mm ^2 is passed and the incident fluence is 2,004,955[ photons/mm ^2/mAs ]

According to the spectral modeling, the number of photons generated when emitting about 1mA of X-rays is about 2,004,955 photons/mm ^2/mAs, i.e., about 2 million photons/mm ^ 2/mAs. Hereinafter, million (M) is used as a unit of million.

In the case of a detector with a dose of about 200mA, the number of photons incident on a unit area of about 1mm ^2 can be about 200 x 2,004,955 photons/mm ^2/s, i.e., about 400M photons/mm ^ 2/s. Furthermore, when the detector is operated at a dose of about 100mA (i.e., when the dose of radiation is reduced by about 50%), the number of photons incident on a unit area of about 1mm ^2 may be about 100 x 2,004,955 photons/mm ^2/s, i.e., about 200M photons/mm ^ 2/s.

Thus, a pixel 600 having a unit area of about 1mm ^2 can absorb and count 200M or more photons per second. Even a minimum of 200M photons/mm 2, for example, may be required to achieve the desired resolution, but accumulating count values of 200M photons per 1mm 2 pixel size can take longer than radiation detectors available in the prior art, which is particularly disadvantageous, but not just in CT scanners where the speed of continuous image registration is critical. In any case, the requirement for a radiation detector capable of registering (register)200M photons per 1mm 2 per second has eluded those skilled in the art to date, as will be explained in more detail below.

Although the minimum of 200M photons/mm 2 per second is indicated here as the current requirement, its value may vary over time, and it should be noted here that: the invention according to the present disclosure is able to meet higher and even higher minimum requirements than the above-mentioned 200M photons per 1mm 2 per second.

Further, one of the 25 count pixels in a pixel (when the value of M is 25) may absorb and count 8M (200M/25) or more photons per second. Subsequently, as described more fully below, the pixel as a whole is more capable of absorbing and counting an impinging dose of 200M photons than any embodiment that utilizes only the entire pixel, which requires a sufficient amount of 200M photons to be absorbed and counted.

Referring to the lower left portion in fig. 6, the pixel 600 shown in the front view 641 may include 36 count pixels (6 × 6 count pixels). That is, the front side 640 of the pixel 600 may be shown in the lower left front view 641 in fig. 6. As described above, when the pixel 600 absorbs and counts about 200M photons per second and includes 36 count pixels, one count pixel can absorb and count 5.56M photons per second (200/36M photons).

Referring to the lower right portion of fig. 6, the pixels 600 may include 24 count pixels (6 × 4 count pixels) as indicated by reference numeral 642. That is, the front surface 640 of the pixel 600 may correspond to reference numeral 642. As described above, when the pixel 600 absorbs and counts about 200M photons per second and includes 24 count pixels, one count pixel can absorb and count 8.336M photons per second (200/24M photons). In addition, pixel 600 may include 25 counting pixels (5 × 5 counting pixels). As described above, when the pixel 600 absorbs and counts about 200M photons per second and includes 25 counting pixels, one counting pixel can absorb and count 8M photons per second (200/25M photons).

As in the above-described embodiments, the number of photons counted at a certain time may be set based on detailed product specifications (e.g., an X-ray apparatus, a tomographic imaging apparatus, etc.) to which the radiation detector is applied, and the number and size of the count pixels included in one pixel may be adjusted based on the set number of photons. For example, the size of a count memory included in the optical processor may be adjusted based on the set number of photons. In addition, here should be noted: in the upper diagram of fig. 6, the pixel 600 is also shown as having a front 610 and a back 620, with a front surface 640 on the front 610 with respect to the direction of radiation incidence.

Fig. 7 is another diagram for describing the count pixel of fig. 4.

Referring to fig. 7A, in a radiation detector, a counting pixel 700 includes a radiation absorbing layer 710, an optical processor 720, and a memory 730. The radiation absorbing layer 710 is the same as the radiation absorbing layer 412 or 416 of fig. 4 or at least comparable to the radiation absorbing layer 412 or 416 of fig. 4.

The radiation absorbing layer 710 converts incident X-ray photons into electrical signals. The radiation absorbing layer 710 may be arranged in front of the counting pixel 700 and may comprise a capacitor formed from CdTe. The radiation absorbing layer 710 may absorb X-ray photons and may be charged with an electrical signal obtained by converting each X-ray photon. The radiation absorbing layer 710 may transmit the charged electrical signals to an optical processor 720. Here, the electric signal obtained by the radiation absorbing layer 710 through conversion may be a voltage signal.

The optical processor 720 may include a comparator 721 and a counter 723.

The comparator 721 compares the electric signal with a reference value to determine whether the electric signal exceeds the reference value. Specifically, when the electric signal is a voltage signal, the comparator 721 compares the electric signal corresponding to the photon with a predetermined reference voltage Sref. When the electric signal is greater than the reference voltage Sref as a result of the comparison, the comparator 721 outputs a signal accumulated and counted by the counter 723 to the counter 723.

Here, the reference voltage Sref is a value corresponding to the energy of photons and may be changed according to the X-ray source. For example, a value that enables determination as to whether or not the electric signal input from the comparator 721 is generated by converting photons may be set as the reference value Sref.

The counter 723 counts the number of photons according to the output signal of the comparator 721.

For example, when the comparator 721 is biased to the + Vh voltage and the-Vh voltage, the comparator 721 may output the + Vh voltage as a logic high level signal and the-Vh voltage as a logic low level signal. The comparator 721 may output a + Vh voltage corresponding to a logic high level when the level of the electrical signal corresponding to the photon is higher than that of the reference voltage Sref, and when the + Vh voltage is input to the counter 723, the counter 723 may increase the count value of the number of incident photons by +1 and thus count the number of photons. On the other hand, when the level of the electric signal corresponding to the photon is lower than the level of the reference voltage Sref, the comparator 721 may output a-Vh voltage, and when the-Vh voltage is input to the counter 723, the counter 723 may continue to accumulate a count value of the number of photons that have formed a logic low level (i.e., the comparator 721 has discriminated for the photons that the electric signal from the radiation absorbing layer 710 does not correspond to the impact of the photons thereon) without increasing the value of the number of photons for the electric signal from the radiation absorbing layer 710 to the optical sub-processor 720.

The memory 730 stores the number of photons counted by the counter 723. Specifically, when n photons are absorbed by one pixel at a specific time, the memory 730 needs to have a storage capacity of a value of n/m. For example, when a pixel absorbs and counts about 200M or more photons per second and includes 25 counted pixels, the memory 730 stores a number of bits corresponding to about 8M to store 8M or more photons per second (200/25M photons). The storage capacity of the memory 730 may be set based on the number of photons counted by one counting pixel at a specific time.

In addition, in fig. 7A, a case where the memory 730 is included in the optical processor 720 is shown as an example, but the memory 730 may be provided separately from the optical processor 720.

In fig. 7B, an alternative embodiment to the configuration of fig. 7A is shown. In the radiation detector according to fig. 7B, the counting pixel 724 comprises a radiation absorbing layer 710, a comparator 721 like the comparator in fig. 7A forming an embodiment of an optical processor 725, and an incremental counter 726. As indicated by the dashed lines in fig. 7B, the optical processor 725 may also include a comparator 721 or may be formed by only an incremental counter 726. The operating system may periodically read out the value accumulated in the increment counter 726 and reset the value accumulated therein to 0, after which a new period of counting impinging photons may begin. Thus, the need for separate counters and memories as in the embodiment of fig. 7A can be avoided in a simple refined way. In particular, a separate memory may be omitted.

Further, the incremental counter 726 may be referred to as a count memory 726. Specifically, the count memory 726 counts and stores the number of photons according to the output signal of the comparator 721. Further, the storage capacity of the count memory 726 may be set based on the number of photons counted by one count pixel at a specific time.

The necessary memory capacity of the memory in fig. 7A or the count value capacity of the incremental counter in fig. 7B may be varied or set or designed according to the amount of photons expected to impinge on the radiation absorbing layer 710 and the number of counted pixels m in each pixel during a measurement period (where the measurement period may be referred to as a "specific time" elsewhere in this disclosure). Any memory (separate or integrated) or incremental counter may preferably be read out in real time. Hereinafter, the expression of the memory capacity is also employed to indicate the maximum count value of the incremental counter of the embodiment of fig. 7B.

The following impression can be obtained from the presentation of fig. 7A and 7B: the comparator, counter and, if present, memory are arranged close to the counted pixels or at least form part of the detector unit, but in fact a counter or other form of optical processor, comparator and, if present, memory may preferably be arranged in the DAS 186 of fig. 1B or will preferably be arranged in the DAS 186 of fig. 1B.

Assuming that the radiation detector is designed with 25 counting pixels 700 per pixel and will be subject to no more than 500M photons per said specific time in use, then a count value of 500M/25 of 20M memory per counting pixel and for applications at lower levels of photon impingement and subsequent lower X-ray tube output are considered sufficient. Within the framework of the present disclosure and claims, it is possible to set a memory size corresponding to a predetermined given number M of counting pixels per pixel (e.g. 25 counting pixels) in view of the maximum incident radiation (e.g. 500M photons). In the case of 200 photons, the memory capacity anyway satisfies (500M/25 ═ 16M, wherein in that case a capacity of (200M/25 ═ 8M would be sufficient. Furthermore, the design trade-off between the required maximum memory capacity of the memory 730 or the maximum count value of the incremental counter 726, the maximum incident radiation n and the number of counted pixels m, the number of sub-pixels m may be changed in order to, for example, allow a higher or lower required maximum memory capacity. For example, in a maximum incident radiation of 500M photons, the required maximum memory capacity can be reduced to a maximum count value (500M/36) of 13.89M by adopting a design with 36 count pixels per pixel instead of 25 count pixels per pixel. However, as explained below, the counting pixel design of the pixels allows combinations of counting pixels to be created that are redefined in use as to the size, shape or other parameters of the pixels, thereby creating new boundaries for the pixels, in particular the number of impinging photons per time unit, depending on the use case. This allows even the following functions: the number of counted pixels in each pixel may be reduced if the required count value is obtained too quickly or increased if the necessary count value is not sufficiently obtained quickly. Similarly, as explained below, as a result of the higher count speed, the exposure time or specific time or period of time during which incident radiation is allowed to impinge on the radiation detector may be reduced. For example, if more incident radiation can be reliably detected, the shorter exposure time (also referred to as the "specific time" in other portions of the disclosure) may be reduced and thus the maximum incident radiation in the specific time will be reduced to allow the reservoir 730 to have a lower required capacity. In more detail, if the specific time is reduced from 1 second to, for example, half of 1 second, the maximum incident radiation is halved to 250M photons in half of one second. Then, the required maximum memory capacity of each memory for each count pixel in the case of a configuration having 25 count pixels may be reduced to 10M, which holds a count value (250M/25 ═ M).

For the latter aspect of reducing the specific time or period or exposure time, it should be noted that: each of the m count pixels in each pixel accumulates its own count value in memory 730 or in the incremental counter 726, or each count pixel. In previous configurations without separate counting of each counting pixel, when, for example, two or more photons actually impinge simultaneously on the front face 640 of the pixel 600 in FIG. 6 at a relatively large distance between the impingement within the surface of the pixel of size 1mm ^2, there is a risk that multiple photons actually impinging simultaneously induce an increase in a single count value of 1, while the count value in the memory of the entire pixel should be increased by the same number as the number of photons actually impinging simultaneously. After all, in that case the electrical signal from the photon absorbing layer will be temporarily higher than a reference value above which the comparator outputs a logic high level, but does not indicate that this is caused by multiple photons impinging at virtually the same time. Thus, a specific time or exposure time is required to obtain enough data to reconstruct an image with sufficient detail and/or at a desired resolution, which may often be significantly longer than 200M pixels actually impinging on previously configured pixels of size 1mm ^ 2.

As a novel feature to reduce the specific time required or exposure time, it is possible to enhance the combination of the radiation absorbing layer, the comparator and the counter to be able to distinguish photons which are actually impinging simultaneously more quickly. However, such an approach may lead to considerable and burdensome design requirements for previously configured radiation absorbing layers, comparators and counters for pixels having a size of 1mm ^2, the requirement for faster photon discrimination may be difficult to meet, although this is not excluded from the present disclosure.

However, in the above-described embodiment, the front surface 640 of the pixel is divided into a region of m counting pixels, where each counting pixel of the pixel counts photons impinging on the respective counting pixel and holds or stores a count value of the impinging photons in its respective memory or incremental counter. Since photons that actually impinge simultaneously may not impinge exactly on the same location of the surface 640 of the pixel 600, but are more likely to impinge on different count pixels of the pixel, distinct photons that actually impinge simultaneously on different count pixels of each pixel will be counted separately, contributing to the speed and total time required to obtain sufficient data to construct an image with sufficient detail and/or resolution. Thus, the required total specific time or exposure time may be reduced and/or the required minimum counting speed may be achieved, and as a result the required maximum capacity for the count values in the memory of the individual counting pixels may be reduced. Thus, a higher speed counting radiation detector may be provided.

Fig. 8 is another diagram for describing the count pixel of fig. 4.

In the radiation detector 400, the number of photons counted by at least one of the count pixels may correspond to one image pixel value of the restored image. In detail, at least one of the count pixels may be grouped, and the radiation detector 400 may generate one image pixel value of the CT image by using the number of photons counted by one group including a plurality of count pixels. Here, the grouped count pixels are referred to as a count pixel group. In particular, the total number of photons counted by one counting pixel group (e.g., 821) may correspond to one image pixel value of the restored CT image. In the radiation detector 400, the number of photons counted by one count pixel may also correspond to the one image pixel value of the restored image.

In fig. 8A, a case where one pixel 810 includes 24 count pixels (6 × 4 pixels) is shown as an example. In fig. 8B, a case where two adjoining

pixels

850 and 870 each include 36 count pixels (6 × 6 pixels) is shown as an example.

Referring to fig. 8A, the one pixel 810 includes a plurality of count pixels arranged adjacent to each other, and a plurality of count pixel groups 821 to 826 each include the plurality of count pixels. All count pixels belong to one of the groups and all groups are within the boundary of the one pixel 810. The total number of photons counted by the plurality of counting pixel groups 821 to 826 may correspond to one image pixel value of the restored image. In particular, the total number of photons counted by one counting pixel group (e.g., 821) may correspond to one image pixel value of the restored CT image.

In fig. 8A, a case where one count pixel group corresponding to one image pixel value includes four count pixels is shown as an example. In this case, when one pixel includes 24 counted pixels (6 × 4 pixels), the one pixel may be divided into 6 counted pixel groups (2 × 2 counted pixel groups), and the one pixel may generate 6 image pixel values from the restored image. Specifically, referring to fig. 8A, one pixel 810 includes 6 count pixel groups 821 to 826. Here, the count pixel groups 821 to 826 may construct an image pixel that generates one pixel value, and thus, the one pixel 810 includes 6 image pixels. Accordingly, the number of counting pixel groups included in the radiation detector 400 may be equal to or more than the number of pixels included in the radiation detector 400. Further, the size of the counted group of pixels (e.g., 821) can be equal to or less than the size of the pixels 810.

As an alternative example, 24 count pixels included in one pixel may be divided into four count pixel groups of 3 × 2 count pixels, and the one pixel may generate a single combined count value as an image pixel value or generate 4 different image pixel values for the restored image (one image pixel value for each of the 4 count pixel groups).

Referring to fig. 8B, two

adjacent pixels

850 and 870 are shown.

In the radiation detector 400, a plurality of count pixels included in the plurality of pixels may be divided into a plurality of groups, and the number of photons counted by one of the divided groups may correspond to one image pixel value of the restored image.

Referring to fig. 8B, 72 count pixels included in two

adjacent pixels

850 and 870 may be divided into 6 groups 881 to 886. In detail, one image pixel value of the restored image may be determined according to the total number of photons counted by 12 count pixels included in one group (e.g., 881). Count pixel group 881- & 886 crosses the boundary between

pixels

850 and 870. Thus, a pixel may be redefined as a group of counted pixels that is smaller than the total number of counted pixels in the original pixel 810, also allowing for a faster acquisition of a desired count value of the number of photons incident, in order to quickly arrive at a measurement that allows image restoration at a desired level of detail or resolution. The reason for this is that the area of these redefined pixels, which correspond to the counted pixel group and are smaller than the original pixels 810, is smaller than the area of the original pixels, and then a smaller number of photons strike in these redefined pixels in the same amount of time. Since operating the count pixels individually allows for an increase in resolution of the spatial distribution of impinging photons, and the count pixel groups forming the redefined pixels are smaller than the size of the original pixels 810, as described further below, the level of detail or resolution of the result may even be increased.

In addition, referring to fig. 8B, 6 count pixel groups 881 to 886 are included in two

pixels

850 and 870. That is, the size of the pixels (e.g., 850) included in one radiation detector 400 may be equal to or larger than the size of the counting pixel group (e.g., 881).

Fig. 9 is a diagram illustrating a computed tomography apparatus 900 according to an embodiment of the present invention.

The computed tomography apparatus 900 according to an embodiment of the present invention includes a radiation detector 910 and an image processor 950. The radiation detector 910 has the same technical spirit and configuration as those of the radiation detector according to the embodiment of the present invention described above with reference to fig. 2 and 4 to 8, and thus, the same description provided with respect to fig. 2 and 4 to 8 is not repeated.

Further, the image processor 950 may be an element corresponding to the image processor 196 described above with reference to fig. 1B. Alternatively, the image processor 950 may be an element corresponding to the external medical device 136 connected to the CT system 20 through the wired/wireless network 15 as in fig. 2.

Referring to fig. 9, a radiation detector 910 includes a plurality of pixels that detect radiation. Here, the plurality of pixels each include at least one of the count pixels. For example, one pixel may include m count pixels.

The plurality of count pixels each include: a radiation absorbing layer 412 that converts incident X-ray photons into electrical signals; an optical processor 413 that counts the number of photons converted into a plurality of electrical signals; a memory (not shown) storing the number of absorbed photons and having a storage capacity of a value of n/m when the corresponding pixel absorbs n photons.

The image processor 950 reconstructs a CT image based on the number of photons detected by the radiation detector 910. For example, the image processor 950 may generate a CT image, an OCT image, a PET-CT image, or an X-ray image based on the number of photons sensed by the radiation detector 910. Hereinafter, the case where the image processor 950 restores the CT image has been described above as an example.

In detail, the image processor 950 may generate one image pixel value of the CT image by using the number of photons counted by the counting pixel group. Here, the count pixel group includes at least one count pixel included in at least one pixel.

In addition, the image processor 950 may generate one image pixel value of the CT image by using the number of photons counted by one count pixel.

Further, the count pixel group corresponding to one image pixel value in the restored CT image may include a plurality of count pixels arranged adjacent to each other included in one pixel.

Further, the count pixel group corresponding to the one image pixel value in the restored CT image may include a plurality of count pixels arranged adjacent to each other included in the plurality of pixels.

For example, when each pixel includes 24 count pixels as shown in fig. 8A, the image processor 950 may divide the 24 count pixels included in one pixel into 6 groups and generate one image pixel value from the restored CT image by using the number of photons counted by one of the 6 count pixel groups. That is, in this case, photons detected by one pixel are used to generate 6 image pixel values from the restored CT image.

As another example, when each pixel includes 36 count pixels as shown in fig. 8B, the image processor 950 may divide 72 count pixels included in two pixels into 6 groups and generate one image pixel value from the restored CT image by using the number of photons counted by one of the 6 count pixel groups. That is, in this case, photons detected by the counting pixel of the two pixels are used to generate 6 image pixel values from the restored CT image.

The image processor 950 may adjust the number of count pixels used to generate one image pixel value from the restored CT image according to the resolution of the restored CT image. For example, when it is desired to generate a super high resolution CT image, the image processor 950 may generate one image pixel value from the restored CT image by using the number of photons counted by one count pixel. This may or may not have an effect on parameters during a particular time, such as the tube voltage or the length of that time.

The radiation detector 910 may detect incident radiation for a particular time (exposure time) to sample the detected radiation incident during the particular time. For example, the number of incident photons per second in a typical diagnostic radiation detector may be about 500M corresponding to an area of 1mm x 1 mm. Accordingly, the related art radiation detector performing a photon counting operation for each pixel measures the energy of incident photons for a specific sampling time and counts the number of photons having a value equal to or greater than a specific value. When 500M photons per second are incident on an area of 1mm x 1mm, the prior art radiation detector samples one photon in 2 nanoseconds (1/500M seconds). The prior art radiation detector reliably samples one photon per nanosecond according to the nyquist sampling theorem, where one nanosecond corresponds to half of 2 nanoseconds.

However, it is difficult to perform an operation of measuring and comparing the energies of photons to count the number of photons within 1 nanosecond. Even when two photons impinge on completely different locations of the surface 640 of the pixel 600 in fig. 6, there is a risk that the count value may only be increased by 1, even though two photons have impinged. Further, even if the circuit that measures and compares the energies of photons to count the number of photons performs the above-described measurement and comparison operations for the sampling time, it is difficult to adjust the response of the radiation absorbing layer that absorbs radiation to the sampling time. In addition, when several photons are simultaneously incident on the comparator, an operation of comparing the energies of the photons may not be normally performed. Further, although the comparison and counting operation for the energy of one photon is being performed, if another photon is incident, the counting operation may not be normally performed, and the count value for the latter impinging photon may be lost.

Further, the counting detector in the related art counts photons by pixel and includes a memory that stores the number of counted photons by frame or a memory that stores the number of photons counted by each group composed of a plurality of pixels.

As described above, in the radiation detector and the apparatus according to the embodiments of the present invention, the plurality of pixels included in the radiation detector each include m count pixels, wherein each count pixel includes a photo-processor that counts photons and a memory that stores the number of counted photons. In detail, each of the plurality of counting pixels separately performs an operation of counting photons and an operation of storing the number of counted photons. When the corresponding pixel absorbs n photons and counts them, the memory of the corresponding counting pixel has a storage capacity with a value of n/m.

Therefore, in the radiation detector according to the embodiment of the present invention, since the photon counting operation is individually performed for each counting pixel, the number of photons to be processed for each counting pixel is reduced by n/m compared to the number of photons to be processed by the pixel in the related art. Thus, a sampling time of 1/(n/m) seconds can be used for each photon and a sampling time of 1/(n/m) seconds is guaranteed for each photon, again reducing the requirements with respect to the radiation absorbing layer, the optical processor, the comparator and the memory (if provided separately). That is, the radiation detector according to the embodiment of the present invention guarantees a sampling time of 1/(n/m) seconds corresponding to m times of the sampling time of the radiation detector in the related art, as compared with the radiation detector in the related art in which the sampling time for each photon is 1/n seconds. Therefore, the degree of accuracy in counting photons is enhanced, and the radiation absorbing layer can sufficiently count the number of absorbed photons. Furthermore, since the number of photons processed by the comparator and the counter is reduced by n/m, the radiation detector according to the embodiment of the present invention solves the problem that the radiation detector in the related art cannot normally count the number of photons when the photons are incident at substantially the same time (i.e., hit within a time period in which the absorption layer, the comparator, and the counter may not distinguish the separately hitting photons according to the above nyquist sampling theorem).

Based on the above disclosure of embodiments, it is clear to a person skilled in the art that: the pixels are referred to as the minimum building blocks for providing the number of impinging photons and producing a reconstructed image therefrom. By dividing the contribution over M counting pixels to the number of counts of impinging photons, a more accurate and faster accumulation of the required number of photons can be achieved for the pixel, for example to achieve a desired high count value in a specific or available time period, a count value of for example 200M per second being required to date elusive to a person skilled in the art. Furthermore, embodiments of the present disclosure allow for the recombination of a selected number of counted pixels to redefine the pixels, thus taking into account the changing circumstances for e.g. photon injection and/or even reducing the number of counted pixels per pixel to increase the acquisition time of the desired or required count value.

Further, the memory of each count pixel is designed to have a storage capacity of a value of n/m, and therefore, the size of the memory included in each count pixel is minimized. Thus, according to embodiments of the present invention, the radiation detector is implemented by including one memory in each counting pixel.

Further, the radiation detector and the apparatus (such as a computed tomography apparatus) according to the embodiment of the present invention generate one image pixel value of the restored image by using the total number of photons counted by at least one counting pixel, thereby achieving the quality of the image according to the image resolution desired by the user.

The embodiments of the present invention can be written as computer programs and can be implemented in general-use digital computers that execute the programs using a computer readable recording medium.

Examples of the computer readable recording medium include magnetic storage media (e.g., ROM, floppy disks, hard disks, etc.), optical recording media (e.g., CD-ROMs, or DVDs), and the like.

It should be understood that: the embodiments described herein are to be considered in all respects only as illustrative and not restrictive. Descriptions of features or aspects within each embodiment should generally be considered as available for other similar features or aspects in other embodiments.

Although one or more embodiments of the present invention have been described with reference to the accompanying drawings, those of ordinary skill in the art will understand that: various changes in form and details may be made therein without departing from the scope of the invention as defined by the appended claims.

Claims

1. A radiation detector for producing an image, the detector sensing radiation comprising a plurality of pixels, each pixel of the plurality of pixels comprising a plurality of counting pixels,

wherein each of the plurality of count pixels includes:

a radiation absorbing layer that converts incident photons incident on respective ones of the plurality of counting pixels into electrical signals; and

an optical processor for counting the number of photons based on the electrical signal transmitted from the radiation absorbing layer, and

wherein the content of the first and second substances,

when the count pixel group includes at least one count pixel:

the optical processor is configured to perform a photon counting operation and a photon number storing operation separately for each of the plurality of counting pixels,

the number of photons counted by the at least one counting pixel comprised in the counting pixel group is used for calculating a pixel value in the image, and

the number of counted pixel groups included in the detector is smaller than the number of the plurality of counted pixels included in the detector.

2. The radiation detector of claim 1, wherein the optical processor counts the number of photons based on the electrical signal according to a direct method of directly converting incident photons to electrical charge to detect photons.

3. The radiation detector according to claim 1, wherein when the plurality of count pixels included in each of the plurality of pixels constitute one count pixel group, each of the plurality of pixels is a count pixel group for calculating the one pixel value included in the image.

4. The radiation detector of claim 1, wherein the optical processor comprises: and a count memory for counting and storing the number of photons smaller than the number of photons incident on the corresponding count pixel group within a specific time.

5. The radiation detector of claim 1, wherein the optical processor comprises:

a comparator to compare the electrical signal to a reference value to determine whether the electrical signal exceeds the reference value; and

and a count memory for counting and storing the number of photons exceeding the reference value based on the comparison result of the comparator.

6. The radiation detector as set forth in claim 1, wherein each of the plurality of counting pixels further includes: and a count memory for counting and storing the number of photons smaller than the number of photons incident on the corresponding count pixel group within a specific time.

7. The radiation detector of claim 1, wherein the plurality of counting pixels included in each of the plurality of pixels is divided into at least one counting pixel group, and the number of photons counted by the at least one counting pixel group is used to calculate one image pixel value in the image.

8. The radiation detector of claim 7, wherein the number of counting pixel groups is equal to or greater than the number of pixels.

9. The radiation detector of claim 7, wherein a size of the at least one counting pixel group is equal to or smaller than a size of the pixel.

10. The radiation detector of claim 1, a plurality of count pixels included in a plurality of adjacent pixels are divided into a plurality of count pixel groups, and a number of photons counted by each of the plurality of count pixel groups is used to calculate an image pixel value in the image.

11. The radiation detector of claim 1, wherein the radiation detector is a radiation detector for generating a tomographic image.

12. The radiation detector of claim 1, wherein the radiation detector is a radiation detector for producing an X-ray image.

13. A tomographic imaging apparatus comprising:

a radiation detector for generating a tomographic image, wherein the detector sensing radiation includes a plurality of pixels, each of the plurality of pixels including a plurality of count pixels; and

an image processor to reconstruct the tomographic image based on the number of photons sensed by the radiation detector,

wherein each of the plurality of count pixels includes:

wherein the content of the first and second substances,

when the count pixel group includes at least one count pixel:

the number of photons counted by the at least one count pixel included in the count pixel group is used to calculate one pixel value in the tomographic image, and