Patent Description:
Medical diagnostic imaging systems generate images of an object, such as a patient, for example, through exposure to an energy source, such as X-rays passing through a patient, for example. The generated images may be used for many purposes. Often, when a practitioner takes X-rays of a patient, it is desirable to take several X-rays of one or more portions of the patient's body from a number of different positions and angles, and preferably without needing to frequently reposition the patient. To meet this need, C-arm X-ray system has been developed. The term C-arm generally refers to an X-ray imaging device having a rigid and/or articulating structural member having an X-ray source and an image detector assembly that are each located at an opposing end of the structural member so that the X-ray source and the image detector face each other. The structural member is typically "C" shaped and so is referred to as a C-arm. In this manner, X-rays emitted from the X-ray source can impinge on the image detector and provide an X-ray image of the object or objects that are placed between the X-ray source and the image detector.

The C-arm X-ray system includes various electrical and optical components which have gain and offset values suitable for only a certain range of operating conditions. The gain drift of the electrical and optical components introduces undesirable artifacts in the final image produced by the C-arm X-ray system. Therefore, to remove these artifacts, it is necessary to calibrate the C-arm X-ray system regularly. Further, even without drift, the calibration is needed to establish the relationship between the X-ray technique being used and the image detector output. Moreover, the calibration helps in compensating for the non-ideal response of the X-ray tube and the X-ray detector.

<CIT> in the abstract states: "A system (<NUM>) includes a stationary gantry (<NUM>) and a rotating gantry (<NUM>), wherein the rotating gantry is rotatably supported by the stationary gantry. The rotating gantry (<NUM>) includes a primary source (<NUM>) that emits primary radiation and a detector array (<NUM>) having at least one row of detector elements (<NUM>) extending along a longitudinal axis. The primary source and the detector array are located opposite each other, across an examination region, and the primary radiation traverses a path (<NUM>) between the primary source and the detector array and through an examination region (<NUM>) and illuminates the at least one row of detector elements of the detector array, which detects the primary radiation. The system further includes a supplemental source (<NUM>), wherein the supplemental source is affixed to a nonrotating portion of the system and emits radiation that traverses a sub-portion of the path and illuminates the at least one row of detector elements of the detector array, which detects the secondary radiation.

<CIT> in the abstract states: "A computed tomography system has an x-ray source, x-ray detector, a gantry, and a controller configured to automatically initiate an air-calibration using the source and detector with air in an air space and to determine gain values for channels of the detector from the automatically initiated air-calibration. A computed tomography system having a patient bore, a controller, an x-ray source, and a detector is calibrated by receiving a setting for a medical scan to be performed by the computed tomography system for a particular patient, the setting being one of a plurality of optional values, scanning air and not the patient in the patient bore with the source and the detector using the setting, determining a gain value based on the scanning with the setting and not the other optional values, and scanning the patient with the source and the detector using the setting and the gain value.

<CIT> in the abstract states: "In order to provide a data processing device and the like, capable of performing highly accurate reference correction even in a case where an object protrudes in reference channels in most of the measurement views, an image processing device (data processing device) of an X-ray CT apparatus calculates a unit air calibration reference value which is a value per unit tube current of an air calibration reference value which is reference data measured during air calibration, calculates a reference value (estimated reference value) corresponding to an X-ray condition in the main scanning on the basis of an output tube current value in the main scanning and a unit air calibration reference value, and corrects normalized reference data obtained by normalizing a measured reference value in the main scanning with the estimated reference value, to be included in an allowable error range, so as to remove the influence of protrusion.

In accordance with an example of the present technique, a system for imaging an object is provided. The system includes an X-ray source operative to transmit X-rays through the object and a detector operative to receive the X-ray energy of the X-rays after having passed through the object and to generate corresponding object X-ray intensity. The system further includes a controller operative to measure a detector entrance dose with no object being placed on the X-ray beam path and to determine a relationship between an X-ray tube electrical parameter and the detector entrance dose. The controller is further operative to determine a relationship between the X-ray tube electrical parameter, the detector entrance dose and a detector average pixel intensity and to obtain a normalized air map as a function of the X-ray tube electrical parameter based on calibration image data. The controller is also operative to generate an air map based on the normalized air map, the detector entrance dose and the detector average pixel intensity and to reconstruct an image of the object based on the air map and the measured object X-ray intensity.

In accordance with another example of the present technique, a method for imaging an object is provided. The method includes transmitting X-rays from an X-ray source to the object and acquiring measurement data related to the object. The method also includes measuring a detector entrance dose with no object being placed on the X-ray beam path and determining a relationship between an X-ray tube electrical parameter and the detector entrance dose. The method further includes determining a relationship between the X-ray tube electrical parameter, detector entrance dose and a detector average pixel intensity and obtaining a normalized air map as a function of the X-ray tube electrical parameter based on calibration image data. Finally, the method includes generating an air map based on the normalized air map, the detector entrance dose and the detector average pixel intensity and reconstructing an image of the object based on the air map and the measurement data related to the object.

When introducing elements of various embodiments of the present embodiments, the articles "a," "an," "the," and "said" are intended to mean that there are one or more of the elements. Furthermore, any numerical examples in the following discussion are intended to be non-limiting, and thus additional numerical values, ranges, and percentages are within the scope of the disclosed embodiments. Furthermore, the terms "circuit" and "circuitry" and "controller" may include either a single component or a plurality of components, which are either active and/or passive and are connected or otherwise coupled together to provide the described function.

<FIG> illustrates an exemplary radiologic imaging system <NUM>, for example, for use in interventional medical procedures. In one embodiment, the system <NUM> may include a C-arm radiography system <NUM> configured to acquire projection data from one or more view angles around a subject, such as a patient <NUM> positioned on an examination table <NUM> for further analysis and/or display. To that end, the C-arm radiography system <NUM> may include a gantry <NUM> having a mobile support such as a movable C-arm <NUM> including at least one radiation source <NUM> such as an X-ray tube and a detector <NUM> positioned at opposite ends of the C-arm <NUM>. In exemplary embodiments, the radiography system <NUM> can be an X-ray system, a positron emission tomography (PET) system, a computerized tomosynthesis (CT) system, an angiographic or fluoroscopic system, and the like or combination thereof, operable to generate static images acquired by static imaging detectors (e.g., CT systems, MRI systems, etc.) prior to a medical procedure, or real-time images acquired with real-time imaging detectors (e.g., angioplastic systems, laparoscopic systems, endoscopic systems, etc.) during the medical procedure, or combinations thereof. Thus, the types of acquired images can be diagnostic or interventional.

In certain embodiments, the radiation source <NUM> may include multiple emission devices, such as one or more independently addressable solid-state emitters arranged in one or multi-dimensional field emitter arrays, configured to emit the X-ray beams <NUM> towards the detector <NUM>. Further, the detector <NUM> may include a plurality of detector elements that may be similar or different in size and/or energy sensitivity for imaging a region of interest (ROI) of the patient <NUM> at a desired resolution. In one embodiment, a dosimeter <NUM> is provided near the detector <NUM> to measure the X-ray dose per frame at the entrance of the detector <NUM>.

In certain embodiments, the C-arm <NUM> may be configured to move along a desired scanning path for orienting the X-ray source <NUM> and the detector <NUM> at different positions and angles around the patient <NUM> for acquiring information for 3D imaging of dynamic processes. Accordingly, in one embodiment, the C-arm <NUM> may be configured to rotate about a first axis of rotation. Additionally, the C-arm <NUM> may also be configured to rotate about a second axis in an angular movement with a range of about plus or minus <NUM> degrees relative to the reference position. In certain embodiments, the C-arm <NUM> may also be configured to move forwards and/or backwards along the first axis and/or the second axis.

Accordingly, in one embodiment, the C-arm system <NUM> may include control circuitry <NUM> configured to control the movement of the C-arm <NUM> along the different axes based on user inputs and/or protocol-based instructions. To that end, in certain embodiments, the C-arm system <NUM> may include circuitry such as tableside controls <NUM> that may be configured to provide signals to the control circuitry <NUM> for adaptive and/or interactive control of imaging and/or processing parameters using various input mechanisms. The imaging and/or processing parameters, for example, may include display characteristics, X-ray technique and frame rate, scanning trajectory, and gantry motion and/or position.

In certain embodiments, the detector <NUM> may include a plurality of detector elements <NUM>, for example, arranged as a 2D detector array for sensing the projected X-ray beams <NUM> that pass through the patient <NUM>. In one embodiment, the detector elements <NUM> produce an electrical signal representative of the intensity of the impinging X-ray beams <NUM>, which in turn, can be used to estimate the attenuation of the X-ray beams <NUM> as they pass through the patient <NUM>. In another embodiment, the detector elements <NUM> determine a count of incident photons in the X-ray beams <NUM> and/or determine corresponding energy.

Particularly, in one embodiment, the detector elements <NUM> may acquire electrical signals corresponding to the generated X-ray beams <NUM> at a variety of angular positions around the patient <NUM> for collecting a plurality of radiographic projection views for construction of X-ray images, such as to form fluoro image(s). To that end, control circuitry <NUM> for the system <NUM> may include a control mechanism configured to control position, orientation and/or rotation of the gantry <NUM>, the C-arm <NUM> and/or the components mounted thereon in certain specific acquisition trajectories.

In certain embodiments, the X-ray source <NUM> and the detector <NUM> for interventional imaging may be controlled using an X-ray controller <NUM> in the control mechanism <NUM>, where the X-ray controller <NUM> is configured to provide power and timing signals to the radiation source <NUM> for controlling X-ray exposure during imaging. Further, the control mechanism <NUM> may also include a gantry motor controller <NUM> that may be configured to control the rotational speed, tilt, view angle, and/or position of the gantry <NUM>. In certain embodiments, the control mechanism <NUM> also includes a C-arm controller <NUM>, which in concert with the gantry motor controller <NUM>, may be configured to move the C-arm <NUM> for real-time imaging of dynamic processes.

In one embodiment, the control mechanism <NUM> may include a data acquisition system (DAS) <NUM> for acquiring the projection data from the detector elements <NUM> and processing the data for image reconstruction by 2D image processor <NUM>, for reconstructing high-fidelity 2D images in real-time for use during the interventional procedure, and/or 3D image processor/reconstructor <NUM>, for generating 3D cross-sectional images (or 3D volumes), and subsequent illustration of the images on display <NUM>. Moreover, in certain embodiments, the data obtained by the DAS <NUM> may be input to a computing device <NUM>. Alternatively, in certain embodiments, the computing device <NUM> may store the projection data in a storage device <NUM>, such as a hard disk drive, a floppy disk drive, a compact disk-read/write (CD-R/W) drive, a Digital Versatile Disc (DVD) drive, a flash drive, or a solid-state storage device for further evaluation.

In one embodiment, the system <NUM> may include an operator console <NUM> that may be configured to allow selection and display of scanning modes, FOV, prior exam data, and/or intervention path. The operator console <NUM> may also allow on-the-fly access to 2D and 3D scan parameters and selection of an ROI for subsequent imaging, for example, based on operator and/or system commands.

Further, in certain embodiments, the system <NUM> may be coupled to multiple displays, printers, workstations, a picture archiving and communications system (PACS) <NUM> and/or similar devices located either locally or remotely, for example, within an institution or hospital, or in an entirely different location via communication links in one or more configurable wired and/or wireless networks such as a hospital network and virtual private networks.

In operation, during a 3D scan of the object (e.g., patient), the X-ray detector measures the image data to generate an actual object X-ray intensity, It after X-rays pass through the object. The 3D image processor/reconstructor <NUM> utilizes this actual object X-ray intensity It to generate the 3D image of the object. This actual object intensity It is related to the unattenuated X-ray intensity, I<NUM>, along the path from the X-ray source to the X-ray detector pixel by the Beer-Lambert law: <MAT> where t is the thickness of the object and µ is the attenuation coefficient of the object.

In accordance with an embodiment of the present technique, an air calibration is performed on the C-arm radiography system <NUM> by measuring the X-ray intensity with no object in the path of the X-ray beam <NUM> i.e., unattenuated X-ray intensity I<NUM>. The air calibration compensates for the X-ray field non-uniformity, X-ray detector pixel gain including analog to digital (A/D) converter gain non-uniformity, as well as the tube-detector alignment variation from view to view. The outcome of the calibration is a series of two-dimensional maps called as an air map which represents the unattenuated X-ray intensity I<NUM>. In general, the air calibration is designed to obtain the unattenuated intensity, I<NUM>, to normalize the scan data to the unobstructed beam intensity, which is used to determine the amount of attenuation caused by an object in the beam path. Based on the determined unattenuated intensity, I<NUM>, and measured actual object X-ray intensity It , the X-ray attenuation caused by an object in the beam may be calculated. The X-ray attenuation of the object is further used to generate or reconstruct the 3D image of the object.

<FIG> illustrates a flow chart <NUM> depicting an exemplary method for imaging an object in accordance with an embodiment of the present technique. In one embodiment, the method of flow chart <NUM> may be implemented in computing device <NUM> of <FIG>. Embodiments of the exemplary method may include computer executable instructions on a computing system or a processor. Generally, computer executable instructions may include routines, programs, objects, components, data structures, procedures, modules, functions, and the like that perform particular functions or implement particular abstract data types. Embodiments of the exemplary method including computer executable instructions may also be practiced in a distributed computing environment where optimization functions are performed by remote processing devices that are linked through a wired and/or wireless communication network. In the distributed computing environment, the computer executable instructions may be located in both local and remote computer storage media, including memory storage devices.

Embodiments of the present method describe techniques for enhanced imaging of high-quality 3D cross-sectional images using a C-arm system <NUM>. To that end, at step <NUM>, a detector entrance dose is measured with no object being placed on the path of the X-ray beam <NUM>. In one embodiment, the detector entrance dose is measured by dosimeter <NUM> placed at the entrance of the detector <NUM>. In one embodiment, the dosimeter measures the detector entrance dose in a unit of microgray/frame (µGy/Frame).

At step <NUM>, a relationship between an X-ray tube electrical parameter and the detector entrance dose is determined based on the measured dose data. The X-ray tube electrical parameter includes the electrical parameter applied to the X-ray tube such as a tube voltage, a tube current or combinations thereof. As will be appreciated by those skilled in the art, the unit for tube voltage is Kilovoltage peak (kVp) and the unit for tube current is milliampere (mA). In general, the tube voltage controls energy and quality of the X-ray beam produced by the X-ray tube whereas the tube current controls the quantity of the X-ray beam.

<FIG> shows a graphical plot <NUM> of an exemplary relationship between an X-ray tube electrical parameter and a detector entrance dose. A horizontal axis <NUM> of the plot <NUM> represents a tube voltage in kVp with a given tube current in mA whereas a vertical axis <NUM> represents detector entrance dose in µGy/Frame. In general, as the tube voltage of the X-ray tube is increased the detector entrance dose at the entrance of the detector entrance is measured by the dosimeter at the surface of the detector.

It should be noted that in one embodiment, the detector entrance dose may be measured while both the tube voltage and the tube current are being changed. In which case, the detector entrance dose becomes the function of both tube voltage and tube current. The tube voltage and the tube current may be changed as a predefined pair in accordance with a 3D automatic brightness system (ABS) table. The 3D ABS is used to keep the brightness of the displayed image at a constant level during X-ray examinations. In one embodiment, the 3D ABS also adjusts a digital gain to adjust the brightness of the image. The kV and mA may be adjusted as a pair depending on the patient and the part of the anatomy of the patient being examined. Note that the tube current mA needs to be adjusted such that the pixel values are around the middle of the dynamic range of the detector to avoid the non-linearity near the saturation. In such a case, the horizontal axis <NUM> can represent a tube voltage and tube current (kVp/mA) pair instead of the tube voltage (kVp) which is fixed for the given tube current (mA).

In one embodiment, a linear interpolation of the measurement points of the plot <NUM> in <FIG> is performed to find out the relationship between the X-ray tube electrical parameter and the detector entrance dose. For example, let the kVp be the actual kVp in the projection and let kVp(n) and kVp(n + <NUM>) be the two consecutive kVp values obtained during the air calibration i.e., n is an index number. Also assume that D(n) and D(n + <NUM>), respectively, are the two detector entrance dose values in µGy/frame corresponding to kVp(n) and kVp(n + <NUM>) obtained from the air calibration. The detector entrance dose D in µGy/frame corresponding to kVp in the projection is then given by the following linear interpolation: <MAT>.

Turning back to <FIG>, at step <NUM>, a relationship between the X-ray tube electrical parameter, the detector entrance dose and a detector average pixel intensity as measured by the detector <NUM> when no object is placed on the X-ray beam path is determined. The average pixel intensity is calculated by averaging pixel values of all the detector elements. In one embodiment, the average pixel value is determined by a central region of interest (e.g., <NUM> x <NUM> pixels, specifically, out of a <NUM> x <NUM> pixel image) instead of using all the pixels from the detector since the X-ray beam is collimated in the corners (creating a "squircle") For a given X-ray spectrum, the average pixel intensity of an X-ray detector is known to be linearly proportional to the detector entrance dose as measured by the dosimeter <NUM>. In other words, the detector pixel intensity is linearly proportional to the detector entrance dose for any given X-ray tube electrical parameters such as tube voltages kVps. Therefore, for a given tube voltage kVp, we can establish the relationship between the tube voltage kVp and the average detector pixel intensity by dividing the average detector pixel intensity by the detector entrance dose and the ABS digital gain at the time of the measurement of the average detector pixel intensity.

<FIG> shows a graphical plot <NUM> of an exemplary relationship between the X-ray tube electrical parameter and an average pixel value normalized by detector entrance dose. A horizontal axis <NUM> of the plot <NUM> represents a tube voltage in kVp whereas a vertical axis <NUM> represents average pixel value with unity dose in pixel counts. The average pixel value with unity dose is determined by dividing the average pixel value by the detector entrance dose and the ABS digital gain at the time of the measurement of the average detector pixel intensity.

In one embodiment, the plot <NUM> between the X-ray tube electrical parameter with unity dose as well as unity gain and average pixel value is represented in terms of a mathematical equation. For example, as earlier, let the kVp be the actual kVp in the projection and let kVp(n) and kVp(n + <NUM>) be the two consecutive kVp values obtained during the air calibration. Also assume that P(n) and P(n + <NUM>), respectively, are the two detector average pixel values corresponding to kVp(n) and kVp(n + <NUM>) obtained from the air calibration. The detector average pixel value P for the projection corresponding to kVp in the projection is then given by the following linear interpolation: <MAT>.

Turning back to <FIG>, at step <NUM>, a normalized air map based on calibration image data is obtained as a function of the X-ray tube electrical parameter. An air map is basically an X-ray intensity map proportional to the unattenuated X-ray intensity I<NUM>. In general, based on the detector entrance dose D (from Eq. <NUM>) and the average detector pixel value P (from Eq. <NUM>), we can obtain the average pixel value for any triumvirate {kVp, mA, K} defined by the 3D ABS table, where the kVp is the X-ray tube voltage, mA is the X-ray tube current and K is the ABS digital gain. However, in order to include the pixel to pixel variation due to detector pixel gain including the A/D gain, detector scintillator non-uniformity, and X-ray field non-uniformity, we also need to obtain the two-dimensional air map or X-ray intensity map for the X-ray detector.

<FIG> shows a graphical plot <NUM> of a normalized air map with respect to the average pixel value for X-ray tube voltage equal to 80kVp. A horizontal axis <NUM> of the plot <NUM> and a vertical axis <NUM> of the plot represents the coordinates of the image pixels respectively in terms of the location of the pixel. To obtain the normalized air map, for a given tube voltage kVp, the tube current mA is adjusted such that the detector output pixel values are around the middle of the detector measurement range when no object is placed on the X-ray beam path. A sequence of calibration images (i.e., calibration image data) is acquired with the same operating condition. The resulting images are firstly averaged to generate one air image with reduced image noise. As will be appreciated by those skilled in the art, image averaging works on the assumption that the noise in the image is truly random. This way, random fluctuations in the image data gradually even out as one averages more and more images. Then the average pixel value of the obtained air image is calculated. Finally, the normalized air map is obtained by the average pixel value divided by the air image.

In one embodiment, the normalized air map may be represented in terms of a mathematical equation. For example, as earlier, let the kVp be the actual kVp in the projection and let kVp(n) and kVp(n + <NUM>) be the two consecutive kVp values obtained during the air calibration. Also assume that <MAT> and <MAT>, respectively, are the two normalized air maps corresponding to kVp(n) and kVp(n + <NUM>) obtained from the air calibration. The normalized air map <MAT> corresponding to kVp of the projection is then given by the following linear interpolation: <MAT>.

Turning back to <FIG>, at step <NUM>, an air map is generated based on the detector entrance dose D (from Eq. <NUM>), the average detector pixel value P (from Eq. <NUM>) and the normalized air map <MAT> (from Eq.<NUM>) using the following equation: <MAT> where I<NUM>(i,j) is the air map, and (i, j) are the coordinates of the image pixels. Finally, at step <NUM>, a 3D image of the subject is reconstructed based on the air map and the measured object X-ray intensity It(i,j). Together with the actual image data It(i,j), the air map I<NUM>(i,j) is used to determine the x-ray attenuation µt(i,j): <MAT> The x-ray attenuation µt(i,j) is then further used to generate the reconstructed 3D image by the 3D image processor/reconstructor <NUM> as will be appreciated by those skilled in the art.

<FIG> shows a pictorial diagram <NUM> of a comparison of a reconstructed image and a ground truth image in accordance with an embodiment of the present technique. In general, the image <NUM> represents a reconstructed image and the image <NUM> represents the ground truth image, both with an offset of +<NUM>. In other words, both the reconstructed image and the ground truth image are represented in a "shifted" Hounsfield Units (sHU). As can be seen, the reconstructed image <NUM> is almost the same as the ground truth image <NUM>. For example, the ground truth image shows the CT number for the bone as <NUM>, for air as <NUM> and for water as <NUM>, whereas the reconstructed image shows the CT numbers for these same elements as <NUM>, <NUM> and <NUM> respectively.

It may be noted that the foregoing examples, demonstrations, and process steps that may be performed by certain components of the present systems, for example, by the control mechanism <NUM>, the DAS <NUM>, the computing device <NUM>, the processor <NUM> and/or the image reconstructor <NUM> may be implemented by suitable code on a processor-based system, such as a general-purpose or special-purpose computer. It may also be noted that different implementations of the present technique may perform some or all of the steps described herein in different orders or substantially concurrently, that is, in parallel.

Additionally, the functions may be implemented in a variety of programming languages, including but not limited to Ruby, Hypertext Preprocessor (PHP), Perl, Delphi, Python, C, C++, or Java. Such code may be stored or adapted for storage on one or more tangible, machine-readable media, such as on data repository chips, local or remote hard disks, optical disks (that is, CDs or DVDs), solid-state drives, or other media, which may be accessed by the processor-based system to execute the stored code.

Claim 1:
A system (<NUM>) for imaging an object comprising:
an X-ray source (<NUM>) operative to transmit X-rays through the object;
a detector (<NUM>) operative to receive the X-ray energy of the X-rays after having passed through the object and to generate corresponding object X-ray intensity;
a dosimeter (<NUM>) operative to measure a detector entrance dose at an entrance of the detector (<NUM>); and
a controller (<NUM>) operative to:
measure, using the dosimeter (<NUM>), the detector entrance dose with no object being placed on the X-ray beam path;
determine a relationship between an X-ray tube electrical parameter and the detector entrance dose;
determine a relationship between the X-ray tube electrical parameter, the detector entrance dose and a detector average pixel intensity with no object being placed on the X-ray beam path;
obtain a normalized air map as a function of the X-ray tube electrical parameter based on a normalized average of a series of calibration images acquired by the detector (<NUM>) with no object being placed on the X-ray beam path;
generate an air map based on the normalized air map, the detector entrance dose and the detector average pixel intensity; and
reconstruct an image of the obj ect based on the air map and the measured object X-ray intensity.