Abstract:
A method, system, and machine readable media are provided for correcting scatter in an image. The method comprises sequentially emitting radiation from a subset of radiation sources toward a detector array and measuring radiation on areas of the detector array not exposed to the emitted primary radiation at the time of measurement. Scatter is estimated from the measured radiation. Furthermore, the scatter estimates are subtracted from the measured data and images with improved image quality are reconstructed.

Description:
BACKGROUND 
   The present invention relates generally to the field of non-invasive imaging and more specifically to the field of computed tomography. In particular, the present invention relates to improving image quality by estimating and reducing scatter in an X-ray imaging system. 
   CT scanners operate by emitting fan-shaped or cone-shaped X-ray beams from an X-ray source towards a detector. The X-ray source emits X-rays at numerous angular positions relative to an object being imaged, such as a patient, which attenuates the X-ray beams as they traverse the object. The attenuated X-ray beams are detected by a set of detector elements, which produce signals representing the attenuation of the incident X-ray beams. The signals are processed to produce data corresponding to the line integrals of the attenuation coefficients of the object along X-ray paths connecting the source and detector elements. These signals are typically called “projection data” or just “projections”. By using reconstruction techniques, such as filtered backprojection, useful images may be formulated from the projections. The images may in turn be associated to form a volume rendering of a region of interest. In a medical context, pathologies or other structures of interest may then be located or identified from the reconstructed images or rendered volume. 
   It is generally desirable to develop CT scanners with high spatial and temporal resolution, good image quality, and good coverage along the z-axis, i.e., the longitudinal or rotational axis of the CT scanner. To meet some or all of these objectives, it may be desirable to increase the coverage provided by the detector, thereby allowing greater scan coverage in one or more dimensions. For example, z-axis coverage of the detector may be lengthened by increasing the number of rows of detector elements in the detector. 
   However, various physical factors associated with the X-ray imaging process may lead to artifacts in the resulting images or to blurring or generally poor image quality. For example, X-rays photons emitted through the imaging volume may pass through the patient or other object being imaged or be absorbed by the patient or object and thus never reach the detector. The amounts of X-ray photons passing through the patient and the amount attenuated are useful to produce the desired radiographic images as this information is indicative of the composition and structure of the patient or object undergoing imaging. At operating voltages of typical X-ray systems, less than 1 megavolt, three dominate absorption processes contribute to the mass attenuation coefficient of the object: photoelectric absorption, Rayleigh scattering, and Compton scattering. Photoelectric absorption is a mechanism where the energy of the photon is absorbed by the material&#39;s electrons and liberated. Rayleigh scattering is an interaction between the photon and material&#39;s electrons, where the photon direction is slightly altered, without any loss of energy. Compton scatter is an interaction where the material absorbs part of the energy of the photon; however, the photon continues to traverse the object or patient along an altered direction. Unlike X-ray photons that are photo-electrically absorbed or undergo Rayleigh scattering, an X-ray photon that is attenuated by the Compton scattering mechanism may eventually reach the detector apparatus but typically along a different trajectory. As a result, a scattered X-ray photon may impact the detector at a location or from a direction that conveys no useful composition or structural information about the patient or object undergoing imaging. As a result, the scattered X-ray photons may lead to blur within the resulting radiographic image or otherwise reduce the image quality, such as CT number nonuniformity or a reduction in the contrast-to-noise ratio in a reconstructed image. The likelihood of such scattering may be increased in imaging systems employing multiple X-ray sources or emission points or increased coverage on the patient or object being imaged. 
   In order to reduce scatter, collimators or anti-scatter grids may be used, which are focally aligned to the X-ray beams from the sources to the detector elements, with a corresponding increase in mechanical complexity and cost of the overall CT system. Further, use of collimators with higher resolution detectors has proven challenging due to the small size of the detector elements or pixels. An alternative method of estimating scatter by attempting to extrapolate scatter signals from detector elements at opposing lateral sides of the detector array has proven difficult and does not provide a reliable estimate for scatter across the full axial volume. A technique for reducing scatter in X-ray imaging while reducing the mechanical complexity and cost of the imaging system is therefore desirable. 
   BRIEF DESCRIPTION 
   A method for correcting scatter in an image is provided. The method includes the act of emitting radiation from one or more sources of radiation towards a detector array and measuring radiation on areas of the detector array that are not exposed to the emitted primary radiation at the time of measurement. Scatter is estimated from the measured radiation. Corresponding claims to tangible, machine readable media comprising code executable to perform these acts are also provided. 
   An imaging system is provided. The imaging system includes one or more radiation sources along the z-axis configured to emit a beam of radiation. The imaging system also includes a detector array comprising a plurality of detector elements. The detector array may generate one or more signals in response to the respective beams of radiation. The imaging system also includes a system controller configured to control the radiation sources, including activating subsets of the radiation sources sequentially. In addition, the system controller is configured to acquire the one or more signals from the plurality of detector elements, including detector elements not exposed to the primary beams of radiation at the time of acquisition. The imaging system also includes a computer system configured to estimate scatter over the entire detector array. 

   
     DRAWINGS 
     These and other features, aspects, and advantages of the present invention will become better understood when the following detailed description is read with reference to the accompanying drawings in which like characters represent like parts throughout the drawings, wherein: 
       FIG. 1  is a diagrammatical view of an exemplary imaging system in the form of a CT imaging system for use in producing processed images, in accordance with one aspect of the present technique; 
       FIG. 2  depicts the activation of a subset of X-ray emission points positioned along the z-axis of an exemplary CT imaging system in accordance with one aspect of the present technique; 
       FIG. 3  depicts the activation of a second subset of X-ray emission points positioned along the z-axis of an exemplary CT imaging system in accordance with one aspect of the present technique; 
       FIG. 4  depicts the activation of a third subset of X-ray sources positioned along the z-axis of an exemplary CT imaging system in accordance with one aspect of the present technique; and 
       FIG. 5  is a flowchart depicting exemplary actions for processing images in accordance with the present technique for estimating and reducing scatter. 
   

   DETAILED DESCRIPTION 
     FIG. 1  illustrates diagrammatically an imaging system  10  for acquiring and processing image data. In the illustrated embodiment, system  10  is a computed tomography (CT) system designed to acquire X-ray projection data, to reconstruct the projection data into an image, and to process the image data for display and analysis in accordance with the present technique. Though the imaging system  10  is discussed in the context of medical imaging, the techniques and configurations discussed herein are applicable in other non-invasive CT imaging contexts, such as baggage or package screening. In the embodiment illustrated in  FIG. 1 , CT imaging system  10  includes a source  12  of X-ray radiation. As discussed in detail herein, the source  12  of X-ray radiation may be any source configured to emit X-rays from one or more z-locations or emission points  13 . For example, the X-ray source  12  may consist of multiple X-ray tubes arranged at different locations along the z-axis. Similarly, the X-ray source  12  may include one or more addressable solid-state sources. Such solid-state sources may be configured as arrays of field emitters, including one-dimensional arrays, i.e., lines, and two-dimensional arrays. Although three emission points  13  are shown in  FIG. 1 , source  12  may include one or more emission points  13 . Moreover, the emission points  13  are shown for illustration purposes only. 
   The X-ray source  12  may be positioned proximate to a collimator  14 . The collimator  14  may consist of one or more collimating regions, such as lead or tungsten shutters, for each emission point of the source  12 . The collimator  14  typically defines the size and shape of the one or more beams of radiation  16  that pass into a region in which a human patient  18  is positioned. A beam of radiation  16  may be generally fan or cone-shaped depending on the configuration of the detector array, as discussed below, as well as the desired method of data acquisition. An unattenuated portion of the radiation  20  passes through the subject, which provides the attenuation, and impacts a detector array, represented generally at reference numeral  22 . 
   The detector  22  is generally formed by a plurality of detector elements, which detect the X-rays that pass through or around a subject of interest. Each detector element produces an electrical signal that represents the intensity of the X-ray beam at the position of the element during the time the beam strikes the detector. Typically, signals are acquired at a variety of angular positions around the subject of interest so that a plurality of radiographic views may be collected. These signals are acquired and processed to reconstruct an image of the features within the subject, as described below. 
   The X-ray source  12  is controlled by a system controller  24 , which furnishes power, focal spot location, control signals and so forth for CT examination sequences. Moreover, the detector  22  is coupled to the system controller  24 , which commands acquisition of the signals generated in the detector  22 . The system controller  24  may also execute various signal processing and filtration functions, such as for initial adjustment of dynamic ranges, interleaving of digital image data, and so forth. In general, system controller  24  commands operation of the imaging system to execute examination protocols and to process acquired data. In the present context, system controller  24  also includes signal-processing circuitry and associated memory circuitry. The associated memory circuitry may store programs and routines executed by the computer, configuration parameters, image data, and so forth. The system controller  24  may also be a general purpose or application-specific computer system. 
   In the embodiment illustrated in  FIG. 1 , system controller  24  may control the movement of a linear positioning subsystem  28  and rotational subsystem  26  via a motor controller  32 . In imaging systems  10  in which the source  12  and/or the detector  22  may be rotated, the rotational subsystem  26  may rotate the X-ray source  12 , the collimator  14 , and/or the detector  22  through one or multiple turns around the patient  18 . It should be noted that the rotational subsystem  26  might include a gantry. The linear positioning subsystem  28  enables the patient  18 , or more specifically a patient table, to be displaced linearly. Thus, the patient table may be linearly moved within a gantry that includes a rotating source  12  and detector  22  or within a stationary source  12  and/or detector  22  configuration to generate images of particular areas of the patient  18 . In embodiments comprising a stationary source  12  and a stationary detector  22 , the rotational subsystem  26  may be absent. Similarly, in embodiments in which the source  12  and the detector  22  are configured to provide extended coverage along the z-axis, i.e, the axis associated with the main length of the patient  18 , the linear positioning subsystem  28  may be absent. 
   As will be appreciated by those skilled in the art, the distributed source  12  of radiation may be controlled by an X-ray controller  30  disposed within the system controller  24 . The X-ray controller  30  may be configured to provide power and timing signals to the X-ray source  12  or emission points  13  therein. In addition, the X-ray controller may be configured to selectively activate the distributed X-ray source  12  such that tubes or emission points at different locations along the z-axis may be activated individually or in subsets. 
   Further, the system controller  24  may comprise a data acquisition system  34 . In this exemplary embodiment, the detector  22  is coupled to the system controller  24 , and more particularly to the data acquisition system  34 . The data acquisition system  34  receives data collected by readout electronics of the detector  22 . The data acquisition system  34  typically receives sampled analog signals from the detector  22  and converts the data to digital signals for subsequent processing by a processor-based system, such as a computer  36 . 
   The computer  36  is typically coupled to or incorporates the system controller  24 . The data collected by the data acquisition system  34  may be transmitted to the computer  36  for subsequent processing and reconstruction. For example, the data collected from the detector  22  may undergo pre-processing and calibration at the data acquisition system  34  and/or the computer  36  to process the data to represent the line integrals of the attenuation coefficients of the scanned objects. The processed data, commonly called projections, may then be filtered and backprojected to formulate an image of the scanned area. In one exemplary embodiment, the computer  36  uses data collected from the detector  22  to estimate scatter, such as due to Compton scattering, and then process the scatter signal. Before filtering and backprojection, the computer  36  may subtract the scatter estimate from intensity measurements prior to computing the projection data. Once reconstructed, the image produced by the system of  FIG. 1  reveals an internal region of interest of the patient  18  which may be used for diagnosis, evaluation, and so forth. Alternately, the estimated scatter signals may be fed into an iterative reconstruction algorithm, which incorporates the scatter estimate into a forward model for the data acquisition, thereby implicitly correcting for the scatter in the data. 
   The computer  36  may comprise or communicate with a memory  38  that can store data processed by the computer  36  or data to be processed by the computer  36 . It should be understood that any type of computer accessible memory device capable of storing the desired amount of data and/or code may be utilized by such an exemplary system  10 . Moreover, the memory  38  may comprise one or more memory devices, such as magnetic or optical devices, of similar or different types, which may be local and/or remote to the system  10 . The memory  38  may store data, processing parameters, and/or computer programs comprising one or more routines for performing the processes described herein. 
   The computer  36  may also be adapted to control features enabled by the system controller  24 , i.e., scanning operations and data acquisition. Furthermore, the computer  36  may be configured to receive commands and scanning parameters from an operator via an operator workstation  40  which may be equipped with a keyboard and/or other input devices. An operator may thereby control the system  10  via the operator workstation  40 . Thus, the operator may observe the reconstructed image and other data relevant to the system from computer  36 , initiate imaging, and so forth. 
   A display  42  coupled to the operator workstation  40  may be utilized to observe the reconstructed images. Additionally, the scanned image may be printed by a printer  44  which may be coupled to the operator workstation  40 . The display  42  and printer  44  may also be connected to the computer  36 , either directly or via the operator workstation  40 . Further, the operator workstation  40  may also be coupled to a picture archiving and communications system (PACS)  46 . It should be noted that PACS  46  might be coupled to a remote system  48 , radiology department information system (RIS), hospital information system (HIS) or to an internal or external network, so that others at different locations may gain access to the image data. 
   One or more operator workstations  40  may be linked in the system for outputting system parameters, requesting examinations, viewing images, and so forth. In general, displays, printers, workstations, and similar devices supplied within the system may be local to the data acquisition components, or may be remote from these components, such as elsewhere within an institution or hospital, or in an entirely different location, linked to the image acquisition system via one or more configurable networks, such as the Internet, virtual private networks, and so forth. 
   The CT imaging system  10  described above may be modified or configured in a variety of ways to improve spatial and temporal resolution, to improve image quality, and/or to improve z-axis coverage. Indeed, various source  12  and detector  22  configurations may be implemented which improve one or more of these parameters. In the example embodiment, the distributed X-ray source  12  includes two to fifteen emission points, such as X-ray tubes or field emitters, along the z-axis (typically the axis of rotation or the axis running through the bore of the scanner) to improve or increase z-axis coverage. As will be explained in  FIGS. 2-4 , the emission points are sequentially activated in interleaved subsets. The number of emission points in each subset may be determined by dividing the total number of emission points by any integer between two and the total number of emission points. In the exemplary embodiment depicted in  FIGS. 2-4 , each subset consists of one-third of the total number of emission points in the embodiment. The subsets of emission points are activated such that radiation beams  16  generated by the different emissions points do not concurrently impact the same detector elements of the detector  22 . As a result, some areas on the detector array  22  will be outside the area exposed to the primary radiation beams. As used herein, the term primary radiation refers to radiation emitted from the emission points that would be incident on the detector array without undergoing any of the absorption or scatter processes described above. The only radiation received on those areas of the detector array outside the areas exposed to the primary radiation is scattered radiation, and the signals from these areas, when read by the data acquisition system  24 , are used by the computer  36  to estimate scatter. After every subset of emission points has been activated, the scatter estimations can then be interpolated to provide a scatter estimate along the entire 2D array of detector elements. 
     FIGS. 2-4  illustrate the sequential activation of emission point subsets. In the example embodiment in  FIGS. 2-4 , there are six emission points and three subsets, each subset consisting of two emission points. Although shown as such, any configuration of two or more emission points can be operated in the manner described herein. Beginning with  FIG. 2 , the first sequence of emission point subset activation is shown, including the X-ray source  12  comprising six emission points. The emission points may be collectively enclosed within the same vacuum enclosure or separately contained within separate vacuum vessels. The active emission points  50  represent the first subset of emission points and are the only emission points emitting radiation in the first sequence. The radiation beams  16  pass through the patient or object  18  to be scanned. As discussed above, the majority of the radiation beams impact the detector array  22  at the primary radiation exposure area  52 , but there are areas  54  of the detector array not within the path of the primary radiation beams  16 . Some scattered radiation  56  from the radiation beams  16  impacts the areas  54  of the detector array  22 , and the incidence of X-rays on the areas  54  is then used to estimate scatter of the radiation beams  16 . 
   Turning now to  FIG. 3 , the second activation sequence is shown. The active emission points  58  represent the second subset of emission points. The beams of radiation  16  from the active emission points  58  pass through the patient or object  18  and impact the detector array  22 . The area  54  on the detector array represents an area not within the path of the primary radiation beams  16 , and the areas  54  receive only the scattered radiation  56  from the radiation beams  16 . Signals acquired at the areas  54  are used to estimate scatter during the second activation of the sequence. 
   Referring now to  FIG. 4 , the active emission points  60  represent the third and final subset of emission points in the example. The active emission points  60  emit beams of radiation  16 , which pass through the patient or object  18  and impact the detector array  22 . As all three subsets have now been activated, the entire axial volume has been illuminated by beams of radiation. As in the preceding examples, the areas  54  on the detector array represent areas not within the path of the primary radiation beams  16  that receive only the scattered radiation  56  from the radiation beams  16 . The incidence of X-rays on the areas  54  is then used to estimate scatter during the third activation of the sequence. 
   In  FIG. 5  exemplary acts for reducing the effects of X-ray scatter using the system described in  FIG. 1  are depicted. The acts described by the flowchart in  FIG. 5  may be performed in any configuration of the system described above, so that the scatter estimation may be performed for any number of emission points, subsets, or any combination thereof. In the depicted example an interleaved subset of one or more emission points is sequentially activated (block  70 ) as described in  FIGS. 2-4 . While the subsets of emission points are activated in sequence, scatter is detected (block  72 ) and estimated (block  74 ) for the entire detector array based on the signal strength seen at the respective areas  54  outside the path of the primary radiation beams  16  during each activation sequence. The scatter may be processed (block  76 ) by smoothing the scatter to remove noise, such as with a low-pass filter or other filtering technique. In one implementation the scatter signal is interpolated to provide a scatter estimate across the entire detector on a view-by-view basis. The scatter estimate is subtracted (block  78 ) from the measured intensity data (block  80 ). A final image is reconstructed (block  82 ) from the corrected intensity data. The intensity data may be processed into projection data before reconstruction into a final image. Since the characteristics of scatter do not vary drastically within projection data acquired at adjacent angular positions of the gantry for CT imaging, it may be possible to estimate scatter for every other view position or every third view position, etc., and interpolate the scatter signal to reduce the computational complexity of the scatter correction technique. Alternately, the scatter can be incorporated into a forward model for an iterative reconstruction algorithm. 
   In view of the techniques described above, scatter can be reduced in radiographic images without the use of hardware collimators associated with the detector array  18 , thereby reducing mechanical complexity and cost. Alternatively, scatter can be further reduced in radiographic imaging from system configurations that incorporate hardware collimators alone with detector array  18 . Further, other methods of scatter estimation or reduction for use with wide-cone CT systems, such as narrowing the cone beam or providing two-dimensional scatter grids, may be unnecessary. Collimation, anti-scatter grids, or other methods for reducing scatter or improving image quality can be combined with the techniques above, and such methods may be simplified to reflect the efficacy of the present technique, such as by collimating the radiation beams in only one dimension. 
   While only certain features of the invention have been illustrated and described herein, many modifications and changes will occur to those skilled in the art. For example, though the present discussion has been in the context of medical imaging using radiographic systems, one of ordinary skill in the art will appreciate that the present techniques are equally applicable to radiographic and tomosynthesis systems and also to non-medical imaging applications employing X-ray sources that may move relative to the detection apparatus. For example, the present techniques may also be applied to non-invasive and/or non-destructive imaging techniques used for security and quality control applications in the fields of baggage and package screening, manufacturing quality control, security screening and so forth. Additionally, although the techniques and X-ray source topologies described herein consider multiple spots along the z-axis, the techniques can equally be applied to multiple spots distributed within the axial plane, as long as scatter estimates can be made at a finite number of sample positions on the detector. It is, therefore, to be understood that the appended claims are intended to cover all such modifications and changes as fall within the true spirit of the invention.