Patent Publication Number: US-2011060566-A1

Title: Method and apparatus for scatter correction

Description:
The present application relates generally to the imaging arts and more particularly to an apparatus and method for scattered photon correction. It finds use in X-ray imaging (using X-ray photons), Computer Tomography or CT imaging (using X-ray photons), and other kinds of systems such as image-guided radiation therapy systems. 
     Such imaging processes generally include a radiation source which produces imaging photons. The photons pass through the imaged subject to be collected or counted by a photon detector. Data generated by the photon detector is then electronically processed to generate an image of the subject. Two types of photons reach the photon detector. The first are “primary” photons, which are generated by the photon source and travel on a straight line path through the imaged subject to reach the photon detector. The second are “scattered” photons, including photons which are generated by the photon source but which get redirected off of a straight line path during their travel to the photon detector, and also including extraneous background photons which were not actually generated by the photon source. Scattered photons can introduce error into the image reconstruction process. Therefore, to generate highly accurate images of the subject, data generated by the photon detector as a result of scattered photons is typically discounted or corrected for during the image reconstruction process. 
     According to one aspect of the present invention, a method and apparatus are provided for improved photon scatter correction. 
     According to a particular aspect of the present invention, an imaging method is provided. A direct physical measurement of scattered photons, as well as a model of the photon scattering process, are used in conjunction during image reconstruction to correct for photon scatter in generating an image. This method may additionally provide a correction for low frequency drop. 
     According to another aspect of the present invention, an imaging apparatus is provided. The imaging apparatus has a photon source and a photon detector. The photon detector has two regions. A first, imaging region of the photon detector receives photons traveling along flight paths leading on a straight line path back to the photon source. A second, scatter region of the photon detector is closed to such photons by a shutter, but is open to other photons. The measurement of scattered photons received by the second, scatter region of the photon detector may then be used in conjunction with a model of the photon scattering process during image reconstruction to correct for scattered photons in the imaging data collected from the first, imaging region of the photon detector. 
     One advantage resides in a more accurate and robust scatter correction, reducing the risk of visible scatter artifacts appearing in images. Another advantage resides in producing more useful X-ray, CT, PET, SPECT or other images. Numerous additional advantages and benefits will become apparent to those of ordinary skill in the art upon reading the following detailed description of preferred embodiments. 
    
    
     
       The invention may take form in various components and arrangements of components, and in various process operations and arrangements of process operations. The drawings are only for the purpose of illustrating preferred embodiments and are not to be construed as limiting the invention. 
         FIG. 1  is a schematic representation of an imaging system; 
         FIG. 2  schematically illustrates the X-ray detector and shutter used in the imaging system of  FIG. 1 ; 
         FIG. 3  schematically illustrates an alternative X-ray detector and shutter combination; 
         FIG. 4  illustrates a process to correct for scattered photons in generating images; 
         FIGS. 5A to 5D  are representative images which may be used in association with the process of  FIG. 4 . 
     
    
    
     The imaging method and apparatus of the present application are directed generally to any imaging system which corrects for scattered photons. One example of such an apparatus is the imaging system  100  shown in  FIG. 1 , which is particularly useful in generating CT images. As already mentioned, the imaging method and apparatus disclosed here have application in various other kinds of imaging systems. 
     As illustrated in  FIG. 1 , a couch or other suitable object support  102  supports an object under examination  104  in an examination region  106 . An X-ray source  108  such as an X-ray tube, and an X-ray detector  110  such as a flat panel area detector, are provided. The X-ray source  108  and X-ray detector  110  are mounted on a common rotating gantry (not shown) having a center of rotation  112 . The X-ray source  108  and X-ray detector  110  together rotate with the gantry around the support  102  and the imaged subject  104 . In that way, imaging measurements may be taken of the transverse field of view (FOV)  114 , the center of which corresponds to the center of rotation  112 . The X-ray beam  116  generated by the X-ray source  108  has a central ray or projection  118  which is perpendicular to the transverse center  120  of the X-ray detector  110  and is displaced from the center of rotation  112  by a distance d. If d is greater than 0, as shown for example in  FIG. 1 , then the X-ray detector  110  is in an “offset” configuration. If d equals 0, then the central ray  118  passes through the center of rotation  112 , and the X-ray detector  110  is in a “central” configuration. 
     A collimator  122  is mounted proximate to the X-ray detector  110 , between the detector  110  and the examination region  106 , to reduce the amount of scattered photons received by the detector  110 . In general, collimators operate to filter the streams of incoming photons so that only photons traveling in a specified direction are allowed through the collimator. Which direction(s) are permitted through which portion(s) of the collimator is determined in accordance with the data type being collected (for example, whether the X-ray source  108  or other photon source is configured to produce a parallel beam, fan beam, and/or cone beam). The collimator  122  shown in  FIG. 1  includes a plurality of lamellae focused on the X-ray source  108 . If the X-ray source  108  is a line source extending generally parallel to the rotation axis  112 , then the X-ray beam  116  will be a “fan beam.” In that event, the lamellae of the collimator  122  will be transversely symmetric with respect to the transverse center  120  of the detector  110 . If the X-ray source  108  is a point source, then the X-ray beam  116  will be a “cone beam.” In that event, the lamellae of the collimator  122  will vary in both the transverse and axial directions to point back to the point source. 
     The X-ray detector  110  may include, for example, a scintillator that emits a secondary flash of light or photons in response to the incident X-ray photons  116 , or optionally can be a solid state direct conversion material (e.g. CZT). In the former instance, an array of photomultiplier tubes or other suitable photodetectors in the detector  110  receives the secondary light and converts it into electrical signals. The X-ray detector  110  records multiple two dimensional images (also called projections) at different points around the imaged subject  104 . That X-ray projection data is stored by an imaging data processor  124  in a memory  126 . Once all the X-ray projection data is gathered, it may be electronically processed by the imaging data processor  124 . The processor  124  generates an image of the subject  104 , according to a mathematical algorithm or algorithms, which can be displayed on an associated display  128 . A user input  130  may be provided for a user to control the processor  124 . 
     The aforementioned functions can be performed as software logic. “Logic,” as used herein, includes but is not limited to hardware, firmware, software and/or combinations of each to perform a function(s) or an action(s), and/or to cause a function or action from another component. For example, based on a desired application or needs, logic may include a software controlled microprocessor, discrete logic such as an application specific integrated circuit (ASIC), or other programmed logic device. Logic may also be fully embodied as software. 
     “Software,” as used herein, includes but is not limited to one or more computer readable and/or executable instructions that cause a computer or other electronic device to perform functions, actions, and/or behave in a desired manner. The instructions may be embodied in various forms such as routines, algorithms, modules or programs including separate applications or code from dynamically linked libraries. Software may also be implemented in various forms such as a stand-alone program, a function call, a servlet, an applet, instructions stored in a memory such as memory  126 , part of an operating system or other type of executable instructions. It will be appreciated by one of ordinary skill in the art that the form of software is dependent on, for example, requirements of a desired application, the environment it runs on, and/or the desires of a designer/programmer or the like. 
     The systems and methods described herein can be implemented on a variety of platforms including, for example, networked control systems and stand-alone control systems. Additionally, the logic, databases or tables shown and described herein preferably reside in or on a computer readable medium such as the memory  126 . Examples of different computer readable media include Flash Memory, Read-Only Memory (ROM), Random-Access Memory (RAM), programmable read-only memory (PROM), electrically programmable read-only memory (EPROM), electrically erasable programmable read-only memory (EEPROM), magnetic disk or tape, optically readable mediums including CD-ROM and DVD-ROM, and others. Still further, the processes and logic described herein can be merged into one large process flow or divided into many sub-process flows. The order in which the process flows herein have been described is not critical and can be rearranged while still accomplishing the same results. Indeed, the process flows described herein may be rearranged, consolidated, and/or re-organized in their implementation as warranted or desired. 
     As already discussed, the collected projection data generally contains inaccuracies caused by scattered X-rays. The imaging system  100  geometry shown in  FIG. 1  can be highly susceptible to X-ray scattering, for at least two reasons. First, the X-ray detector  110  is offset from the X-ray source  108  by a distance d which is greater than zero. Second, the X-ray source  108  emits X-rays in a cone-beam configuration. The X-ray scattering of such a configuration, and other configurations, may be corrected for as follows. 
     A mathematical algorithm is applied to the projection data collected by the X-ray detector  110  to correct for X-ray scatter and generate sufficiently accurate CT images. That mathematical algorithm applies a model of photon scattering. The model may be a physical model, based on assumptions or estimates regarding the physical space between the X-ray source  108  and the X-ray detector  110 , including the subject  104 . One such algorithm is disclosed in PCT Application Publication WO 2007/148263 entitled “Method and System for Error Compensation.” That application is incorporated herein by reference for its disclosure of photon scatter compensation based on a physical model. Other algorithms may be used to correct for photon scatter, including the disclosures of:
     J. Wiegert, M. Bertram, G. Rose and T. Aach, “Model Based Scatter Correction for Cone-Beam Computed Tomography”, Medical Imaging 2005: Physics of Medical Imaging, Proceedings of SPIE Vol. 5745 (2005), at 271-82;   J. Wiegert, M. Bertram, D. Schafer, N. Noordhoek, K. de Jong, T. Ach and G. Rose, “Soft Tissue Contrast Resolution Within the Head of Human Cadaver by Means of Flat Detector Based Cone-Beam CT”, Medical Imaging 2004: Physics of Medical Imaging, Proceedings of SPIE Vol. 5368 (2004), at 330-37;   M. Bertram, J. Wiegert and G. Rose, “Scatter Correction for Cone-Beam Computed Tomography Using Simulated Object Models”, Medical Imaging 2006: Physics of Medical Imaging, Proceedings of SPIE Vol. 6142 (2006), at C-1 to C-12.   L. A. Love and R. Kruger, “Scatter Estimation for a Digital Radiographic System Using Convolution Filtering”, Med. Phys. Vol. 14, No. 2 (March/April 1987), at 178-85;   B. Ohnesorge, T. Flohr and K. Klingenbeck-Regn, “Efficient Object Scatter Correction Algorithm for Third and Fourth Generation CT Scanners”, Eur. Radiol. Vol. 9 (1999), at 563-69;   M. Zellerhoff, B. Scholz, E. P. Rührnschopf and T. Brunner, “Low Contrast 3D-Reconstruction from C-Arm Data”, Medical Imaging 2005: Physics of Medical Imaging, Proceedings of SPIE Vol. 5745 (2005), at 646-55;   V. Hansen, W. Swindell and P. Evans, “Extraction of Primary Signal from EPIDs Using Only Forward Convolution”, Med. Phys. Vol. 24, No. 2 (September 1997), at 1477-84;   J. Seibert and J. Boone, “X-Ray Scatter Removal by Deconvolution”, Med. Phys. Vol. 15, No. 4 (July/August 1988), at 567-75; and   L. Spies, M. Ebert, B. Groh, B. Hesse and T. Bortfeld, “Correction of Scatter in Megavoltage Cone-Beam CT”, Phys. Med. Biol. Vol. 46 (2001), at 821-33.   

     Those sources are hereby incorporated by reference for their respective disclosures of photon scatter correction models and algorithms. Such models and algorithms may be applied using the processor  124  and memory  126  described above. 
     Such scatter correction models and algorithms may be used in conjunction with a direct physical measurement of scattered photons. For example, as shown in  FIG. 1 , a shutter mechanism  132  may be disposed between the X-ray source  108  and the examination subject  104 . The shutter mechanism  132  operates to block the X-ray beam  116  except for an aperture  134  provided in the shutter mechanism  132 . The size of the aperture  134  may be adjustable. In the configuration illustrated in  FIG. 1 , the shutter mechanism  132  prevents the X-ray beam  116  from reaching a lateral border  136  of the X-ray detector  110 . That border  136  is positioned approximately behind the center of the imaged subject  104  and near the center of rotation  112 . 
     Turning now to  FIG. 2 , the X-ray detector  110  of  FIG. 1  is shown with the shutter mechanism  132 . The collimator  122  is not shown in  FIG. 2 . The X-ray detector  110  is divided into two regions: an imaging region  210  and a scatter region  220 . The imaging region  210  of the X-ray detector  110  corresponds to the aperture  134  in the shutter mechanism  132 , so it receives photons traveling along flight paths leading on a straight line path back to the X-ray source  108 . In other words, the imaging region  210  is open to primary photons as well as to scattered photons which approach the collimator  122  and detector  110  along the same flight paths as primary photons. The shutter mechanism  134  prevents the scatter region  220  of the X-ray detector  110  from receiving primary photons, but the scatter portion  220  is open to other photons. Thus, the scatter region  220  is open to scattered photons but not to primary photons. 
     There is not necessarily any difference in structure or operation of the X-ray detector  110  in the imaging region  210  and the scatter region  220 . Rather, the imaging region  210  of the X-ray detector  110  will count primary photons as well as scattered photons which approach the X-ray detector  110  along the same flight path as primary photons. And the scattered region  220  of the X-ray detector  110  will count scattered photons, but not primary photons. Of course, alternatively the X-ray detector  110  may be two separate X-ray detectors with one in each region  210 ,  220 . 
     As shown in  FIG. 2 , the scatter region  220  of the X-ray detector  110  is one contiguous region of the detector  110 , extending across the entire width W and a portion of the length L. When the detector  110  is placed in the imaging system  100 , the scatter region  220  may be advantageously positioned approximately behind the center of the imaged subject  104  and near the center of rotation  112  along the lateral border  126 , as illustrated in  FIG. 1 . 
     The scatter region of the X-ray detector need not be entirely contiguous like the representative scatter region  220  shown in  FIG. 2 . For example,  FIG. 3  shows an X-ray detector  300  having an imaging region  310  and a scatter region  320  including two non-contiguous sub-portions  320   a  and  320   b . The sub-portions  320   a ,  320   b  are disposed at opposing lateral borders of the detector  300 . This configuration is especially useful for an X-ray detector  300  meant for use in a CT apparatus with a center detector arrangement, such as for example a C-arm arrangement. Any number of non-contiguous sub-portions may be used to form a scatter portion in a photon detector. 
     Yet other configurations are of course possible. The scatter region of the photon detector may be located along the entire border of the detector (e.g., all four sides of a rectangular detector). Or it may be a polka dot pattern, for example. The amount of overall detector area devoted to the scatter region should optionally be large enough to help compensate for low frequency drop or LFD (discussed further below) yet small enough to leave a sufficiently large area remaining for the imaging region to generate a useful image. It has been found that, in a rectangular detector  110  such as shown in  FIG. 2  wherein L equals about 38 cm and W equals about 29 cm, a scatter region  220  extending along the entire width and about 2 cm of the length is sufficient. 
     The direct physical measurement of scattered photons striking the scatter region of the photon detector may be used during image reconstruction to correct for scattered photons in the imaging data recorded in the imaging region of the photon detector. Generally, the scatter region of the photon detector collects substantially only scattered photons. The scatter region of the photon detector then generates an electronic signal reflecting only such scattered photons. The direct physical measurement of scattered photons may be used to estimate the contribution of scattered photons to other areas of the photon detector. That estimate may then be subtracted or divided from the signal produced by the photon detector in those areas to correct for scattered photons and generate a more accurate image. 
     For example, such a process  400  is shown in  FIGS. 4 and 5A  to  5 D. Initially, as shown in  FIG. 4 , raw image data  410  is collected by rotating the X-ray source  108  and X-ray detector  110  with the collimator  122  around the imaged subject  104 . The data  410  is a collection of several two-dimensional projection images recorded by the X-ray detector  110  at various imaging positions disposed around the subject  104 . One of those projections is then selected to undergo the process  400  to correct the selected projection&#39;s imaging data  420  for scatter. Once all such projections have been corrected for scatter, the projections are then processed together as a whole to generate a final image. 
     Often, a single projection image  420  may initially be corrected for low frequency drop (LFD) within the X-ray detector  110  to obtain an LFD-corrected projection image  430 . LFD results from photons scattering within the scintillator component of the X-ray detector  110 . LFD can strongly falsify the signals recorded by the X-ray detector  110 , especially portions of the detector nearby large incident X-ray intensity. Although LFD corrections may be made in the imaging region  210  and in the scatter region  220  of the X-ray detector  110 , they are especially useful in the scatter region  220  due to the relatively low amounts of photons in that region  220 . Thus, it is typically advantageous to place the scatter region  220  in an area of the X-ray detector  110  which is sufficiently far from areas with high incident X-ray intensity. Using the geometry shown in  FIG. 1 , that condition is usually met for the lateral border  126  of the X-ray detector  110  positioned approximately behind the center of the imaged subject  104  and near the center of rotation  112 . That border  126  is subject to a relatively low intensity of X-rays because it lies in the shadow of the object support  102  and/or the object  104 . Thus, the X-ray detector  110  of  FIG. 2  is particularly useful in connection with the imaging system  100  of  FIG. 1  if the scatter region  220  lies along the border  126 . Other configurations will be better suited for use in connection with other imaging system geometries. For example, the X-ray detector  300  of  FIG. 3  can be well suited for use in connection with a center detector arrangement such as for example a C-arm arrangement. 
       FIG. 5A  shows a representative projection image  420  or  430 , taken using a CT system having the geometry of the system  100  and using a shutter  132  and X-ray detector  110 . The dotted region  510  in the image  420  or  430  corresponds to the scatter region  220  of the X-ray detector  110  used to generate the image  420  or  430 . 
     Once a raw image is selected  420 , and LFD corrections have been made to that image (if desired), then a physical or empirical model of the photon scattering process  440  is employed. Representative examples of such a physical model are provided above. Such a physical model  440  advantageously covers at least a portion of the imaging region  210  and at least a portion of the scatter region  220  of the X-ray detector  110 . Using the physical model  440 , a scatter estimate  450  corresponding to the scatter region  220  is calculated for the projection  420  or  430 .  FIG. 5B  shows a representative example of such a scatter estimate  450 , generated using the physical model of WO 2007/148263. The modeled scatter estimate  450  corresponding to the scatter region  220  for the selected projection  420  (e.g., dotted region  520  in  FIG. 5B ) is then compared with the measured data  420  or  430  from the scatter region  220  for the selected projection  420  (e.g., dotted region  510  in  FIG. 5A ). 
     Based on that comparison, the scatter model  440  is globally adjusted over the entire X-ray detector region  210  and  220  to obtain an updated physical scatter model  460 . This adjustment is made in such a way that maximum correspondence is obtained in the scatter region  220  between the updated physical scatter model  460  and the measured data  420  or LFD-corrected data  430 . This may be achieved, for example, by multiplying the initial scatter estimate  450  with a scaling factor that is chosen in such a way so as to minimize the root mean square difference between the scatter estimate  450  and the measured data  420  or  430  in the scatter region  220 . The scaling factors may be weighted to rely more heavily on portions of the region  220  which are believed to be more accurate than other portions.  FIG. 5C  shows a representative example of an updated scatter model  460 , based on the same imaging data used to generate  FIGS. 5A and 5B . 
     Once the improved scatter model  460  is calculated for a particular projection  420  or  430 , that improved model  460  is applied to the imaging projection data  420  or  430  to correct for scattered photons and generate a scatter-corrected projection image  470 . This correction may be carried out, for example, in a subtractive or a multiplicative manner.  FIG. 5D  shows a representative example of such a scatter-corrected projection image  470 . The dotted region  530  in the image  470  corresponds to the imaging region  210  of the X-ray detector  110 . It is the scatter-corrected data corresponding to that region  210  which is later used by the image processor  124  to generate an image of the subject  104 . 
     Once all the projections in the data acquisition have been adjusted according to the process  400  of  FIG. 4 , the scatter-corrected projections  470  are reconstructed together to obtain a tomographic image of the scanned subject  104 , as will be well understood by one of ordinary skill in this art. 
     While the present scatter correction technique is particularly useful in a cone-beam CT apparatus with an offset detector as shown in  FIG. 1 , it has application in other contexts as well. For example, it may be employed to correct for scatter photons in a cone-beam CT apparatus with a centered detector, such as for example C-arms. 
     The invention has been described with reference to the preferred embodiments. Obviously, modifications and alterations will occur to others upon reading and understanding the preceding detailed description. It is intended that the invention be construed as including all such modifications and alterations insofar as they come within the scope of the appended claims or the equivalents thereof. The invention may take form in various components and arrangements of components, and in various steps and arrangements of steps. The drawings are only for purposes of illustrating the preferred embodiments and are not to be construed as limiting the invention.