Abstract:
A transmission radiation source ( 30   a ) radiates at a plurality of energy levels within a specified energy range. The energy range is divided into two or more energy subranges. Detectors ( 20   a   -20   c ) detect the position or trajectory and energy of transmitted radiation and emitted radiation. A sorter ( 48 ) sorts the detected radiation into the appropriate energy subrange. Data for each subrange is stored in a plurality of transmission data memories ( 50   a   -50   n ). Reconstruction processors ( 52   a   -52   n ) generate a transmission image representation ( 54   a   -54   n ) representative of each energy subrange. A combine processor ( 60 ) weights each energy subrange image representation with an assigned weighting factor ( 64   a   -64   n ) to provide enhancement of at least one feature when the images are combined to generate weighted image representations ( 72, 74, 76 ). The plurality of transmission images are also combined with equal weighting to generate an image representation ( 70 ) used to generate attenuation correction factors ( 80 ) for correcting the emission data ( 46 ). A reconstruction processor ( 84 ) generates a corrected emission image representation ( 86 ). The emission image can be combined with one of the feature-enhanced structural images ( 72, 74, 76 ) using a combiner ( 92 ) and displayed, allowing the functional emission image to be located with respect to structural or anatomical features. Also, a feature-enhanced structural image ( 72, 74, 76 ), can advantageously be used to register the emission image ( 86 ) with an image ( 100 ) from another modality, such as a computed tomography (CT) image.

Description:
BACKGROUND OF THE INVENTION 
     The present invention relates generally to the art of nuclear medicine. It finds particular application to nuclear imaging techniques and apparatuses employing emission and transmission tomography. Although the present invention is illustrated and described herein primarily in reference to positron emission tomography (PET) and single photon emission computed tomography (SPECT), it will be appreciated that the present invention is also amenable to other noninvasive investigation techniques and other diagnostic modes in which a subject or patient is examined with transmitted radiation. 
     Diagnostic nuclear imaging is used to study a radionuclide distribution in a subject. Typically, one or more radiopharmaceuticals or radioisotopes are injected into a subject. The radiopharmaceuticals are commonly injected into the subject&#39;s blood stream for imaging the circulatory system or for imaging specific organs which absorb the injected radiopharmaceuticals. Gamma or scintillation camera detector heads are placed adjacent to a surface of the subject to monitor and record emitted radiation. For SPECT imaging, collimators are typically placed on the detector heads. For PET imaging, a coincidence detector detects concurrent receipt of a radiation event on two oppositely disposed heads. Often, the detector heads are rotated or indexed around the subject to monitor the emitted radiation from a plurality of directions. The monitored radiation data from the multiplicity of directions is reconstructed into a three-dimensional image representation of the radiopharmaceutical distribution within the subject. Such images typically provide functional and metabolic information. 
     Positron emission tomography (PET) is a branch of nuclear medicine in which a positron-emitting radiopharmaceutical such as  18 F-fluorodeoxyglucose (FDG) is introduced into the body of a subject. Each emitted positron reacts with an electron in what is known as an annihilation event, thereby generating a pair of 511 keV gamma rays. The gamma rays are emitted in directions approximately 180° apart, i.e., in opposite directions. 
     A pair of detectors registers the position and energy of the respective gamma rays, thereby providing information as to the position of the annihilation event and hence the positron source. Because the gamma rays travel in opposite directions, the positron annihilation is said to have occurred along a line of coincidence connecting the detected gamma rays. A number of such events are collected and used to reconstruct a clinically useful image. 
     The energy spectrum for clinical positron annihilation imaging is typically characterized by a photopeak at 511 keV. Similarly, Compton scattered radiation contributes to counts having energies ranging as high as the Compton edge. In coincidence imaging, a dual energy window detection scheme is sometimes used. A window around the photopeak and a window in the vicinity of the Compton region are identified. A coincidence event is counted if both detectors detect temporally simultaneous events within the photopeak window, or if one detector observes an event in the photopeak window while the other simultaneously detects an event in the Compton window. In each case, a memory location is incremented to note the event and its location such that the respective events are weighted equally. Events in which both detectors observe Compton events are discarded. 
     Single photon emission computed tomography (SPECT) is another nuclear imaging technique used to study the radionuclide distribution in subjects. Typically, one or more radiopharmaceuticals are injected into a subject. The radiopharmaceuticals are commonly injected into the subject&#39;s blood stream for imaging the circulatory system or for imaging specific organs which absorb the injected radiopharmaceuticals. Gamma or scintillation camera heads are placed closely adjacent to a surface of the subject to monitor and record emitted radiation. In SPECT imaging, the detector head or heads are rotated or indexed around the subject to monitor the emitted radiation from a plurality of directions. The monitored radiation emission data from the multiplicity of directions is reconstructed into a three-dimensional image representation of the radiopharmaceutical distribution within the subject. 
     One of the problems with nuclear imaging techniques such as PET and SPECT is that photon absorption and scatter by portions of the subject between the emitting radionuclide and the camera head(s) distort the resultant image. One solution for compensating for photon attenuation is to assume uniform photon attenuation throughout the subject. That is, the subject is assumed to be completely homogenous in terms of radiation attenuation, with no distinction made for bone, soft tissue, lung, etc. This enables attenuation estimates to be made based on the surface contour of the subject. However, human subjects do not cause uniform radiation attenuation, especially in the chest. 
     In order to obtain more accurate radiation attenuation measurements, a direct measurement is made using transmission computed tomography techniques. In this technique, radiation is projected from a radiation source through the subject. Radiation that is not attenuated is received by detectors at the opposite side. The source and detectors are rotated to collect transmission data concurrently or sequentially with the emission data through a multiplicity of angles. This transmission data is reconstructed into a transmission image representation using conventional tomography algorithms. The radiation attenuation properties of the subject from the transmission image representation are used to correct for radiation attenuation in the emission data. 
     It is desirable to precisely locate the emission images relative to other anatomical details. In so doing, the diagnostic accuracy of the nuclear medicine image is increased. This is particularly so in the area of oncology, in which precise localization of nuclear medicine images aids in surgical and/or radiotherapeutic planning and for assessment of lesion progression and treatment effectiveness. 
     While transmission data has largely been successful in determining attenuation correction factors for the correction of the emission image data, the transmission image data itself has generally been of less than ideal resolution. The coarseness of the images could create uncertainties when localizing the emission image with respect to anatomical features. 
     One method of localizing the functional information is to merge the emission image representation with an image representation generated with another imaging modality that provides anatomical or structural details, such as x-ray computed tomography (CT), magnetic resonance (MR), or ultrasound image representations. When fusing images of different modalities, image registration of the two images is required to correct for any differences in geometric relations between the two images. Any misalignment of the two images impairs the diagnostic value of the fused images. Image registration can be performed by a number of techniques, such as using discrete extrinsic or intrinsic landmarks known to bear a constant relationship to the subject&#39;s anatomy during the two studies, and using three-dimensional surface identification algorithms to construct numerical models of the external surface of the images. Such techniques allow the images to be aligned and oriented with respect to each other by translating, rotating, and descaling one or both of the image representations to allow the images to be superimposed or fused. However, given the lack of structural detail in nuclear medicine emission images and the low resolution of typical transmission images, significant uncertainties can remain when combining a nuclear medicine image with an image from a different modality. Moreover, inconvenience, cost, and multiple scans are generally required when obtaining scans from multiple modalities. 
     Imaging devices which combine a CT-like device with a gamma camera are known in the art. Such a device can reduce scan times by using correlated acquisition of nuclear medicine image data and CT image data. However, such a combined device is a less than optimal solution to the problem of nuclear medicine image localization due to cost and for logistical reasons. Also, although different modalities are combined on a single machine, this type of device retains the conventional approach of addressing separately the need for attenuation correction and the need for precise nuclear medicine image localization. 
     Transmission image quality can also be increased through increasing the number of counts, i.e., by increasing the source activity, increasing the imaging time, or both. Increasing the source activity, however, has the disadvantage of increasing cost and shielding requirements. Increasing the imaging time is generally undesirable for patient handling reasons. Also, both increasing source activity and increasing imaging time undesirably increase the dose of radiation received by the subject. 
     Accordingly, the present invention contemplates a new and improved nuclear medicine imaging method and apparatus which overcome the above-referenced problems and others. 
     SUMMARY OF THE INVENTION 
     In accordance with a first aspect of the present invention, a method of generating an image comprises providing a transmission radiation source emitting gamma rays at a plurality of energy levels and directing the emitted gamma rays through a subject to be imaged, the subject attenuating the transmission of the gamma rays and defining an energy range encompassing the emission energy levels. The energy range is divided into a plurality of energy subranges and gamma rays passing through the subject and falling within the defined energy range are detected. Detector head positions or trajectories and energies of the detected gamma rays are determined and this information is logged into a plurality of image data subsets based on the determined energy of the detected gamma rays, wherein each image data subset corresponds to one of the energy subranges. The steps of detecting, determining, and logging are repeated for a plurality of transmitted rays. The image data subsets are compared to determine one or both of: (1) variations in attenuation between different tissue types of the subject as a function of energy, and (2) variations in attenuation within each tissue type of the subject as a function of energy. Based on the determined attenuation variations, a weighting factor is assigned to each of the image data subsets and the image data subsets are combined in accordance with their assigned weighting factors to produce a weighted image data set, the weighting factors being assigned so as to enhance at least one structural feature in the weighted image data set. A feature-enhanced transmission image representation representative of the weighted image data set is then generated. 
     According to another aspect, a method of diagnostic imaging comprises transmitting radiation with a defined energy spectrum through a subject and converting the transmission radiation which has traversed the subject into electronic transmission data indicative of transmission radiation detector head position or trajectory and energy. In accordance with the energy data, the transmission trajectory data is sorted into a plurality of energy windows and the transmission trajectory data in each window is reconstructed into a corresponding electronic transmission image representation. The electronic transmission image representations are each weighted and the weighted transmission image representations are combined. 
     In yet a further aspect, the present invention provides a gamma camera comprising a transmission radiation source for generating radiation im a selected energy range and a detector for detecting emission radiation emitted from within a subject and transmission radiation from the transmission radiation source which has traversed a subject to be imaged, the subject attenuating the radiation, the detector generating position or trajectory and energy data. The gamma camera further comprises energy discrimination circuitry connected with the detector, the energy discrimination circuitry sorting detected transmission radiation in accordance with a plurality of energy subranges within the selected energy range; an electronic storage medium connected with the energy discrimination circuitry, the electronic storage medium storing a plurality of transmission data subsets, the data subsets comprising data grouped by energy in accordance with the plurality of energy subranges; at least one reconstruction processor connected with the electronic storage medium which generates a transmission image representation for each of the plurality data subsets; and a combine processor connected with the reconstruction processor which weights the transmission image representations and combines the plurality of weighted image representations to produce at least one weighted image representation, the weighting being selected to enhance at least one selected feature in each weighted image representation. 
     One advantage of the present invention is that it provides a transmission image providing increased anatomic structural detail without increasing imaging time or increasing radiation source activity. 
     Another advantage is that the transmission image data is maintained in a proper format for attenuation correction of the emission image data. 
     Another advantage of the present invention is that it provides enhanced transmission image data for improved registration of an emission image representation with an image representation of a different modality. 
     Still further advantages and benefits of the present invention will become apparent to those of ordinary skill in the art upon reading and understanding the following detailed description of the preferred embodiments. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     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 preferred embodiments and are not to be construed as limiting the invention. 
     FIGS. 1A and 1B, taken together as FIG. 1, are a diagrammatic illustration of a first embodiment of a gamma camera in accordance with the present invention. 
     FIGS. 2A and 2B, taken together as FIG. 2, are a diagrammatic illustration of a second embodiment of a gamma camera in accordance with the present invention. 
     FIGS. 3 and 4 are flow charts outlining two exemplary methods of the present invention. 
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
     With reference to FIGS. 1 and 2, a diagnostic imaging apparatus includes a subject support  10 , such as a table or couch, which supports a subject  12  being examined and/or imaged. The embodiments illustrated in FIGS. 1 and 2 can be, but are not necessarily, implemented in a gamma camera system capable of imaging in both PET and SPECT modes. Alternatively, the embodiments of FIGS. 1 and 2 can be implemented in dedicated PET and SPECT gamma camera systems, respectively. The subject  12  is injected with one or more radiopharmaceuticals or radioisotopes such that emission radiation is emitted therefrom optionally, the subject support  10  is selectively height adjustable so as to center the subject  12  at a desired height, e.g., so that the volume of interest is centered. A first or stationary gantry  14  rotatably supports a rotating gantry  16 . The rotating gantry  16  defines a subject receiving aperture  18 . In certain embodiments embodiment, the first gantry  14  is moved longitudinally along the subject support  10  so as to selectively position regions of interest of the subject  12  within the subject receiving aperture  18 . Alternatively, the subject support  10  is advanced and retracted to achieve the desired positioning of the subject  12  within the subject receiving aperture  18 . 
     Detector heads  20   a ,  20   b , and  20   c  are individually positionable on the rotating gantry  16 . The detector heads  20   a - 20   c  also rotate as a group about the subject receiving aperture  18 , and the subject  12  when received, with the rotation of the rotating gantry  16 . The detector heads  20   a - 20   c  are radially, circumferentially, and laterally adjustable to vary their distance from the subject and spacing on the rotating gantry  16 . Separate translation devices  22   a ,  22   b , and  22   c , such as motors and drive assemblies, independently translate the detector heads radially, circumferentially, and laterally in directions tangential to the subject receiving aperture  18  along linear tracks or other appropriate guides. 
     Each of the detector heads  20   a - 20   c  has a radiation receiving face facing the subject receiving aperture  18 . Each head includes a scintillation crystal, such as a large doped sodium iodide crystal, that emits a flash of light or photons in response to incident radiation. An array of photomultiplier tubes receives the light flashes and converts them into electrical signals x,y,z. A resolver circuit resolves the x, y-coordinates of each flash of light and the energy of the incident radiation. That is to say, radiation strikes the scintillation crystal causing the scintillation crystal to scintillate, i.e., emit light photons in response to the radiation. The photons are received by the photomultiplier tubes and the relative outputs of the photomultiplier tubes are processed and corrected to generate an output signal indicative of (i) a position coordinate on the detector head at which each radiation event is received, and (ii) an energy of each event. The energy is used to differentiate between various types of radiation such as multiple emission radiation sources, stray and secondary emission radiation, scattered radiation, transmission radiation, and to eliminate noise. In SPECT imaging, a projection image representation is defined by the radiation data received at each coordinate on the detector head. In PET imaging, the detector head outputs are monitored for coincident radiation events on two or more heads. From the position and orientation of the heads and the location on the faces at which the coincident radiation is received, a ray between the peak detection points is calculated. This ray defines a line along which the radiation event occurred. The radiation data from a multiplicity of angular orientations of the heads is then reconstructed into a volumetric image representation of the region of interest. 
     For SPECT imaging, the detector heads  20   a - 20   c  include mechanical collimators  24   a ,  24   b , and  24   c  (FIG.  2 ), respectively, removably mounted on the radiation receiving faces of the detector heads  20   a - 20   c . The collimators include an array or grid of lead vanes which restrict the detector heads  20   a - 20   c  from receiving radiation not traveling along selected rays in accordance with the selected imaging procedure. In this manner, each detector head at each angular position creates a projection image along rays defined by the collimator. For PET imaging, a SPECT camera without collimators on the detector heads can be employed. 
     In specific reference to FIG. 1, a nuclear medicine imaging apparatus is shown in a configuration for PET imaging in accordance with the present invention. In the embodiment shown, two of the detector heads, e.g.,  20   a  and  20   c , are arranged on the rotating gantry  16  on opposite sides of the receiving aperture  18  in facing relation. The receiving faces of the detectors  20   a  and  20   c  are advantageously aligned in generally parallel planes for receiving the coincidence emission counts. A transmission radiation source  30   a  is mounted to the first detector head  20   a  or the rotating gantry  16  and is collimated such that transmission radiation from the radiation source  30   a  is directed toward and received by the detector head  20   b  positioned across the subject receiving aperture from the radiation source  30   a . In this manner, two detector heads, e.g.,  20   a  and  20   c  are used for detecting emission radiation and the third head, e.g.,  20   b  is used for detecting transmission radiation. 
     It will be recognized that the configuration illustrated in FIG. 1 is illustrative and exemplary only and many additional configurations are contemplated. For example, the present invention may be adapted to a two-head gamma camera system, e.g., wherein the detector head  20   b  is absent, or a three-head detector head systems wherein one of the three detector heads, e.g., the detector head  20   b , is not used. In such cases, one (or more) transmission radiation sources are positioned so pass through the subject receiving aperture  18  to be received by an opposing detector head also receiving the emission radiation, with the emission and transmission radiation being distinguished based on their respective energies. Likewise, gamma camera systems including four or more detector heads are also contemplated. 
     Annihilation radiation events are identified by a coincidence detector  40  which identifies simultaneous scintillations in both heads and passes the x,y coordinates on each head and the angular orientation of each head from a position sensor  42  to a ray processor  44 . Noncoincident and piled-up events are discarded. A ray processor identifies the trajectory or ray corresponding to each coincident event from the x,y coordinates of the scintillations on each head and the position of the heads which is stored in an emission memory  46 . Transmission radiation from the transmission radiation source  30   a  are sorted by a sorter  48  on the basis of relative energies z b  and stored in a plurality of transmission memories  50   a - 50   n . Each of the memories corresponds to a preselected energy range or band. 
     In specific reference now to FIG. 2, a three-head embodiment, in a configuration suitable for SPECT imaging, is illustrated. The apparatus includes the first detector head  20   a , the second detector head  20   b , and the third detector head  20   c  arranged on the rotating gantry  16  spaced from one another around the subject receiving aperture  18 . A radiation source  30   a  is mounted to the first detector head  20   a  such that transmission radiation  32   a  therefrom is directed toward and received by the second detector head  20   b . The radiation source  30   a  is preferably collimated at the source. An optional second radiation source  30   b  can be optionally mounted on another detector head, e.g., the detector head  20   b , in like manner such that transmission radiation therefrom can be directed toward and received by the opposing detector head, e.g., the detector head  20   c . Likewise, it is to be appreciated that in still further embodiments, radiation sources can also be mounted to all three detector heads. In still further embodiments, systems having fewer than three detector heads, e.g., one- or two-head systems, are contemplated. Likewise, gamma camera systems including four or more detector heads are also contemplated. Single photon emission events and transmission radiation from the transmission radiation source  30   a  are sorted by a sorter  48  on the basis of relative energies and stored in an emission memory  46  and a plurality of transmission memories  50   a - 50   n , respectively. 
     Referring again to FIGS. 1 and 2, the radiation source  30   a  preferably contains a radioactive line source, preferably a radionuclide held in a shielded steel cylinder  32   a  which is sealed at the ends. In this configuration, the radioactive source generates a radiation fan beam which passes through the subject receiving aperture. The radiation source can be stepped or rotated around the examination volume with the detector heads to obtain coverage of the volume of interest. The steel cylinder can be adjustably mounted onto the corresponding detector head through a pivoting arm mechanism  34   a  for retraction when the transmission source is in use. Alternatively, the radiation source  30   a  is a bar source, point source, flat rectangular source, disk source, or flood source. 
     The radiation source  30   a  emits gamma radiation across a relatively large energy range. In preferred embodiments, a single radioisotope emitting a plurality of specific energy bands is used, although the use of a plurality of radioisotopes emitting at different energy levels is also contemplated. In a particularly preferred embodiment, the transmission source  30   a  uses  133 barium as the radioactive material. Barium-133 emits gamma radiation mainly at 356 keV, but also at 383 keV and 303 keV, thereby providing a relatively large useful energy band. 
     An angular position sensor  42 , which may be, for example, optical, mechanical, or optomechanical, senses or indexes the position of the rotatable gantry  16  and radial, tangential, and circumferential shifts of the heads, and thus the positions of the detector heads  20   a - 20   c  in space, as it rotates about the subject receiving aperture  18  during data acquisition. The head positions recorded are used for transforming the recorded emission and transmission data into a subject coordinate. 
     The sorter or energy discrimination circuitry  48  first sorts the acquired emission data (FIG. 2) from the transmission data on the basis of the relative energies of the detected emission events. Second, the sorter sorts the transmission data into energy segments (FIGS.  1  and  2 ). 
     The position of detected events having an energy associated with emission events, for example 511 keV for positron annihilation in PET imaging, or characteristic emission energies of the particular radiopharmaceutical used for SPECT imaging, are stored in the emission data memory  46 . 
     As stated above, the transmission radiation source  30   a  employed in accordance with this teaching provides a relatively large usable energy band. The transmission energy spectrum is divided into n contiguous energy bands or windows covering the transmission energy spectrum, wherein n is at least two, and is preferably from two to about eight. Detected transmission events are sorted according to the designated energy windows and stored in a corresponding one of n transmission memories  50   a - 50   n . For example, in a preferred embodiment,  133 Ba is used as the transmission radiation source, which emits at 303 keV, 356 keV, and 383 keV. Accordingly, the emission spectrum from about 300 keV to about 400 keV is sampled and segmented into a plurality of contiguous energy windows, preferably 3, each centered about one of the peaks. 
     Transmission radiation data is preferably collected first such that the emission data can be processed into a corrected image as it is collected. However, it is to be appreciated that the emission and transmission acquisition portions of the imaging operation need not be performed in a set order. In addition, emission and transmission radiation data may be acquired simultaneously. 
     Each of the transmission data memories  50   a - 50   n  corresponding to the plurality of defined energy windows, is connected to an associated one of n reconstruction processors  52   a - 52   n , each of which is in turn connected to an associated one of n transmission image memories  54   a - 54   n . Alternately, a single reconstruction processor can reconstruct all of the energy windows on a time shared basis. The reconstruction processors  52   a - 52   n  reconstruct the transmission data stored in the transmission data memories  50   a - 50   n  to generate n transmission image representations which are stored in the transmission image memories  54   a - 54   n . The reconstruction process can change according to the mode of collection, the nature of the study, and the types of collimators used (i.e., fan, cone, parallel beam, and/or other modes). Each of the transmission image representations contained in memories  54   a - 54   n  are then combined using an image combining processor or circuitry  60 . Control circuitry  62  determines weighting factors or functions  64   a - 64   n , one for each of the transmission image representations contained in memories  54   a - 54   n . The weighting factors  64   a - 64   n  are determined in accordance with selected features or tissue types to be enhanced in the resultant weighted image representation. A summing circuit  66  combines some or all of the image representations  54   a - 54   n  in accordance with the determined weighting factors or functions  64   a - 64   n.    
     A non- or equally-weighted combined transmission image representation is generated, e.g., wherein each of the transmission image representations in the memories  54   a - 54   n  is summed or averaged with equal weighting, and stored in a combined image memory  70 . This equally weighted transmission image representation is used to determine attenuation correction factors which are stored in an attenuation factor memory  80 . An emission data correction processor  82  corrects each emission data in accordance with the attenuation factors. For example, for each ray along which emission data is received, an emission correction processor  82  calculates a corresponding ray through the transmission attenuation factors stored in the memory  80 . Each ray of the emission data is then weighted or corrected by the emission data correction processor  82  inversely with the attenuation factors. The corrected emission data are reconstructed by an emission radiation reconstruction processor  84  to generate a three-dimensional emission image representation that is stored in a volumetric emission image memory  86 . Alternately, the transmission data correction is performed as a part of the reconstruction process. A video processor  104  withdraws selected portions of the data from the image memory  86  to generate corresponding human-readable displays on a video monitor  106 . Typical displays include reprojections, selected slices or planes, surface renderings, and the like. 
     In addition to generating an equally weighted sum or average of the n transmission image representations contained in the memories  54   a - 54   n , one or more feature-enhanced image representations are also generated by varying the weighting factors or functions accordingly. The weighting factors or functions  64   a - 64   n  are determined in accordance with (1) the attenuation characteristics of each of the three tissue types generally present in the imaged volume (i.e., bone, soft tissue, and air (lung)) which vary as a function of the energy of the transmission source material, and (2) the attenuation differences for each tissue type, e.g., bone and soft tissue, as a function of the energy of the transmission source material. By examining the differences between the plurality of transmission image representations corresponding to the plurality of energy windows, weighting factors or functions can be determined that allow the n transmission image representations to be combined so as to enhance certain structural features in the resultant image. Accordingly, the plurality of transmission image data sets stored in memories  54   a - 54   n  are statistically analyzed and weighting factors are determined that emphasize a selected tissue type. 
     One or more feature-enhanced image representations, such as soft-tissue-enhanced, bone-enhanced, and/or air enhanced image representations, are generated and stored in memories  72 ,  74 , and/or  76 , respectively. In certain embodiments, a tissue-enhanced image representation is generated. In other embodiments, a bone-enhanced image representation is generated. In still other embodiments, an air-enhanced image representation showing aerated lung boundaries is generated. Optionally, the enhanced images can be used to correct the emission data. When a plurality of feature-enhanced image representations are generated, an image selection control  90  allows an operator to select one of the enhanced transmission image representations from the memories  72 ,  74 , and  76  to be fused with the corrected emission image representation. A combining circuit or processor  92  produces a combined image representation showing the emission or functional image as well as anatomical or structural features from the selected feature-enhanced transmission image data. This allows features of the functional emission image, such as lesions, to be localized with respect to the subject&#39;s anatomy using structural features shown in the feature-enhanced transmission image. The fused image representation is stored in a memory  94 . The video processor  104  withdraws selected portions of the data from the combined image memory  94  to generate corresponding human-readable displays on the video monitor  106 . In certain embodiments, the image selection control  90  also allows the operator to select a feature-enhanced transmission image representation for display without being combined with the emission image, thereby providing CT-like functionality. 
     Referring now to FIG. 3, a flow chart outlining a method of the present invention wherein an emission image representation is combined with a feature-enhanced transmission image representation is illustrated. The method includes an initial acquisition of emission and transmission data (steps  300  and  302 ), as described above. The transmission data is sorted into n energy windows based on the energy of the recorded transmission events (step  304 ), wherein n is an integer of 2 or greater, preferably from 2 to about 8, most preferably 3. A non-weighted transmission image representation is generated (step  308 ) by reconstructing the transmission data from each energy window and taking a non-weighted sum or average of the resultant image representations. The non-weighted transmission image representation is used to generate attenuation correction factors for the emission image data (step  312 ). The emission image representation is reconstructed using the calculated attenuation correction factors (step  316 ). Weighted sums or averages of the reconstructed transmission image representations from the n energy windows are used to provide feature-enhanced transmission image representations (step  320 ), as detailed above. The reconstructed emission image representation and the feature-enhanced transmission image representation are then superimposed (step  324 ). 
     Referring again to FIGS. 1 and 2, an image registration processor  102  registers the emission image representation  86  with a digital image representation  100  acquired from another imaging modality, such as a CT, MR, or ultrasound image representation. In a preferred embodiment, the other modality image representation  100  is a CT image representation. Common structural features are detected in one of the enhanced transmission image representations saved in the memories  72 ,  74 , or  76 , and in the other modality image representation  100 . The detected common features are then used to map or correlate the functional image representation  86  to the other modality image representation  100 , e.g., by calculating appropriate rotation, translation, and scaling factors. Since structural or anatomical features such as air boundaries, soft-tissue features, and bone structure, are more accurately represented in the corresponding enhanced image representations, and the use of the enhanced transmission image representation in accordance with this teaching provide more accurate registration of the emission image representation with the other modality image representation. In preferred embodiments, the bone-enhanced image representation  76  is used to register the emission image representation  86  and the other modality image representation  100 . After registration, a combiner  96  then fuses or superimposes the registered emission and other modality image representations, and the resultant fused image representation is stored in a memory  98 . The video processor  104  withdraws selected portions of the data from the combined image memory  98  to generate corresponding human-readable displays on a video monitor  106 . 
     Referring now to FIG. 4, a flow chart is illustrated outlining a method of the present invention wherein an emission image representation is registered with an image representation from another imaging modality, using a feature-enhanced transmission image representation in accordance with this teaching to perform the registration. The method includes an initial acquisition of emission (step  402 ) and transmission data (step  400 ), as described above. Image data is acquired from another imaging modality (step  403 ) and reconstructed to form an image representation (step  406 ). The transmission data is sorted into n energy windows based on the energy of the recorded transmission events (step  404 ), wherein n is an integer of 2 or greater, preferably from 2 to about 8. A non-weighted transmission image representation is generated (step  408 ) by reconstructing the transmission data from each energy window and taking a non-weighted sum or average of the resultant image representations. The non-weighted transmission image representation is used to generate attenuation correction factors for the emission image data (step  412 ). The emission image representation is reconstructed using the calculated attenuation correction factors (step  416 ). Weighted sums or averages of the reconstructed transmission image representations from the n energy windows are used to provide a feature-enhanced transmission image representation (step  420 ), as detailed above. The feature-enhanced transmission image representation is used to register the emission image representation with the other modality image representation (step  422 ). The other modality image representation and the emission image representation are then superimposed (step  424 ). The fused image is then output to a display (step  428 ). 
     The invention has been described with reference to the preferred embodiments. Modifications and alterations will occur to others upon a reading and understanding of 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.