Abstract: A diagnostic imaging system includes an x-ray source (16), which is rotated around an examination region (20). A subject, disposed on a couch (30), is translated longitudinally through the examination region (20). The x-ray source (16) is pulsed at selected angular location(s), e.g. one or both of 6 and 12 o'clock, to transmit x rays through the subject as it is being translated through the examination region (20). The transmitted radiation is being detected by a radiation detector (22) and is reconstructed by an image processor (52) into a two-dimensional projection pilot scan image. A subject contour is calculated and is used along with the radiation attenuation data by a dose calculator (60) to determine the minimum radiation dose required to produce a constant quality image.

Abstract: A diagnostic imaging system includes an x-ray source (16), which is rotated around an examination region (20). A subject, disposed on a couch (30), is translated longitudinally through the examination region (20). The x-ray source (16) is pulsed at selected angular location(s), e.g. one or both of 6 and 12 o'clock, to transmit x rays through the subject as it is being translated through the examination region (20). The transmitted radiation is being detected by a radiation detector (22) and is reconstructed by an image processor (52) into a two-dimensional projection pilot scan image. A subject contour is calculated and is used along with the radiation attenuation data by a dose calculator (60) to determine the minimum radiation dose required to produce a constant quality image.

Abstract: A radiation detector (30) for a computed tomography scanner (12) includes a plurality of radiation detector modules (32). Each detector module (32) includes an anti-scatter module, at least one radiation absorbing mask (120) and a detector subassembly module (100). The anti-scatter module (32) includes radiation absorbing anti-scatter plates (80). The detector sub-assembly module (100) includes a substrate (102) and an array (104) of detector elements. The radiation absorbing mask (120) is a photoetched grid, formed of a radiation absorbing material and is positioned between the anti-scatter module (78) and the detector elements of array (104). The strip of the grid, that is parallel to the anti-scatter plates (80), is wider than each anti-scatter plate (80). The detector module (32) is aligned with a spatial focus (74) by inserting the alignment pins (160) into the alignment openings (128) of the radiation absorbing mask (120) and the alignment openings (162) of the detector subassembly module (100).

Abstract: A radiation detector module includes a scintillator (62, 62?, 162, 262) arranged to receive penetrating radiation of a computed tomography apparatus (10). The scintillator produces optical radiation responsive to the penetrating radiation. A detector array (66, 66?, 166, 266) is arranged to convert the optical radiation into electric signals. Electronics (72, 72?, 172, 272) are arranged on a side of the detector array opposite from the scintillator in a path of the penetrating radiation. A radiation shield (86, 86?, 100, 100?, 100?, 186, 210, 210?, 286, 286?) is disposed between the detector array and the electronics to absorb the penetrating radiation that passes through the scintillator. The radiation shield includes openings (90, 90?) that communicate between the detector array and the electronics. Electrical feedthroughs (88, 88?, 102, 102?, 102?, 188, 212, 212?, 288, 288?) pass through the radiation shield openings and electrically connect the detector array and the electronics.

Abstract: A radiation detector (30) for a computed tomography scanner (12) includes a support structure (62). An alignment board (60) secures to the support structure (62) and includes photolithographically defined alignment openings (70) arranged to define a spatial focal point (34) relative to the alignment board (60). An anti-scatter element (32) is disposed on the support element (62) and includes one or more protrusions (86) which mate with the alignment openings (70) of the alignment board (60) to align the anti-scatter element (32) with the spatial focal point (34). A detector board (104) includes alignment structures (106) that align the detector board (104) with the anti-scatter element (32).

Abstract: A radiation detector (30) for a computed tomography scanner (12) includes a support structure (62). An alignment board (60) secures to the support structure (62) and includes photolithographically defined alignment openings (70) arranged to define a spatial focal point (34) relative to the alignment board (60). An anti-scatter element (32) is disposed on the support element (62) and includes one or more protrusions (86) which mate with the alignment openings (70) of the alignment board (60) to align the anti-scatter element (32) with the spatial focal point (34). A detector board (104) includes alignment structures (106) that align the detector board (104) with the anti-scatter element (32).

Abstract: X-rays from an x-ray tube (16) pass through an examination region (14) and are detected by a single or two-dimensional x-ray detector (20). The x-ray detector (20) includes an array (22) of photodiodes, CCD devices, or other opto-electrical transducer elements. A matching array (24) of transparent scintillator crystals, e.g., CdWO4, is supported on and optically coupled to the photoelectric transducer array. A layer (26) of a high efficiency scintillator with a good spectral match with the opto-electrical transducer array but with limited light transmissiveness is optically coupled to the transparent scintillator array. The layer (26) is preferably zinc selenide ZnSe (Te). Electrical signals from the transducer array are reconstructed (32) into an image representation and converted into a human-readable display (38). To reduce cross-talk, the zinc selenide layer is etched with pits (40), sliced into strips (26′), cut into rectangles (26″), or has channels (44) cut into it.

Abstract: An x-ray tube assembly (16) includes a vacuum envelope (52) and an x-ray permeable exit window (58). An anode (50) is positioned within the vacuum envelope (52) such that a near side is adjacent to the exit window (58) and a far side is opposite thereof. A cathode assembly (66) is also mounted within the vacuum envelope (52) which directs an electron beam (72) toward a focal spot or point (62) on the far side of the anode (50). The anode further includes a central cavity or indentation (70) which provides a location for mounting a set of radiation attenuating vanes (64) in addition to a shaped x-ray filter or compensator (68). Close placement of the vanes (64) and the filter (68) relative to the focal spot of the anode desirably reduce off focal radiation and allow beam shaping. An externally located collimator (18) further shapes the output x-ray beam.

Abstract: A CT scanner (10) includes a reconstruction processor (32) and a mosaic X-Radiation detector (20). The mosaic detector includes plural detector elements (22, 22, 23, 24, 25, 26) arranged in abutting relationship and configured for the desired imaging application. The detector elements include scintillating crystals (50) in optical communication with a back-illuminated photodiode array (52) or modified top-surface photodiode array (152, 252) for converting emitted light into electrical charge. The photodiode array is mounted on a carrier substrate (58) via bump (56) bonding. The carrier substrate provides a conductive path routing the photodiode array output through to contacts on the back side for connection to readout electronics (60). The carrier substrate and readout electronics are contained within the footprint defined by the photodiode array, allowing the detector elements to be abutted on any and all sides, thus permitting the mosaic detector to be tailored to any desired size and shape.

Abstract: An x-ray source (30, 80, 100) transmits a beam of x-rays through an examination region (E). A receiver (28, 82, 102), in an initial spatial orientation relative to the source (30, 80, 100), receives the beam and generates a view of image data indicative of the intensity of the beam received. A sensor, such as an accelerometer, detects motion in a selected portion of a mechanical structure (M) supporting the source (30, 80, 100) and the receiver (28, 82, 102). Upon detection of motion, the sensor generates a motion signal. In one embodiment, a first accelerometer (40, 90) is associated with the receiver (28, 82) and a second accelerometer (42, 88) is associated with the source (30, 80). A position calculator (58, 60) mathematically calculates a position of both the source and receiver based on the acceleration data generated by the accelerometers.

Abstract: An x-ray tube assembly (16) includes a vacuum envelope (52) and an x-ray permeable exit window (54). An anode (50) is positioned within the vacuum envelope (52) such that a near side is adjacent to the exit window (54) and a far side is opposite thereof. A cathode assembly (74) is also mounted within the vacuum envelope (52) which directs an electron beam (78) toward a focal spot or point (58) on the far side of the anode (50). The anode further includes a central cavity or indentation (70) which provides a location for mounting a set of radiation attenuating vanes (64) in addition to a shaped x-ray filter or compensator (68). Close placement of the vanes (64) and the filter (68) relative to the focal spot of the anode desirably reduce off focal radiation and allow beam shaping. An externally located collimator (18) further shapes the output x-ray beam.

Abstract: A CT scanner (10) includes a reconstruction processor (32) for reconstructing an image from digital signals from detector arrays (20). Each detector array includes scintillation crystals (22) arranged in an array for converting x-ray radiation into light. An array of back-illuminated photo diodes (24) is mounted beneath the scintillation crystal array for converting the light emitted from the scintillation crystals into electrical charge. The electrical charge from the back-illuminated photodiodes is transmitted via a path orthogonal to the detector array (20, 40) to signal processing circuitry (66). The back-illuminated photodiode has a backside (26) which is in optical communication with the crystal array (22) and which is optically transmissive to photons of light emanating from the crystal. The converted electrical charge leaves the photodiode via electrical connections (28) or bump bonds (62, 72) on the front side of the photodiode.

Abstract: An x-ray radiation stabilization system is provided including an x-ray tube (20) which emits x-ray radiation (22). The x-ray tube (20) has an anode (52), a cathode (50), and a vacuum envelope (54) which houses the anode (52) and the cathode (50). A high-voltage generator (40) is connected to the x-ray tube (20). It supplies a high-voltage electric potential between the cathode (50) and anode (52) such that an electron beam flows therebetween. The electron beam strikes the anode (52) producing the x-ray radiation (22). A reference radiation detector (60) samples a representative portion of the x-ray radiation (22) emitted by the x-ray tube (20) and generates an error signal in response to an intensity of the sampled x-ray radiation (22). A feedback circuit (80) is connected between the reference radiation detector (60) and the high-voltage generator (40).

Abstract: An x-ray source (30) transmits a beam of x-rays through an examination region (E). A detector (28), in an initial spatial orientation relative to the source, receives the beam and generates a view of image data indicative of the intensity of the beam received. A first accelerometer (40), capable of generating acceleration data in at least one dimension, is associated with the detector. A second accelerometer (42), capable of generating acceleration data in at least one dimension is associated with the source. A position calculator (60) mathematically calculates a position of both the source and detector based on the acceleration data generated by the accelerometers. An image reconstructor (62) receives the relative position data; electronically corrects for any misalignment or change in beam travel distance, and reconstructs the views into a volumetric image representation.

Abstract: A gantry (10) includes a large diameter bearing having an outer race (12) and an inner race (16) which surrounds an examination region. An x-ray tube (18) and collimator (52) are mounted to the inner race, as is a flat panel detector (20) and a mechanical mechanism (50) for moving the flat panel detector closer to and further from the examination region. A timing and control circuit (30) controls a motor (22) which indexes the inner race around the examination region, an x-ray power supply (32) which pulses the x-ray tube in a fluoroscopic mode at discrete positions around the examination region, and a read out circuit (34) which reads out a frame of data after each pulse of the x-ray tube. The read out frames of data are stored in a frame memory (36) and reconstructed by a reconstruction processor (38) into a volumetric image representation for storage in a volume image memory (40).

Abstract: A radiographic scanner (10) has a stationary gantry portion (12) defining a subject receiving region (16) and a rotating gantry portion (20) on which an imaging x-ray tube (22) is mounted. The rotating gantry portion (20) is rotatably mounted to the stationary gantry portion (12) for rotation about the subject receiving region (16). A slip ring assembly extending around the subject receiving region (16) connected with the stationary and rotating gantry portions, includes a scintillating optical fiber (44) mounted around the patient receiving region (16) to one of the rotating and stationary gantry portions. A communication x-ray tube (40) is mounted to the other gantry portion and directed such that radiation therefrom enters the scintillating optical fiber (44) from a lateral direction. The scintillating optical fiber (44) converts the incident x-rays (52) to light (58) and transmits the light (58) along its longitudinal axis.

Abstract: A toroidal x-ray tube (I) is supported and selectively positioned by a gantry (II). The x-ray tube includes a toroidal housing (A) in which a rotor (20) is rotatably mounted. One or more cathodes (C) are mounted on the rotor for generating an electron beam which strikes an anode (B) to generate a beam of x-rays which pass through a patient aperture (62) to strike a detector (60). The x-ray tube includes pre-collimators (70, 74) having slots (72, 76) for passage of the x-ray subsequent to generation thereof and prior to being collimated by the collimator (90). A ring collimator (90) collimates an x-rays formed into a fan shaped beam. The collimator (90) includes a first ring (92) and a second ring (94) which are concentric. The distance between the first and second rings (92, 94) is adjustable to adjust the slice thickness of the final image.

Abstract: A toroidal x-ray tube (I) is supported and selectively positioned by a gantry (II). The x-ray tube includes a toroidal housing (A) in which a rotor (30) is rotatably mounted. One or more cathodes (C) are mounted on the rotor for generating an electron beam which strikes an anode (B) to generate a beam of x-rays which passes through a window (20) and strikes an annular ring of detectors (160). A grid bias control circuit (100) selectively applies a continuously adjustable bias to a grid (36) for regulating the electron current, hence the intensity of the x-ray beam. A scintillating optical fiber (110) extends around the exterior of the window. The scintillation optical fiber includes fluorescent dopant (116) which convert a very small fraction of the x-rays into optical light which is transmitted along the fibers to an opto-electric transducer (118). The opto-electric transducer is connected with the grid bias control circuit.

Abstract: A source (A) of images, such as a CT scanner (10), a magnetic resonance imaging apparatus (12), and the like produces a plurality of basis images (I.sub.0, I.sub.1, I.sub.2, I.sub.3 . . . ). Two of the basis images are subtracted and divided (70, 72) by a number of interpolation increments (L.sub.1) to form a first differential image (I.sub..DELTA.1). The first and the third basis images are subtracted and divided (76, 78) by a number of available interpolation increments (L.sub.2) to form a second differential image (I.sub..DELTA.2). Four differential images are selectively combined and divided by a product of the first and second available increments (82, 84) to form a second order differential image (I.sup.2.sub..DELTA.12). An array of adders (D) selectively adds the first differential image to a currently displayed image stored in an image memory E each time a track ball (104) moves a cursor one increment in a horizontal position.

Abstract: A CT scanner (A) non-invasively examines a volumetric region of a subject and generates volumetric image data indicative thereof. An object memory (B) stores the data values corresponding to each voxel of the volume region. An affine transform algorithm (60) operates on the visible faces (24, 26, 28) of the volumetric region to translate the faces from object space to projections of the faces onto a viewing plane in image space. An operator control console (E) includes operator controls for selecting an angular orientation of a projection image of the volumetric region relative to a viewing plane, i.e. a plane of the video display (20). A cursor positioning trackball (90) inputs i- and j-coordinate locations in image space which are converted (92) into a cursor crosshair display (30) on the projection image (22). A depth dimension k between the viewing plane and the volumetric region in a viewing direction perpendicular to the viewing plane is determined (74).