Abstract: An imaging detector includes a radiation sensitive region having first and second opposing sides. One of the first or second sides senses impinging radiation. The detector further includes electronics located on the other of the first or second sides of the radiation sensitive region. The electronics includes a thermal controller that regulates a temperature of the imaging detector.

Abstract: A radiographic imaging apparatus (10) comprises a primary radiation source (14) which projects a beam of radiation into an examination region (16). A detector (18) converts detected radiation passing through the examination region (16) into electrical detector signals representative of the detected radiation. The detector (18) has at least one temporally changing characteristic such as an offset B(t) or gain A(t). A grid pulse means (64) turns the primary radiation source (14) ON and OFF at a rate between 1000 and 5000 pulses per second, such that at least the offset B(t) is re-measured between 1000 and 5000 times per second and corrected a plurality of times during generation of the detector signals. The gain A(t) is measured by pulsing a second pulsed source (86, 100, 138) of a constant intensity (XRef) with a second pulse means (88). The gain A(t) is re-measured and corrected a plurality of times per second during generation of the detector signals.

Abstract: A computed tomography method includes rotating an electron beam along an anode (104) disposed about an examination region (112) for a plurality of sampling intervals in which x-ray projections are sampled. The electron beam is swept during each sampling interval to generate a plurality of successive focal spots at different focal spot locations during each sampling interval, wherein the focal spots generated in a given sampling interval include a sub-set of the focal spots generated in a previous sampling interval. The x-ray projections radiated from each of the plurality of focal spots is sampled during each sampling interval. The resulting data is reconstructed to generate volumetric image data.

Abstract: An ionizing radiation detector module (22) includes a detector array (200), a memory (202), signal processing electronics (208), a communications interface (210), and a connector (212). The memory contains detector performance parameters (204) and detector correction algorithms (206). The signal processing electronics (208) uses the detector performance parameters (204) to correct signals from the detector array (200) in accordance with the detector correction algorithms (206).

Abstract: A medical imaging system includes an x-ray source (112) having a focal spot that emits radiation that traverses a examination region (108). The position of the focal spot along a longitudinal direction is a function of a temperature of one or more x-ray source components. The system further includes a detector (120) that detects the radiation and a collimator (116), disposed between the x-ray source (112) and the examination region (108), that collimates that radiation along the longitudinal direction. A focal spot position estimator (132) dynamically computes an estimated position of the focal spot along the longitudinal direction based on the temperature of one or more x-ray source components. A collimator positioner (128) positions the collimator (116) along the longitudinal direction based on the estimated focal spot position prior to performing a scan.

Abstract: A computed tomography method includes rotating an electron beam along an anode (104) disposed about an examination region (112) for a plurality of sampling intervals in which x-ray projections are sampled. The electron beam is swept during each sampling interval to generate a plurality of successive focal spots at different focal spot locations during each sampling interval, wherein the focal spots generated in a given sampling interval include a sub-set of the focal spots generated in a previous sampling interval. The x-ray projections radiated from each of the plurality of focal spots is sampled during each sampling interval. The resulting data is reconstructed to generate volumetric image data.

Abstract: A radiographic imaging apparatus (10) comprises a primary radiation source (14) which projects a beam of radiation into an examination region (16). A detector (18) converts detected radiation passing through the examination region (16) into electrical detector signals representative of the detected radiation. The detector (18) has at least one temporally changing characteristic such as an offset B(t) or gain A(t). A grid pulse means (64) turns the primary radiation source (14) ON and OFF at a rate between 1000 and 5000 pulses per second, such that at least the offset B(t) is re-measured between 1000 and 5000 times per second and corrected a plurality of times during generation of the detector signals. The gain A(t) is measured by pulsing a second pulsed source (86, 100, 138) of a constant intensity (XRef) with a second pulse means (88). The gain A(t) is re-measured and corrected a plurality of times per second during generation of the detector signals.

Abstract: A computed tomography system (100) includes an x-ray source (112) that rotates about an examination region (108) and translates along a longitudinal axis (120). The x-ray source (112) remains at a first location on the longitudinal axis (120) while rotating about the examination region (108), accelerates to a scanning speed and performs a fly-by scan of a region of interest (220) in which at least one hundred and eighty degrees plus a fan angle of data is acquired. At least one detector (124) detects x-rays radiated by the x-ray source (112) that traverses the examination region (108) and generates signals indicative thereof. A reconstructor (132) reconstructs the signals to generate volumetric image data.

Abstract: A radiation detector module (22) particularly well suited for use in computed tomography (CT) applications includes a scintillator (200), a photodetector array (202), and signal processing electronics (205). The photodetector array (202) includes a semiconductor substrate (208) having a plurality of photodetectors and metalization (210) fabricated on non-illuminated side of the substrate (208). The metalization routes electrical signals between the photodetectors and the signal processing electronics (205) and between the signal processing electronics (205) and an electrical connector (209).

Abstract: An ionizing radiation detector module (22) includes a detector array (200), a memory (202), signal processing electronics (208), a communications interface (210), and a connector (212). The memory contains detector performance parameters (204) and detector correction algorithms (206). The signal processing electronics (208) uses the detector performance parameters (204) to correct signals from the detector array (200) in accordance with the detector correction algorithms (206).

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 data measurement system (DMS) (30) for a computed tomography (CT) scanner (12) includes a plurality of connectorized detector sub-array modules (32). Each detector sub-array module (32) includes: a scintillator (40) that produces scintillation events responsive to irradiation by x-rays; a photodetector array (42) arranged to detect the scintillations; and two symmetrically arranged signal connectors (541, 542) that transmit the photodetector signals. Symmetrically mounted pipeline cards (60) mate with the signal connectors (54) of each side of groups of the detector sub-array modules (32) to receive the photodetector signals. A processor (64) communicating with the pipeline cards (60) receives the photodetector signals from the pipeline cards (60) and constructs a DMS output from the photodetector signals.

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 data measurement system (DMS) (30) for a computed tomography (CT) scanner (12) includes a plurality of connectorized detector sub-array modules (32). Each detector sub-array module (32) includes: a scintillator (40) that produces scintillation events responsive to irradiation by x-rays; a photodetector array (42) arranged to detect the scintillations; and two symmetrically arranged signal connectors (541, 542) that transmit the photodetector signals. Symmetrically mounted pipeline cards (60) mate with the signal connectors (54) of each side of groups of the detector sub-array modules (32) to receive the photodetector signals. A processor (64) communicating with the pipeline cards (60) receives the photodetector signals from the pipeline cards (60) and constructs a DMS output from the photodetector signals.

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 third generation CT scanner includes a rotating x-ray source (18) and a detector array (16). Each sampling of the detector array generates a source fan data line of data values that are converted (34) to attenuation values in a logarithmic domain and subject to preliminary corrections (36). Attenuation values from a plurality of adjoining source fans are converted (94) back to a non-logarithmic domain as intensity values. A corresponding deconvolution function (88) from a deconvolution function look-up table (90) corresponding to the detector whose intensity value is being corrected is deconvolved with a line (84) of the intensity values which spans a plurality of adjoining source fan data lines to remove the intensity attributable to off-focal radiation (30). The intensity data is converted (102) back into attenuation values in the logarithmic domain and reconstructed (106) into an image representation for display on a monitor (112).

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.