Abstract: When generating a 3D image of a subject or patient, a cone beam X-ray source (20a, 20b) is mounted to a rotatable gantry (14) opposite an offset flat panel X-ray detector (22a, 22b). A wedge-shaped attenuation filter (24a, 24b) of suitable material (e.g., aluminum or the like) is adjustably positioned in the cone beam to selectively attenuate the beam as a function of the shape, size, and density of a volume of interest (18) through which X-rays pass in order to maintain X-ray intensity or gain at a relatively constant level within a range of acceptable levels.

Abstract: When performing nuclear (e.g., SPECT or PET) and CT scans on a patient, a volume cone-beam CT scan is performed using a cone-beam CT X-ray source (20) and an offset flat panel X-ray detector (22). A field of view of the X-ray source overlaps a field of view of two nuclear detector heads (18), and the offset of the X-ray detector (22) minimizes interference with nuclear detector head movement about a rotatable gantry (16). Additionally, a locking mechanism (80) provides automatically locking of the X-ray detector (22) in each of a stowed and operation position, improving safety and CT image quality.

Abstract: A computed tomography acquisition geometry provides an increased field of view (218). A radiation source (202, 702) such as an x-ray source and a radiation detector (204, 704) are displaced from the imaging center (214). In one implementation, the central ray (216) of a radiation beam (212) is parallel to the plane of the detector (204, 704) at the detector midpoint (219, 719), but is displaced from the imaging center.

Abstract: When generating a 3D image of a subject or patient, a cone beam X-ray source (20a, 20b) is mounted to a rotatable gantry (14) opposite an offset flat panel X-ray detector (22a, 22b). A wedge-shaped attenuation filter (24a, 24b) of suitable material (e.g., aluminum or the like) is adjustably positioned in the cone beam to selectively attenuate the beam as a function of the shape, size, and density of a volume of interest (18) through which X-rays pass in order to maintain X-ray intensity or gain at a relatively constant level within a range of acceptable levels.

Abstract: When performing nuclear (e.g., SPECT or PET) and CT scans on a patient, a volume cone-beam CT scan is performed using a cone-beam CT X-ray source (20) and an offset flat panel X-ray detector (22). A field of view of the X-ray source overlaps a field of view of two nuclear detector heads (18), and the offset of the X-ray detector (22) minimizes interference with nuclear detector head movement about a rotatable gantry (16). Additionally, a locking mechanism (80) provides automatically locking of the X-ray detector (22) in each of a stowed and operation position, improving safety and CT image quality.

Abstract: A computed tomography acquisition geometry provides an increased field of view (218). A radiation source (202, 702) such as an x-ray source and a radiation detector (204, 704) are displaced from the imaging center (214). In one implementation, the central ray (216) of a radiation beam (212) is parallel to the plane of the detector (204, 704) at the detector midpoint (219, 719), but is displaced from the imaging center.

Abstract: A radiation detector package includes a radiation-sensitive solid-state element (10) having a first electrode (12) and a pixelated second electrode (14) disposed on opposite principal surfaces of the solid-state element. An electronics board (20) receives an electrical signal from the solid-state element responsive to radiation incident upon the radiation-sensitive solid-state element. A light-tight shield (40, 40?) shields at least the radiation-sensitive solid-state element from light exposure and compresses an insulating elastomer and metal element connector (30, 32) between the pixilated electrode (14) and contact pads (24) on the electronics board.

Abstract: A nuclear camera system includes a detector (12) for receiving radiation from a subject (14) in an exam region (16). The detector (12) includes a scintillation crystal (20) that converts radiation events into flashes of light. An array of sensors (22) is arranged to receive the light flashes from the scintillation crystal (20). Each of the photomultiplier sensors (22) generates a respective sensor output value in response to each received light flash. A processor (26) determines when each of the radiation events is detected. At least one of an initial position and an energy of each of the detected radiation events is determined in accordance with respective distances (d1 . . . d19) from a position of the detected event to the sensors (22). An image representation is generated from the initial positions and energies.

Abstract: A nuclear camera system includes a detector (12) for receiving radiation from a subject (14) in an exam region (16). The detector (12) includes a scintillation crystal (20) that converts radiation events into flashes of light. An array of sensors (22) is arranged to receive the light flashes from the scintillation crystal (20). Each of the photomultiplier sensors (22) generates a respective sensor output value in response to each received light flash. A processor (26) determines when each of the radiation events is detected. At least one of an initial position and an energy of each of the detected radiation events is determined in accordance with respective distances (d1 . . . d19) from a position of the detected event to the sensors (22). An image representation is generated from the initial positions and energies.

Abstract: A nuclear camera system includes a detector (12) for receiving radiation from a subject (14) in an exam region (16). The detector (12) includes a scintillation crystal (20) that converts radiation events into flashes of light. An array of sensors (22) is arranged to receive the light flashes from the scintillation crystal (20). Each of the photomultiplier sensors (22) generates a respective sensor output value in response to each received light flash. A processor (26) determines when each of the radiation events is detected. At least one of an initial position and an energy of each of the detected radiation events is determined in accordance with respective distances (d1 . . . d19) from a position of the detected event to the sensors (22). An image representation is generated from the initial positions and energies.

Abstract: A nuclear medicine imaging device includes a radiation camera (10) including a plurality of photo multiplier tubes (28). Each photo multiplier tube (28) is configured with an analog to digital converter (30) which converts a detected scintillation event (50) to sampled digital values. A storage device (66) is preloaded with an estimator function which can be derived from a calibration scintillation events. A processor (14) in communication with both the camera (10) and the storage device (66), detects an event and combines the digital values which are sampled together to arrive at a total area or energy of the scintillation event. Alternately, if a second pulse (52) is detected before the first scintillation event (50) has ended, the area combining (A1) of the first event is stopped and a pulse tail is estimated (A2) from the estimator functions stored. This estimated tail is then added to the combined data values taken until the time of pile-up.

Abstract: In positron emission imaging, coincident gamma ray pairs are acquired and processed to generate an image. Random gamma ray pairs in the acquired coincidence data degrade the quality of the resultant image. The coincident gamma ray pairs are re-paired to generate non-coincident gamma ray pairs. The non-coincident pairs are used to correct for randoms in the acquired coincidence data. Alternately, singles gamma rays may be detected and paired with non-coincident single gamma rays to generate non-coincident pairs. These pairs may be used to correct for randoms in the acquired coincidence data.

Abstract: A DMA control device (10) is connected with an n-bit address bus (12) by way of a bidirectional internal n-bit bus (14). The m most significant bits of signals received on the bidirectional bus (14) are reserved for carrying codes which identify or enable the DMA device to respond, to generate a load signal, to generate a count signal, and to generate an output signal. The remaining bits are reserved for address data. The load signal causes the remaining bit addresses to be loaded into counters (22) or registers (40). The count signal causes the counters (22) or a latched incrementor (44) to increment. The output signal controls three-state buffers (24, 42, 46) which cause the current address to be outputted on the bidirectional bus. In this manner, the DMA control device has only a single bus and in the embodiment of FIG. 2 replaces the counter array with a register array.

Abstract: Pixel values f(m,n) of an image are reduced (A) by predicting each pixel value based on preceding pixel values and retaining the deviation e(m,n) between the actual and predicted values. The reduced values e(m,n) are serially fed to an input latch (20). A latched sizer (22) determines the larger of the number of bits of the value in the input latch and the largest previously received reduced value e(m,n). A comparator (42) compares a number of bits from the latched sizer with the number of bits per field indicated by a look-up table (44). Each time another reduced value e(m,n) is received, a state counter (46) indexes one state, i.e., increases the number of bit fields per output word and, when necessary causes the look-up table (44) to reduce the number of bits per field.

Abstract: A network (A) carries large image blocks among medical diagnostic equipment (20, 22, 24), archive computer (26), and data handling and display stations (28, 30). Four kilobyte packets of 4 megabyte image blocks are moved from transmit buffers (38) to queuing buffers (42). The order in which packets from the queuing buffers are transmitted on the network medium (10) is determined by a combination of an assigned priority, duration in the buffer, and a statistical availability of the addressed receiving node. A data link (52) at the receiving node includes an elasticity buffer (122) which stores a small plurality bits, e.g. 5, to accommodate for variances in the clock speed of the transmitting and receiving nodes. A buffer table (60) monitors memory addresses at which preceding data packets corresponding to the same image are stored and provides address information to send each subsequently received packet.