The invention relates to a combined imaging detector for detection of gamma and x-ray quanta comprising an x-ray detector (31) for generating x-ray detection signals in response to detected x-ray quanta and a gamma detector (32) for generating gamma detection signals in response to detected gamma quanta. The x-ray detector (31) and the gamma detector (32) are arranged in a stacked configuration along a radiation-receiving direction (33). The gamma detector (32) comprises a gamma collimator plate (320) comprising a plurality of pinholes (321), and a gamma conversion layer (322, 324) for converting detected gamma quanta into gamma detection signals.

A PET detecting module may include a scintillator array configured to receive a radiation ray and generate optical signals in response to the received radiation ray. The scintillator array may have a plurality of rows of scintillators arranged in a first direction and a plurality of columns of scintillators arranged in a second direction. A first group of light guides may be arranged on a top surface of the scintillator array along the first direction. The light guide count of the first group of light guides may be less than the row count of the plurality of rows of scintillators. A second group of light guides may be arranged on a bottom surface of the scintillator array. The light guide count of the second group of light guides may be less than the column count of the plurality of columns of scintillators.

Provided is a PET detector for reducing the number of silicon photomultipliers in use, which is characterized in that: the detector comprises layers respectively formed by a scintillation crystal array unit and a silicon photomultiplier (4) array unit, the scintillation crystal array unit and the silicon photomultiplier (4) array unit are rectangular cross sections in plan view, and the scintillation crystal array unit and the silicon photomultiplier (4) array unit have the same area of the rectangular cross sections in plan view; the scintillation crystal array unit consists of a plurality of scintillation crystal strips (1) parallel to each other, free of gaps and attached to each other on sides, the scintillation crystal strips (1) are all cuboids with uniform length, width and height; the silicon photomultiplier (4) array unit is an array assembly, which is formed by M silicon photomultiplier (4) arrays and has the rectangular cross section in plan view.

A scalable medical imaging detector arrangement is provided having interchangeable sensor tiles with fixed outer dimensions for a fixed or universal mechanical, electrical, and cooling interface. Different sensor tile types with different performance grades and production costs care configured with a common interface for coupling to the medical imaging device, while the rest of the imaging system can remain unchanged.

A PET detector and method thereof are provided. The PET detector may include: a crystal array including a plurality of crystal elements arranged in an array and light-splitting structures set on surfaces of the plurality of crystal elements, the light-splitting structures jointly define a light output surface of the crystal array; a semiconductor sensor array, which is set in opposite to the light output surface of the crystal array and is suitable to receive photons from the light output surface, the semiconductor sensor array comprises a plurality of semiconductor sensors arranged in an array.

Described here are tileable block detectors for use in nuclear medicine applications, such as in positron emission tomography (PET] systems and positron emission mammography (PEM] systems. The block detectors described here are four-side tileable such that seamless arrays of block detectors can be constructed for use in PET or PEM systems. When so arrayed, the block detectors allow for a full-size seamless detector that achieves full coverage of an object (e.g., a gently immobilized breast), improves data collection, and enables high-resolution imaging with a significantly lower radiation dose than with other currently available PEM systems.

A PET detecting module may include a scintillator array configured to receive a radiation ray and generate optical signals in response to the received radiation ray. The scintillator array may have a plurality of rows of scintillators arranged in a first direction and a plurality of columns of scintillators arranged in a second direction. A first group of light guides may be arranged on a top surface of the scintillator array along the first direction. The light guide count of the first group of light guides may be less than the row count of the plurality of rows of scintillators. A second group of light guides may be arranged on a bottom surface of the scintillator array. The light guide count of the second group of light guides may be less than the column count of the plurality of columns of scintillators.

The invention relates to a gamma radiation detector that provides compensation for the parallax effect. The gamma radiation detector includes a plurality of scintillator elements, a planar optical detector array, and a pinhole collimator that includes a pinhole aperture. Each scintillator element has a gamma radiation receiving face and an opposing scintillation light output face. The gamma radiation receiving face of each scintillator element faces the pinhole aperture for generating scintillation light in response to gamma radiation received from the pinhole aperture. The scintillator elements are arranged in groups. Each group has a group axis that is aligned with the pinhole aperture and is perpendicular to the radiation receiving face of each scintillator in that group. The scintillation light output faces of each of the scintillator elements are in optical communication with the planar optical detector array.

A method and apparatus are provided for positron emission imaging to correct a recorded energy of a detected gamma ray, when the gamma ray is scattered during detection. When scattering occurs, the energy of a single gamma ray can be distributed across multiple detector elementsa multi-channel detection. Nonlinearities in the detection process and charge/light sharing among adjacent channels can result in the summed energies from the multiple crystals of a multi-channel detection deviating from the energy that would be measured in single-channel detection absent scattering. This deviation is corrected by applying one or more correction factors (e.g., multiplicative or additive) that shifts the summed energies of multi-channel detections to agree with a known predefined energy (e.g., 511 keV). The correction factors can be stored in a look-up-table that is segmented to accommodate variations in the multi-channel energy shift based on the level of energy sharing.

The present invention provides a detector and an emission tomography device including the detector. The detector comprises: a scintillation crystal array comprising a plurality of scintillation crystals; and a photo sensor array, coupled to an end surface of the scintillation crystal array and comprising multiple photo sensors. At least one of the multiple photo sensors is coupled to a plurality of the scintillation crystals respectively. Surfaces of the plurality of the scintillation crystals not coupled to the photo sensor array are each provided with a light-reflecting layer, and a light-transmitting window is disposed in the light-reflecting layer on a surface among the surfaces adjacent to a scintillation crystal coupled to an adjacent photo sensor. The detector has DOI decoding capability. No mutual interference occurs during DOI decoding, and decoding is more accurate. Moreover, with the number of photo sensor arrays being the same, the decoding capability for the scintillation crystals is significantly improved. With the number of photo sensor arrays being the same, the size of the photo sensor array and the number of channels of a readout circuit of the photo sensors of the present invention can be reduced by three-quarters to eight-ninths.