A handheld radiation image detecting system and an operation method thereof are provided. The handheld radiation image detecting system includes a handheld device including a radiation emitter and a first transceiver and a sensing device including a radiation image sensor and a second transceiver. The first transceiver is coupled to the radiation emitter and used for generating a first wave with directionality. The second transceiver is used for receiving the first wave and for generating a second wave with directionality, and the first transceiver is used for receiving the second wave.

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.

The invention provides novel Compton camera detector designs and systems for enhanced radiographic imaging with integrated detector systems which incorporate Compton and nuclear medicine imaging, PET imaging and x-ray CT imaging capabilities. Compton camera detector designs employ one or more layers of detector modules comprised of edge-on or face-on detectors or a combination of edge-on and face-on detectors which may employ gas, scintillator, semiconductor, low temperature (such as Ge and superconductor) and structured detectors. Detectors may implement tracking capabilities and may operate in a non-coincidence or coincidence detection mode.

A radiation detector for combined detection of low-energy radiation quanta and high-energy radiation quanta has a multi-layered structure. A rear scintillator layer (5) is configured to emit a burst of scintillation photons responsive to a high-energy radiation quantum being absorbed by the rear scintillator layer (5). A rear photosensor layer (6) attached to a back side of the rear scintillator layer (5) is configured to detect scintillation photons generated in the rear scintillator layer (5). A front scintillator layer (3) arranged in front of the rear scintillator layer (5) opposite the rear photosensor layer (6) is configured to emit a burst of scintillation photons responsive to a low-energy radiation quantumbeing absorbed by the front scintillator layer (3). A front photosensor layer (2) attached to a front side of the front scintillator layer (3) opposite the rear scintillator layer (5) is configured to detect scintillation photons generated in the front scintillator layer (3). The high-energy radiation quantum is a gamma ray and the low-energy radiation quantum is an X-ray.

Detector systems for enhanced radiographic imaging incorporate Compton and PET imaging capabilities. The detector designs employ one or more layers of detector modules comprised of edge-on or face-on detectors, or a combination of edge-on and face-on detectors, which may employ structured detectors. The detectors implement tracking capabilities and operate in a non-coincidence or coincidence detection mode.

Uptake of hypoxia-sensitive PET tracers is dependent on tissue transport properties, specifically, distribution volume. Variability in tissue transport properties reduces the sensitivity of static PET imaging to hypoxia. When tissue transport (v.sub.d) effects are substantial, correlations between the two methods of determining hypoxic fractions are greatly reducedthat is, trapping rates k.sub.3 are only modestly correlated with tumour-to-blood ratio (TBR). In other words, the usefulness of dynamic- and static-PET based hypoxia surrogates, trapping rate k.sub.3 and TBR, in determining hypoxic fractions is reduced in regions where diffusive equilibrium is achieved slowly. A process is provided for quantifying hypoxic fractions using a novel biomarker for hypoxia, hypoxia-sensitive tracer binding rate k.sub.b, based on PET imaging data. The same formalism can be applied to model the kinetics of non-binding CT and MT contrast agents, giving histopathological information about the imaged tissue.

A method for generating transmission information in a time-of-flight positron emission tomography (PET) scanner having a patient tunnel and a plurality of PET detector rings. The PET scanner uses continuous bed motion to move a patient bed and patient through the patient tunnel. The patient receives a positron-emitting radioisotope dose prior to undergoing a PET scan. The method includes storing a positron-emitting radioisotope in a radiation shielded container. The method also includes moving the radioisotope into a stationary vessel located adjacent to the PET detectors and within a field of view of the PET scanner at substantially the same time that the patient receives the radioisotope dose to form a stationary transmission source wherein transmission information is generated while the bed undergoes continuous bed motion. Further, the method includes withdrawing the radioisotope from the vessel when the PET scan is complete and storing the radioisotope in the container.

Detector module designs for radiographic imaging include first and second layers of scintillator rods or pixel arrays oriented in first and second directions. The first and second directions are transversely oriented to define a light sharing region between the first and second layers. Encoding features may be disposed in, on or between the first and second layers, and configured to modulate propagation of optical signals therealong or therebetween.

Detector systems for enhanced radiographic imaging incorporate one or more Compton and nuclear medicine imaging, PET imaging, and x-ray CT imaging capabilities. The detector designs employ one or more layers of detector modules comprising edge-on or face-on detectors, or a combination of edge-on and face-on detectors, which can employ gas, scintillator, semiconductor, low temperature (such as Ge and superconductor) or structured detectors. The detectors implement tracking capabilities, and operate in non-coincidence or coincidence detection modes.

A radiation detector (16) having a first detector layer (24) and a second detector layer (26) encircles an examination region (14). Detectors of the first layer include scintillators (72) and light detectors (74), such as avalanche photodiodes. The detectors of the second detector layer (26) include scintillators (62) and optical detectors (64). The scintillators (72) of the first layer have a smaller cross-section than the scintillators (62) of the second layers. A group, e.g., nine, of the first layer scintillators (72) overlay each second group scintillator (62). In a CT mode, detectors of the first layer detect transmission radiation to generate a CT image with a relatively high resolution and the detectors of the second layer detect PET or SPECT radiation to generate nuclear data for reconstruction into a lower resolution emission image. Because the detectors of the first and second layers are aligned, the transmission and emission images are inherently aligned.