Recent Advancements in Radio Nuclear ImagingByBindu Yadav USN:1RV13LBI032ndSemester,BMSP&IRVCEPresentation onR. V.COLLEGE OF ENGINEERING*BMSP&I,RVCE

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Emission Tomographyprojection*BMSP&I,RVCE

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SPECTSingle Photon Emission Computed Tomography*BMSP&I,RVCE

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PETPositron Emission TomographyWhat do we want to detect in PET?

2 photons of 511 keV in coincidence, coming in a straight line from the same annihilation*BMSP&I,RVCE

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How it works

A short lived radioactive tracer isotope, is injected in to the living subject (usually in to blood circulation) . The tracer is chemically incorporated in to a biologically active molecule.

There is a waiting period while the active molecule becomes concentrated in tissues of interest.

As the radioisotope undergoes positron emission decay (also known as positive beta decay), it emits a positron, an antiparticle of the electron with opposite charge.

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After traveling up to a few millimeters the positron encounter an electron.

The encounter annihilates them both, producing a pair of (gamma) photon moving in opposite directions.

These are detected when they reach scintillator in the scanning device creating a burst of light which is detected by photomultiplier tubes.

The technicians can then create an image of the parts of your brain, for example which are overactive.*BMSP&I,RVCE

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Limitations of PET scanTime-consuming.

The resolution of structures of the body with nuclear medicine may not be as clear as with other imaging techniques, such as CT or MRI.

PET scanning can give false results if chemical balances within the body are not normal.

Because the radioactive substance decays quickly and is effective for only a short period of time, it is important for the patient to be on time for the appointment and to receive the radioactive material at the scheduled time.

A person who is very obese may not fit into the opening of a conventional PET/CT unit.

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TOF measurementsThe two 511 keV gamma rays emitted when a positron and electron mutually annihilate each other arrive at PET detectors at different times if they have not been generated at the exact center of two detectors.The annihilation positions of gamma ray pairs (distances from the geometric center to annihilation positions) can be determined using d = c (t1-t2)/ 2.The timing resolutions of available clinical PET detectors are of the order of 600 picoseconds which equates to a positional uncertainty of ~ 9 cm (Conventional PET systems it is 27cm).

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TOF information can be used to reduce background noise in back-projection based reconstruction images. This is because the back-projected events can be confined within areas around estimated annihilation positions

SNR enhancements during TOF PET reconstruction, PET images of better quality.Better lesion-detection performanceReduce scan times Reduce the radiopharmaceutical doses administered.

Fig. (4). Background noise propagation in (A) TOF reconstruction and (B) conventional non-TOF reconstruction.*BMSP&I,RVCE

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Fig (5) : Benefit of TOF in FDG PET studies of heavy patients. Better lesion detection performance can be obtained in TOF PET relatively to non-TOF PET*BMSP&I,RVCE

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Resolution Degrading Factors in PETPositron range(the distance traveled by a positron before it undergoes annihilation) and displacement errors, which are caused by the non-collinear annihilation of gamma ray photons.when light signals from scintillation crystal arrays are not read individually, but rather multiplexed using the light sharing or resistive charge division methods to reduce the number of readout channels.At the peripheral region of the FOV in ring type PET scanners, parallax error makes the spatial resolutions worse and non-uniform.Furthermore, the use of longer crystals to enhance sensitivity increases parallax errors, which affect a large proportion of FOVs in small ring PET cameras. Therefore, the spatial resolution of PET is limited and not uniform across the FOV.

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Resolution Recovery ModelingIterative statistical reconstruction is now a standard PET reconstruction method for clinical PET data, and spatial resolution can be enhanced by incorporating a resolution degradation model in the iterative reconstruction procedure.In conventional method, the system matrix is determined using the intersection lengths or areas between lines of response (LOR) and voxels.Here we use analytical calculations, Monte Carlo simulation, and point source measurements.Enhance spatial resolution and reduce statistical noise (at the expense of computation time).Improved lesion to background contrast and lesion detection performance.

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RESOLUTION RECOVERY RECONSTRUCTIONFig. (6). Iterative statistical reconstruction incorporating a resolution degradation modelFig. (7). Effects of resolution recovery reconstruction: reconstructed images of the point sources (A) without and (B) with resolution recovery modeling.*BMSP&I,RVCE

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DEPTH-OF-INTERACTION MEASUREMENTSDOI measurements provide another useful means ofimproving detection sensitivity and spatial resolution,especially those of small bore ring-type PET cameras (e.g.,brain, breast, and animal PET) because the parallax error can be corrected using DOI information.The most direct method involves reading the light output from each crystal block using thin semiconductor photosensorsDOI information can be also obtained using light emitted from the rear surfaces of stacked, optically coupled, multiple crystal blocksThe relative offset method - DOI information using stacked crystal layers and a single-end readout.One of two crystal layers is shifted by half a crystal pitch in both the horizontal and vertical directions. Visible light photons emitted from the crystals are read by position sensitive PMT or by multi-channel PMT, which provide the centroid of the light distribution, which is also shifted between the crystal layers.

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Fig. (8). Discrete DOI encoding methods. (A) Light readout from each individual crystal layer. (B) Pulse shape discrimination. (C) Relative offset method. (D) Combination of (B) and (C).

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Application specific PETsLarge-Bore PET

The Philips GEMINI TF Big bore PET/CT has a bore size of 85cm which is larger than the standard TOF PET/CT scanner(70 cm bore size) offered by Philips.

Fig. (9). Large Bore PET (Courtesy of Philips Medical).*BMSP&I,RVCE

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Breast dedicated PETEnhanced anatomical and functional informationIt provides clearer images of suspected lesions by overcoming the tissue superposition problem of mammography Higher spatial resolutions and tumor contrast than conventional CT

Fig. (10). Positron emission mammography (PEM).*BMSP&I,RVCE

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Brain-dedicated PETBrain-dedicated PET scanners are design to provide detailed anatomical information and improved quantitative accuracy. They have smaller ring diameters (40~50 cm) and crystal sizes (2~4 mm) than whole-body PET/CT scanners to increase spatial resolution, sensitivity, and image quality. For example, the Siemens HRRT PET scanner has a ring diameter of 46.9 cm and an axial field-of-view (FOV) of 25.2 cm.

Figure (11) Brain-dedicated PET scanners. (A) Siemens HRRT. (B) Philips G-PET. (C) Photo Detection Systems NeuroPET. (D) NIRS jPETD4*BMSP&I,RVCE

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Fig. (12). Transformable PET (A) Brain mode. (B) Whole-body mode.Transformable PETTwo ring geometries, namely, a diameter of 83 cm and an axial FOV of 13 cm for whole-body imaging or a diameter of 54 cm and axial FOV of 21 cm for brain/breast imaging

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ScintillatorsThallium-activated sodium iodide, NaI(Tl) is still the most commonly used scintillator for SPECT applications. In most clinical SPECT systems, the NaI(Tl) detector is one large crystal as described above. However, NaI(Tl) can be pixelated and several small-field-of-view devices, including small-animal SPECT systemsThallium-activated cesium iodide competes well with NaI(Tl) in terms of efficiency and has a significantly longer scintillation time, However, the scintillation light from CsI(Tl) has a longer wavelength that is not as well matched to PMTs This approach is used by Digirad in their Cardius product.cesium iodide, CsI(Na), is similar to CsI(Tl), but with an emission that is better matched to PMTs. It is currently being used as the detector in the LinoView small-animal SPECT systemLaBr3 has a very high light output with a correspondinglyimproved energy resolution .In addition, LaBr3 is a very fast detector and potentially very useful for approaches such as Compton g-cameras, where high counting rates are likely to be encountered. It is also being used as the detector in are recently reported small-animal SPECT system

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Photon TransducersPMTs have been used to convert scintillation light into electronic signals .PMTs have very large electronic gains with relatively low noise. However, their quantum conversion efficiency is low (;20%), leading to a significant loss of signal that affects both energy resolution and intrinsic spatial resolution. It is difficult to maintain the long-term stability of PMTs because they are susceptible to environmental influences such as temperature, humidity and magnetic fields and their properties change as they age. In addition, they are bulky as well as expensive.Improved replacements for conventional PMTs: positionsensitivePMT (PSPMT), the avalanche photodiode (APD) and the silicon photomultiplier

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Position Sensitive PMTsConventional PMTs have no intrinsic localizationthat is, there is no way of knowing what portion of the photocathode interacts with the scintillations. PSPMTs are essentially a matrix of PMT elements that combine high gain capabilities with the ability to localize events to several millimeters.PSPMTs have been primarily used in small-field-of-view devices with individual, pixilated detectors .In addition to their position sensitivity and high gain, PSPMTs are more compact and more efficiently sample the detector surface than individual PMTs. Limitations: However, there are wide gain variations across the field, which makes this device better suited for pixelated detectors. PSPMTs are also influenced by magnetic fields and other environmental factors

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APDsAPDs are solid-state photon converters .They can be conceptualized as a light-sensitive diode with a very high reverse bias . Incoming photons liberate charge carriers(electrons and holes) in the cathode, which are accelerated through the diode depletion zone with enough energy to create additional charge carriers. The final electronic signal is proportional to the initial number of light photons that were detected. APDs have several advantages over PMTs. They are more rugged, very compact, and largely immune to environmental factors such as magnetic fields. They operate at a lower voltage than PMTs and have a much higher quantum conversion efficiency. However, the maximum gain of currently available APDs is about 250 (compared with 106 for PMTs). A large gain is desirable for separating the scintillation signal from noise, which is an important factor in energy resolution.APDs are especially well suited for scintillators that emit light with longer wavelengths, such as CsI(Tl), and are best suited for pixelated detectors.

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silicon photomultiplierPhotodiodes are analogous to ionization chambers except that the charge carriers can be liberated with significantly lower energy. The APD operates much like a proportional counter as the output signal is proportional to the initial number of photons. Within this proportional region, only the electrons have enough energy to cause an avalanche. If the bias voltage is increased to this configuration, this proportionality is lost because the holes also contribute avalanches, leading to discharge. This is referred to as the Geiger region.As in gas ionization chambers, the Geiger region is associated with a total discharge of the device regardless of how many charge carriers were initially detected. A device with a single Geiger region would not have much value as a photon detector, but in the silicon PMT, a large array of Geiger regions is closely packed together. Current devices are available with a 24 X 24 array of Geiger photodiodes packaged into a 1 X 1 mm area.Silicon photomultipliers have all of the advantages of the APDs (compactness, low sensitivity to environmental factors, high quantum efficiency) with the added advantage of very high gains at relatively low bias levels.

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FIGURE (13) . Photon transducers for converting scintillationsinto an electronic pulse. (A) PMT. (B) Position-sensitive PMT. (C)Avalanche photodiode. (D) Silicon photomultiplier.*BMSP&I,RVCE

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REFERENCESMark T. Madsen, Recent Advances in SPECT Imaging, The Journal of Nuclear Medicine, 2012.Jae Sung Lee , Technical Advances In Current PET And Hybrid Imaging Systems, The Open Nuclear Medicine Journal, 2010.Craig S. Levin, PhDa,*, Habib Zaidi, PhD, PDb ,Current Trends in Preclinical PET System Design, Elsevier Inc, 2007 .

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