Nuclear imaging_ by hcj

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									Nuclear imaging Instrumentation part
Prepared by : Dr. Ali Saad, College of applied medical Science Dept of Biomed. Technology.

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Nuclear imaging,
ِAll procedures involving the detection of and image formation from the emissions of radiopharmaceuticals introduced into patients for diagnostic purposes. Nuclear scanning started with the use of rectilinear scanners, systems which had a single detector with pencil-like detection sensitivity. By moving this detector over a body region of a patient in a meandering fashion, and printing the local gamma- ray intensity, images of the distribution of radioactivity could be obtained. While such systems are still in use, they have been superseded by two-dimensional detection devices called gamma cameras. The most widely used gamma cameras are the socalled Anger cameras, in which a series of phototubes detects the light emissions of a large single crystal, covering the field of view of the camera. Alternatively, individual crystals arranged in a two-dimensional array are used as detector systems. Such detector systems may well be the detectors of the future as with modern technology, large arrays of such detectors can be produced. If the detectors are arranged in ringdetector banks, cross-sectional images can be obtained, and if several rings adjacent to each other are used, data acquisition is three-dimensional. While such SPECT imaging systems have been devised, their cost and poor flexibility have resulted in single- or multiple-head gamma cameras which rotate around the patient, thereby acquiring the projections necessary for reconstruction of axial slices. PET imaging, the most recent nuclear imaging method introduced into clinical practice is also based on ring detector systems, but recently, manufacturers also have started to fit dualhead camera systems with the coincident detection circuitry necessary for PET imaging.

Radiopharmaceutical, Substance consisting of a molecule to which a radionuclide is bound. There are many types of radiopharmaceuticals such as chelates, blood pool agents, substances which are excreted by the liver, and molecules subject to very selective binding or uptake mechanisms such as iodine (thyroid) or markers of receptor systems in the brain and other organs. These can be imaged with nuclear imaging methods because these methods can detect nano- to picomolar concentrations of substances. Radionuclide, It is an isotope which is radioactive and thus undergoes radioactive decay.

Isotope, Isotopes are families of atomic elements which have a fixed atomic number (number of protons) and a variable number of neutrons and thus of nucleons. Isotopes of the same atom share common chemical characteristics as they contain an identical number of

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to counterbalance the fixed nuclear protonic charge. As an example, the element carbon C has four isotopes, C-11, C-12, C-13 and C-14. C-12 and C-13 are stable, while C-11 and C-14 are beta emitters (see beta decay). C-14 decays by emitting an electron and C-11 decays by emitting a positron. C-13 has a nuclear spin and is used in MR spectroscopy.
electrons

Radioactive decay Also called radioactive disintegration, radioactive transformation or simply decay or disintegration. Fundamental physical process by which some atomic nuclei disintegrate (see isotope). There are several types of radioactive decay, classified as alpha decay, beta decay and gamma decay. Another type of decay is the socalled electron capture EC . Many radioactive isotopes, particularly heavy ones such as uranium, disintegrate by a series of radioactive decays (radioactive series) until they have been transformed into stable atoms. Radioactive decay obeys the radioactive decay law: an exponential decay. If we call the initial number of radioactive isotopes N0 and the number remaining after a time t, Nt, the decay law is given by
Nt = N0 exp(-t) where  is the radioactive decay constant. The half life t is defined as the time during which the number of radioactive nuclei decays to half its initial value: = Nt/N0 =exp(- t), hence t =  ln2 The half life of isotopes extends from the extremely short times observed only in experimental atomic reactions to times in the minute-to-hour range used in modern nuclear imaging to the extremely long half lives, encountered in isotopes such as uranium-238, which has a half life of several billion years. The way in which radioactive isotopes decay are usually represented in so called decay schemes or disintegration schemes

Beta decay type of radioactive decay in which a nucleus ejects a beta particle, either an electron or a positron. In a beta (-) decay a neutron gets converted into a proton and an electron. Hence, the atomic number of the nucleus increases by one, the number of nucleons stays constant and the electron leaves the nucleus as a beta (-) particle. In a beta (+) or positron decay a proton is converted into a neutron and a positron or beta (+) particle, which leaves the nucleus. The atomic number of the nucleus thus decreases by one, while the number of nucleons stays constant. Beta (-) decays are favored in small radioactive nuclei when atoms have an excess of neutrons relative to protons, while beta (+) decays are favored when the "mother" nuclide has an excess of protons relative to neutrons. 3

Alpha decay
process during which a radioactive nucleus also called an alpha emitter emits a helium He nucleus (alpha particle, alpha ray) consisting of two protons and two neutrons. A "mother" nucleus decaying with an alpha decay will therefore end up as a "daughter" nucleus with an atomic number that is by two smaller than the "mother" nucleus, and a nucleonic number that is by four smaller. Alpha particles strongly interact with matter. Thus they can easily be shielded and are almost quantitatively absorbed by even a sheet of paper. However, their effect on biological tissue is very strong as they have a high linear energy transfer LET .
radioactive decay

Gamma decay
in which a nucleus emits a high-energy photon or gamma ray. In such a decay, a nucleus that has undergone another type of radioactive decay remains in an excited or metastable state for a prolonged time eventually relaxing back to the ground state by emitting the gamma ray. A typical example is the decay of technetium 99m to technetium-99.
radioactive decay

Gamma Camera
imaging device used in nuclear scanning. By far the most widely used gamma camera was invented by H. Anger in the 1960s and thus is also frequently called the Anger camera. The Anger camera principle is also used in one type of PET camera (see PET imaging), either using six crystals arranged in a hexagon configuration or retrofitting a dual head gamma camera with an electronic coincidence detection circuitry. With this arrangement, the lead collimator is not needed because of the electronic collimation provided by the PET detection process. Digital gamma cameras are currently under development which use arrays of small crystals digitally interfaced to a computer. It is too early to say how they will compare with the Anger camera principle in clinical practice. Image reconstruction, The process of generating an image from the raw data, or set of unprocessed measurements, made by the imaging system. In general, there is a well defined mathematical relationship between the distribution of physical properties (e.g. density, acoustic impedance, magnetization) in an object and the measurements made by the imaging system. Image reconstruction is the process which inverts this mathematical process to generate an image from the set of measurements. In an ideal X ray imaging system, the X-ray intensity is exponentially attenuated by the object density along the X-ray path. The logarithm of the ratio of the incident intensity to the transmitted intensity is equal to the line integral or raysum of the 4

distribution of linear attenuation coefficient within the object along the path. An image of the object density distribution can be created using a projection reconstruction algorithm. In an ideal computed tomography CT system image reconstruction would be accomplished by a simple projection reconstruction algorithm, where the projection measurements are simply filtered by an appropriate convolution kernel and then back-projected . In an actual CT scanner, there are a variety of imperfections which must be corrected to minimize artifacts in the reconstructed images. Offset correction and gain correction are used to eliminate errors between detector channels. Beam hardening correction is used to minimize the effects of a broad X-ray energy spectrum. Other compensation techniques, such as ring removal in the third generation CT scanner are also used to minimize scanner geometry specific artifacts.

Anger camera, imaging device used in nuclear imaging and invented by H. Anger in the 1960s.The principle by which it functions is at the basis of almost all cameras used in current day nuclear imaging.

Anger camera, Fig. 1 Photons are selected by a collimator and produce light flashes which are detected by the photomultipliers. See text.

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Anger camera, Fig. 2 Light flash producing different responses in the detecting photomultipliers.

An Anger camera functions as follows (Fig.1). It consists of a collimator, placed between the detector surface and the patient. This is made out of lead and serves to suppress gamma rays which deviate substantially from the vertical and thus acts as a type of "lens". The detector is a single crystal of NaI (sodium iodide), which produces light flashes of multiple photons when an impinging gamma ray interacts with the crystal (scintillation camera). The bursts of light flashes are detected by an array of photomultiplier tubes which are optically coupled to the surface of the crystal (Fig.2). The output signal from the photomultipliers is currents, which are proportional to the energy of the gamma ray. Depending on the position of the event, the phototubes are variably activated. Hence, the entire system response yields positional information which is relatively accurate. With actual systems the intrinsic resolution (see full width half maximum FWHM ) of two radiation sources placed immediately on the crystal surface without the collimator is in the order of 1 mm. The positional information is recorded using an analogue output onto film or a digital image is stored in a computer coupled to the camera. Since the total amount of energy deposited by a single gamma ray and thus the net current generated by all photomultiplier tubes is close to constant, a deviation of the net current below a preset threshold signifies that the detected ray was degraded by Compton scattering. Use of a discrimination circuitry therefore permits suppression of Compton-scattered photons.

Collimator
device made of a highly absorbing material such as lead which selects X- or gamma-rays along a particular direction. In an X ray tube and computed tomography CT collimators are placed so as to ensure that only a limited region is irradiated, but in CT they are also placed in front of the detector bank to eliminate stray radiation. In nuclear imaging, they serve to suppress scatter but also to select a ray orientation (see Anger camera (I), Fig. 1).

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The simplest collimators contain parallel holes (Fig.1), but various other arrangements are

in use.
Collimator, Fig. 1

Parallel hole collimator, typically made of lead, with "honeycomb"-like structure. Note that the thickness of lead shielding between adjacent holes is minimal every 60.

Photomultiplier tube
(PMT) an extremely sensitive photocell used to convert light signals of a few hundred photons into a usable current pulse without significantly increasing noise. A PMT consists of two major elements; a photocathode coupled to an electron multiplier structure contained within a glass envelope. The photocathode (Fig.1) comprises a photosensitive layer that converts as many of the incident light photons as possible into low-energy electrons. The number of photoelectrons produced will be comparable to the number of incident light photons and thus, the charge on the photoelectrons will be too small to provide a detectable electrical signal. The electron multiplier section consists of an arrangement of dynodes (Fig. 1) that serves both as an efficient collection geometry for the photoelectrons and a nearly ideal amplifier, that greatly increases the number of electrons. After amplification, a typical scintillation pulse will give rise to 107-1010 electrons, sufficient to generate a charge signal that can be collected at the anode. PMTs in general, perform a highly linear charge amplification producing an output pulse which, over a wide range, is proportional to the number of original photoelectrons. Much of the timing information of the original light pulse is also retained. When illuminated by a very short duration light pulse, most tubes will produce an electron pulse with a time width of a few nanoseconds after a delay of only 20-50 ns.

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Photomultiplier tube, Fig. 1:Schematic representation of a photomultiplier tube (PMT) Compton scattering (Arthur H. Compton, 18921962, American physicist, Nobel laureate 1927). During Compton scattering, a photon impinges on an electron in matter, and in this process transfers part of its energy to it. The excited electron is termed a Compton electron and is ejected or moved into an excited atomic state, while due to the law of conservation of energy the photon energy is reduced.

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SPECT imaging
Single photon emission (computed) tomography (SPECT or SPET): tomographic nuclear imaging technique producing cross-sectional images from gamma ray emitting radiopharmaceuticals (single photon emitters or positron emitters). SPECT data are acquired according to the original concept used in tomographic imaging: multiple views of the body part to be imaged are acquired by rotating the Anger camera detector head(s) around a craniocaudal axis. Using backprojection, cross-sectional images are then computed with the axial field of view FOV (the area which an imaging modality is set to visualize. Its linear dimensions are given in cm or mm. Typical FOVs vary from a few cm up to the size of the human body in whole-body imaging.) determined by the axial field of view of the gamma camera. SPECT cameras are either standard gamma cameras which can rotate around the patient's axis or consist of two or even three camera heads to shorten acquisition time (Fig.1). A few systems use ring detector arrays similar to computed tomography CT but this type of system is not in widespread use.

SPECT camera Spect imaging, Fig. 1 Schematic

representation of triple head SPECT camera

Data acquisition is over at least half a circle (180) (used by some for heart imaging), but mostly over a full circle. Data reconstruction has to take into account the fact that the emitted rays are also attenuated within the patient, i. e. photons emanating from deep inside the patient are considerably attenuated by surrounding tissues. While in CT absorption is the essence of the imaging process, in SPECT attenuation degrades the images. Thus, data of the head reconstructed without attenuation correction may show substantial artificial enhancement of the peripheral brain structures relative to the deep ones. The simplest way to deal with this problem is to filter the data before reconstruction (Fig.2).

Attenuation correction
correction applied for the scatter and absorption of photons emitted from a radioactive tracer distributed within the body. Unless a correction is applied for absorption and scattering, tracer concentration will appear to be artifically reduced in deep structures of 9

the body, i.e two regions of equal concentration, one located on the surface and one located at depth, will appear to have different concentrations of tracer. Deep structures will be spatially distorted along directions of lower attenuation, and thus measurements of lesion size are also unreliable. To correct for attenuation, an exponential factor is required which corrects for the probability of attenuation along the path of the photon from the point of emission to the face of the detector. This factor is of the form exp{x}, where  is the linear attenuation coefficient (which is closely related to the Hounsfield unit in computed tomography CT ), and x is the distance from the point of emission to the detector. In SPECT imaging, this depth-dependent correction is complicated by the fact that photons originating from different depths are detected at the same position on the detector. A single correction factor for each detector element is therefore not sufficient. A number of approximate, iterative-based algorithms are available to correct for attenuation in SPECT. The current trend is to measure approximate attenuation factors using a scanning line source of photons. In PET imaging, attenuation correction is, in principle, simpler because two factors are required, one to correct for the attenuation of each photon from the coincident pair. The appropriate factor is therefore exp{-x} exp{-x2} = exp{-X} where X=x + x2, is the total path length for the two photons and is therefore independent of the actual point of photon emission. The correction factors exp{-X} can be measured exactly using an external source of 511 keV photons, typically germanium-68. This measurement, which is in addition to the regular emission scan, is called a transmission scan. The ability to correct exactly for photon self-absorption is the reason that PET is a fully quantitative imaging technique.

Spect imaging, Fig. 2 :Head SPECT image (technetium-99m HMPAO), showing a 10

normal brain perfusion A more elegant but elaborate method used in triple head cameras is to introduce a gamma-ray line source between two camera heads (Fig. 1), which are detected by the opposing camera head after being partly absorbed by the patient. This camera head then yields transmission data while the other two collect emission data. Note that the camera collecting transmission data has to be fitted with a converging collimator to admit the appropriate gamma rays (Fig. 1).

Backprojection,
a step in a projection reconstruction algorithm in which the projection measurements are added to all picture elements in an image along the lines of the original projection paths (Fig.1). Backprojection is the second step in the standard computed tomography CT reconstruction algorithm (see filtered backprojection). When the projection measurements are backprojected without any prefiltering, the resulting image is significantly blurred. If only a few view angles are backprojected, then a star-like pattern emanating from dense points in the image is created (Fig. 1, right). When a large number of views are backprojected, originally dense points appear as dense points in the image, but are significantly blurred because their density affects the reconstructed values at all other pixel locations. The blurring kernel is the radially symmetrical function, 1/r, where r is the distance from the dense point. Because backprojection is very computationally intensive, particularly in direct fan-beam reconstruction, special purpose processors called backprojectors have been designed and are used in most CT systems.

Backprojection, Fig. 1 Four views are shown resulting in four projection measurements of a single dense point (left). The four measurements are backprojected across the entire field of view, giving a star-like image of the point (right). When the number of views is increased, the star-like image converges to a radial blur of the point.

Projection reconstruction the process of generating images from a set of projection measurements. In computed tomography CT , the projections refer to the set of line integrals or raysums through the object. There are a variety of mathematical expressions that can be used to accomplish image reconstruction from a set of projections. These include iterative reconstruction techniques and filtered backprojection. Most reconstruction algorithms are based on the 11

assumption that a large number of projection measurements are made such that every point in the object is included in X-ray paths from all angles. Backprojection of all of these measurements would then result in a blurred image. The blurring can be removed by a filtering of the reconstructed image or a prefiltering of the sets of projection measurements.

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PET Imaging
(or positron emission tomography) is a tomographic nuclear imaging procedure, which uses positrons as radiolabels and positron - electron annihilation reaction-induced gamma rays to locate the radiolabels. Current clinical indications for PET imaging include the location of epileptic foci, the grading of brain tumours, the assessment of cerebral and cardiac perfusion, the identification of hibernating myocardium and tumour staging (Fig.1). PET is an important research tool for the assessment of cerebral function, metabolism and receptor ligand systems. These applications as well as the use of PET for the detection of foci of acute inflammation are being intensively investigated at all major PET centres. The principal radiopharmaceuticals used are fluorodeoxyglucose FDG , O 15 water and N 13 ammonia.

PET imaging, Fig. 1Coronal

whole-body FDG-PET scan (thickness 7 mm) of a patient showing normal FDG accumulation in the brain and the renal excretory system as well as a pathological focal accumulation medial to the right tibia due to metastasis of malignant melanoma. The PET principle is as follows. A low dose of a radiopharmaceutical labelled with a positron emitter such as C-11, N-13, O-15 or F-18 is injected into the patient, who is scanned by the tomographic system. Scanning consists of either a dynamic series or a

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static image obtained after an interval during which the radiopharmaceutical enters the biochemical process of interest. The scanner detects the spatial and temporal distribution of the radiolabel by detecting gamma rays during the so-called emission scan. The events leading to this detection are (Fig.2):

Pet imaging, Fig. 2 PET principle showing annihilation reaction between positron and electron, production of two gamma rays and detection in coincidence detection system.

1. the positron is emitted by a beta decay, 2. it is slowed down to small speeds which are necessary for 3. the annihilation reaction between the positron and a shell electron of a neighbouring atoms to occur. The distance the positron travels (mean free path) depends on the energetics of the beta decay but is typically one or a few millimeters. 4. The annihilation reaction produces two 511 keV gamma rays which travel in almost exactly opposite directions (this is due to the conservation of energy and momentum laws). 5. The two gamma rays are detected by a coincidence counting detection system (see below). 6. After proper filtering the collected raw data sinograms are reconstructed into a crosssectional image. In a coincidence counting system, the detection of one gamma ray by one detector results in the opening of an electronic time "window", during which detection of a gamma ray in

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another detector results in the counting of a coincidence event (Fig.2). It is obvious from the geometry of the gamma ray emission that the radioactive decay leading to the detection of such a coincidence event must have occurred along a straight line connecting the two detectors. Since the probability of absorption of the two gamma rays is independent of the position of the event along that imaginary line, PET is an inherently quantitative imaging method allowing the measurement of regional concentrations of the radiopharmaceutical injected after proper calibration. To achieve quantitative detection, several problems have to be overcome. First, a correction has to be introduced for elastic scattering (scatter correction), while Compton scattering can be corrected for to some extent by using an energy discriminator on the gamma ray detectors. Second, a correction for random coincidences has to be introduced. The frequency of such random coincidences (two gamma rays detected which come from different decay events) depends on the imaging system. They are more frequent the higher the concentration of the activity in the patient and are more frequent during three-dimensional data acquisition (see below). Third, an absorption correction matrix has to be acquired to correct for tissue absorption. This is done by acquiring a transmission PET scan prior to or after the injection of the radiopharmaceutical. Typically, a germanium radiation source is rotated around the patient and the attenuation of this radiation by the patient recorded and reconstructed into a transmission scan, which is then used to correct the data of the emission scan. Alternatively, the correction is made by assuming that the attenuation values for the imaged pixels is homogeneous, an assumption which is only reasonable for brain scans. The reconstructed data are transaxial slices of a few millimetres in thickness. Depending on the application they are reconstructed into coronal (Fig. 1), sagittal or oblique sections for clinical reading. Dynamic data acquisition is performed when the data are to be quantified. In this technique, scans are acquired over times as short as 30 seconds. The most important technique for analysis of dynamic data is input deconvolution. Using this mathematical technique, quantitative information on tissue perfusion and other parameters can be extracted. Usually data are acquired in a two-dimensional (2D) acquisition mode with a PET scanner consisting of several detector rings (Fig.3).

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Pet imaging, Fig. 3Two- and three-dimensional data acquisition. With septa in place only coincidences of directly opposing or adjacent rings are counted, while in three-dimensional imaging the septa are retracted and coincidences among all opposing detectors are possible

In 2D imaging the detectors are separated by lead septa which permit coincidences to occur between opposing detectors (red lines, Fig. 3 left) or adjacent detectors (green lines, Fig. 3 left). In the three-dimensional (3D) mode the septa are retracted and coincidences between all possible detector banks are admitted (Fig. 3, right). While data acquisition in 3D mode has become standard, whole-body imaging is currently done in 2D or 3D mode and cardiac imaging in 2D mode. The major problems with 3D data acquisition are the large number of random coincidences counted and the long data reconstruction time, but the latter problem is alleviated by increasing computer performance. Annihilation reaction reaction in which a positron generated by a beta decay and slowed to "thermal speeds", react with an electron to completely disappear (annihilate) and form two gamma rays (photons) which radiate from the point of annihilation in almost exactly opposite directions (conservation of momentum) (Fig.1). This conversion of matter into energy E is in accordance with the famous Einstein equation E = mc2 (mass: m; velocity of light c), which requires that the energy of each gamma ray is 511 keV . The reaction is used in PET imaging.

Annihilation reaction, Fig. 1 : An electron and a positron meet, annihilate and form two gamma rays.

Fluorodeoxyglucose (FDG), Glucose analogue and the compound predominantly used in PET imaging . F-18-2fluoro-2-deoxyglucose (FDG) is a glucose analogue which competes with glucose in the hexokinase reaction, in which glucose is phosphorylated at the 6 position to form glucose-6-phosphate. In contrast to glucose, FDG-6-phosphate does not undergo glycolysis, and does not enter the fructose-pentose shunt or glycogen synthesis pathway. Thus, cellular FDG uptake reflects the overall rate of transmembranous exchange of glucose. The dephosphorylation of FDG is also slow and when modeling the behaviour of FDG (see compartment modelling ), this reaction is negligible as long as the time from FDG injection is less than 60 minutes. Owing to the simple reaction pattern, the behaviour of FDG in brain is amenable to quantitative compartment modelling .

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FDG-PET is useful in imaging the following tumors: it is taken up by bronchial carcinoma, colorectal cancer, esophageal and gastric cancers, testicular, breast and other gynecological cancers, ENT cancers, lymphomas, melanomas and sarcomas. It shows little uptake in most prostate cancers, and cancers of the renal collecting system, where an additional difficulty is distinguishing lesions from the normal urinary FDG excretion. It is probably not very useful in plasmacytoma/multiple myeloma. Additionally, FDG-PET may be an excellent marker for acute and chronic inflammation, as it is avidly taken up by activated granulocytes and macrophages. This cells, when engaged in fighting inflammation, undergo a so called metabolic burst, where glucose and thereby FDG uptake is increased manyfold. Coincidence counting Method of counting employing a coincidence circuit so that an event is recorded only if events are detected in two sensing devices simultaneously. Such counting methods may be used to reduce background noise if a radioisotope emits more than one detectable radiation event in coincidence. The effective background from sources other than this radioisotope can be greatly reduced through the use of coincidence counting. The requirement for a coincidence between two detectors eliminates background counts that occur in only one detector at a time. Coincidence detection of gamma rays is used in PET imaging. When the two gamma rays originating from a single positron-electron annihilation reaction are detected within a specific time window, the annihilation is described as a coincident event. Coincident events identify the presence of the positron-emitting labelled compound. The ability to distinguish true coincidence counts from scattered and random coincidence counts is the basis of the signal in PET imaging. Sinogram, Image representation of raw data obtained when projection-reconstruction imaging is used. The sinogram is a three-dimensional representation of the signal measured at a given angle in the imaging plane at varying distances r along the detector array (Fig.1).

Sinogram, Fig. 1

Schematic representation of raw data acquisition process and representation in a sinogram. The signal intensity measured is a histogram I() for each angle . In X computed -ray tomography CT , where the imaged object results in absorption, I() is low in the

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centre and high when no object is in the beam (dashed line), while in emission tomography (SPECT imaging, PET imaging), the intensity measured is highest in the region where the most emitting radioisotopes are present (solid line). Fig.2 represents four PET sinograms (vertical axis , horizontal axis r) of sections through the body where the narrowest portion of the sinogram in the middle represents the lateral projections (narrowest body projection) and the widest portions at the top and bottom the AP projections (widest body projection). From these data, the cross-sectional images are reconstructed.

Sinogram, Fig. 2 Four sinograms of four transaxial PET sections through a patient' s body.

In PET imaging the term is used somewhat loosely to identify a projection image of I () for a given angle, typically AP, oblique and lateral (0, 30, 60, and 90) with the vertical image axis being the axial direction of the patient. Thus, a projection image of the imaged body region is generated (Fig.3). The sinogram corresponds to the scanogram used in computed tomography

Sinogram, Fig. 3Projection images of a body PET scan data set.

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