Nuclear Emission Imaging
PET & SPECT
By
Abdullah A. Wahbi
Student in Arabian Developing Institute Biomedical Engineering 2010
1 Introduction to
NUCLEAR
EMISSION
1.1 Introduction Introduction
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1.2 Radioisotopes 1.2.1 Atomic Structure Structure and Radioactive Radioactive Decay Decay 1.2.1.1 Atomic Structure Tightly bound protons and neutrons form the nucleus of an atom. Both types of particles within the nucleus receive the generic name of nucleons. According to the Bohr’s atomic model, negatively charged electrons are located around the nucleus in orbits or shells of different binding energy levels. Protons have a positive electrical charge of the same magnitude as that of electrons and mass approximately 1840 times of that of the electron. Neutrons have the same mass as protons, but no electrical charge. Since nucleons include most of the atomic mass, the total number of nucleons in an atom is named the atomic mass number (A). The total number of protons, which is equal to the total number of orbital electrons, defines the unique position of the atom in the periodic table of elements and is named the atomic number (Z).
1.2.1.2 Isotopes Those atoms having the same number of protons but a different total number of nucleons are named isotopes. The notation to represent isotopes is X , where X is the chemical symbol of the element. Since there is a unique relationship between the atomic number and the symbol of the element, only the atomic mass number superscript is used to differentiate the isotopes of the same element.
1.2.1.3 Radioactive Decay Nucleons are packed together in the atomic nucleus by the strong short-range nuclear forces among them. However, in certain combinations of protons and neutrons, electrostatic repulsion forces among protons predominate over nuclear forces and the nucleus becomes unstable. Unstable nuclei spontaneously transform to more stable combinations of neutrons and protons by releasing or absorbing energy in the form of subatomic particles or electromagnetic radiation of high frequency. This This process has been given the name of radioactive transformation or radioactive decay. Upon a radioactive transformation, the resultant nucleus may be stable or it may still be unstable and subsequently transform again. The atomic nucleus prior to radioactive decay is named the parent, while the nucleus after transformation is named the daughter. daughter.
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1.2.1.4 Radioisotopes Those unstable isotopes of an element suffering radioactive decay are named radioisotopes. All nuclei with Z > 82 are radioactive, with the exception exception of 209Bi. Some others others lighter natural nuclei (Z < 82) are also radioactive.
1.2.2 Activity Activity and Half-Life 1.2.2.1 Activity The rate of radioactive decay, or disintegrations per unit of time of a radioactive sample, is named the activity of the sample. The classical unit of activity is the curie (Ci) defined as 10 3.7 x 10 disintegrations per second (dps). The unit of activity of the System International (SI) is the Becquerel (Bq) defined as one disintegration per second. Activities in emission imaging are usually expressed in millicuries (mCi) and megabecquerels (MBq) or microcuries (μCi) and kilobecquerels (kBq). Table 10.1 shows the conversion between both units.
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1.2.2.2 Half-Life The decay of a radioisotope is commonly expressed by its half-life. That is the time necessary for one half of the radioactive atoms of a sample to decay. The relationship between half-life (T1/2) and the decay constant can be easily demonstrated:
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1.2.3 The Energy of Nuclear Radiations 1.2.3.1 Electronvolt The energy of particles or electromagnetic radiation involved in a radioactive decay process is expressed in a special unit named electronvolt (eV). It represents the energy acquired by an electron through an electric field of 1 V of potential difference. One electronvolt is equivalent to 1.6 x 10 –19Joules (J). Since this is a very small unit, the multiples kiloelectronvolt (keV = 103eV) and megaelectronvolt (MeV = 106eV) are commonly used.
1.2.3.2 Gamma Rays Electromagnetic energy emitted by the atomic nucleus or annihilation processes is named gamma radiation. The energy of gamma rays can be from fr om 50 keV to higher than 3 MeV, MeV, but for emission imaging, only gamma rays r ays in the range from f rom 60 to 511 keV are used. Frequency 19 -1 . of gamma rays is usually higher than 3 x 10 s
1.2.3.3 Characteristic X-Rays Electromagnetic energy released from the transition of orbital electrons from an outer- to an inner-shell is named characteristic x-rays. Energy of characteristic x-rays is from 124 eV upward and usually overlaps the energy range of gamma rays. The frequency of characteristic x-rays is higher than 3 x 10 17 s-1
1.2.4 Types of Radioactive Radioac tive Decay
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1.2.4.1 Alpha Decay This is commonly limited to radioisotopes of high atomic number (Z > 82). The alpha particle consists of two protons and two neutrons and is equivalent to the nucleus of a Helium atom. Although the alpha particle is usually emitted with an approximate kinetic energy of 4 MeV, MeV, it is stopped by a few centimeters of air or by a few microns of tissue. Consequently, Consequently, alpha emitters are not used for emission imaging but have been proposed and investigated for internal radiotherapy of cancer tumors. After an alpha decay, decay, the atomic number of the daughter is two units units less than that of the parent and the atomic mass is four units less than that of the parent:
1.2.4.2 Beta Decay This transformation consists in the nuclear emission of a particle of mass and magnitude of the charge equal to that of the electron and a neutral particle without mass. The particle may be negatively charged, in which case it is named negatron or beta minus (β–). During β– disintegration the atomic number of the daughter is increased by one unit:
where ύ represents the neutral, without mass particle named antineutrino.
1.2.4.3 Positron Emission In some radioisotopes , the emitted particle may be positively charged. In this case the particle is named positron (β+) and the atomic number of the daughter is reduced by one unit:
the accompanying uncharged, without mass particle υ is named neutrino. Both particles are stopped by a few millimeters of tissue and cannot be used directly for emission imaging. However, the positron has an important characteristic that makes it useful for emission imaging. After the positron loses most of its kinetic energy in some millimeters millimeter s of tissue, the particle interacts with an electron at rest and the two particles undergo annihilation. In the annihilation process, the rest mass of each particle is converted into electromagnetic radiation in the form of two gamma rays of 511 keV, keV, each emitted in opposite directions (Figure 10.2). The simultaneous detection of these two gamma rays is the foundation of PET imaging.
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FIGURE 10.2 Annihilation of a positron and an electron, and emission of two gamma rays of 511 keV each in opposite directions.
Negatrons and positrons, the particles emitted during beta decay, are characterized by a maximum energy, energy, but most particles are emitted with energies lower than the maximum. The difference in energy between the maximum and the specific energy of each particle is carried away by the neutrino or antineutrino particles. Consequently, beta decay yields a continuous energy spectrum of negatron or positron particles.
1.2.4.4 Electron Capture This is a decay deca y process in which an orbital electron is absorbed by an unstable nucleus. As a consequence of this process, the daughter nucleus reduces its atomic number in one unit, such as in the positron decay. The loss of an orbital electron produces a rearrangement of the orbital electrons to fill the vacancy. This rearrangement is usually accompanied by the e mission of characteristic x-rays, with enough energy in some cases to travel out of the body and be used for emission imaging.
1.2.4.5 Isomeric Transition This decay is produced when energetic excited nuclei transforms to a more stable state by emitting the excess of energy as gamma rays. Nuclear excited states usually occur after alpha, beta or electron capture decays. In isomeric transitions there is no change in the atomic or mass number. To To differentiate the excited from the stable state, the letter le tter m, meaning metastable, is added to the atomic mass of the parent:
An excited nucleus is named metastable when its half-life is higher than 10 –6 sec. Gamma rays emitted in isomeric transitions have one or several specific energies, but not a continuous spectrum of gamma energies. In the above example, 99mTc has a half-life half- life of 6 h 99m and is produced by the beta decay of 99 Mo. The isomeric transition of Tc releases a gamma ray of 140 keV . 6
2 NUCLEAR MEDICINE INSTRUMENTATION
Single Photon Emission Computerized Tomography
(SPECT)
2.1 Introduction Introduction
Single-photon emission computed tomography (SPECT) is a medical imaging modality that combines conventional nuclear medicine (NM) imaging techniques and CT methods. Different from x-ray CT, CT, SPECT uses radioactive-labeled pharmaceuticals, i.e., radio pharmaceuticals, pharmaceuticals, that distribute distribute in different internal tissues tissues or or organs organs instead of of an external x-ray source. The spatial and uptake distributions of the radiopharmaceuticals depend on the biokinetic properties of the pharmaceuticals and the normal or abnormal state of the patient. The gamma photons emitted from the radioactive source are detected by radiation detectors similar to those used in conventional nuclear medicine. The CT method requires projection (or planar) image data to be acquired from different views around the patient. These projection data are subsequently reconstructed using image reconstruction methods that generate cross-sectional images of the internally distributed radiopharmaceuticals. The SPECT images provide much improved contrast and detailed information about the radiopharmaceutical distribution as compared with the planar images obtained from conventional nuclear medicine methods.
2.2 Basic Principles of SPECT Single-photon emission computed tomography (SPECT) is a medical imaging technique that is based on the conventional nuclear medicine imaging technique and tomographic reconstruction methods . General review of the basic principles, instrumentation, and reconstruction technique for SPECT can be found in a few review articles [Barrett, 1986; Jaszczak et al., 1980; Jaszczak and Coleman, 1985a; Jaszczak and Tsui, 1994].
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The conventional nuclear medicine imaging process. Gamma-ray photons emitted from the internally distributed radioactivity may experience photoelectric (a) or scatter (b) interactions. Photons that are not traveling in the direction within the acceptance analog of the collimator (c) will be intercepted by the lead collimator. Photons that experience no interaction and travel within the acceptance angle of the collimator will be detected (d).
2.2.1.1 Collimator a collimator is a device that filters a stream of rays so that only those traveling parallel to a specified direction are allowed through. Collimators are used in neutron, X-ray, and gammaray optics because it is not yet possible to focus radiation with such short wavelengths into an image through the use of lenses as is routine with electromagnetic radiation at optical or nearoptical wavelengths. Collimators are also used with radiation detectors in nuclear power stations for monitoring sources of radioactivit y.
How a Söller collimator filters a stream of rays. Top: Top: without a collimator. Bottom: with a collimator
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2.2.1.2 The Multi-Hole Collimator A multi-hole collimator, made of a high atomic number substance, is attached to the external surface of the detector head . The collimator consists of an array of thousands of holes separated by walls (septa) made out of lead and tungsten. The shape of holes varies in different collimators, but the most common designs are circular and hexagonal shapes. Collimator septa prevent photons from
2.2.1.3 Collimator Types
Types of collimators according to the direction of holes’ angles and number of holes.
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2.2.1.4 Scintillation crystal ( Scintillator ) A scintillator is material which exhibits the property of luminescencewhen excited by ionizing radiation. Luminescent materials, when struck by an incoming particle, absorb its energy and scintillate, i.e. reemit the absorbed energy in the form of a small flash of light, typically in the visible range. (Throughout this article, the word “particle” will be used to mean “ionizing radiation” and can refer to either charged particulate radiation such as electrons and heavy charged particles, or to uncharged radiation such as photons and neutrons, provided that the y have enough energy to induce ionization.) If the reemission occurs promptly, i.e. within the ≈10−8s required for an atomic transition, the process is called fluorescence. Sometimes, the excited state is metastable, so the relaxation back out of the excited state is delayed (necessitating anywhere from a few microseconds to hours depending on the material): the process then corresponds to either one of two phenomena, depending on the type of transition and hence the wavelength of the emitted optical photon: delayed fluorescence or phosphorescence (also called after-glow).
Scintillation crystal surrounded by various scintillation detector assemblies
2.2.1.5 Photomultiplier Photomultiplier tubes (photomultipliers or PMTs for short), members of the class of vacuum tubes, and more specifically phototubes, are extremely sensitive detectors of light in the ultraviolet, visible, and near-infrared ranges of the electromagnetic spectrum. These detectors multiply the current produced by incident light by as much as 100 million times (i.e., 160 dB), in multiple dynode stages, enabling (for example) individual photons to be detected when the incident flux of light is very low. The PMT “multiplies” photons photons because it has a quartz entrance window which is coated to release electrons when it absorbs a light photon and there is a voltage drop; the number of electrons released is proportional to the amount of light that hits the coating. The electrons are guided through a hole and caused to hit the first dynode, which is coated with a special substance to allow it to release electrons when it is hit by an electron. There are a series of dynodes each with a voltage that pulls the electrons from the last dynodes toward it. The surface coating not only releases electrons but also multiplies the electron shower. In a cascade through 10 to 12 dynodes, there is a multiplication of approximately 106 , 11
so that pulses of a few electrons become currents of the order of 10–12 amps. The PMTs PMTs must be protected from other influences, such as stray radioactivity or strong magnetic fields, which might cause extraneous electron formation or curves in the electron path. Without the voltage drop from one dynode to the next, there is no cascade of electrons and no counting.
Schematic drawing of a photomultiplier tube (PMT). Each of the dynodes and the anode is connected to a separate pin in the tube socket. The inside of the tube is evacuated of all gas. Dynodes are typically copper with a special oxidized coating for electron multiplication.
Photomultiplier
Dynodes inside a photomultiplier tube 12
2.3 The SPECT Imaging Process The imaging process of SPECT can be simply depicted as in Fig. 64.6. Gamma-ray Ga mma-ray photons emitted from the internal distributed radiopharmaceutical penetrate through the patient’s body and are detected by a single or a set of collimated radiation detectors. The emitted photons experience interactions with the intervening tissues through basic interactions of radiation with matter [Evans, 1955]. The photoelectric effect absorbs all the energy of the photons and stops their emergence from the patient’s body. body. The other major inter interaction action is Compton interaction, which transfers part of the photon energy to free electrons. The original photon is scattered into a new direction with reduced energy that is dependent on the scatter angle. Photons that escape from the patient’s body include those that have not experienced any interactions and those which have experienced Compton scattering. For the primary photons from the commonly used radionuclides in SPECT, SPECT, e.g., 140-keV of TC-99m and ~70-keV of TI-201, the probability of pair production is zero.
Most of the radiation detectors used in current SPECT systems are based on a single or multiple NaI(TI) scintillation detectors. The most significant development in nuclear medicine is the scintillation camera (or Anger camera) that is based on a large-area (typically 40 cm in diameter) NaI(TI) crystal [Anger, 1958, 1964]. An array of photomultiplier tubes (PMTs) (PMTs) is placed at the back of the scintillation crystal. When a photon hits and interacts with the crystal, the scintillation generated will be detected by the array of PMTs. An electronic circuitry evaluates the relative signals from the PMTs and determines the location of interaction of the incident photon in the scintillation crystal. In addition, the scintillation cameras have built-in energy discrimination electronic circuitry with finite energy resolution that provides selection of the photons that have not been scattered or been scattered within a small scattered angle. The scintillation cameras are commonly used in commercial SPECT systems. In SPECT, projection data are acquired from different views around the patient. Similar to x-ray CT, image processing and reconstruction methods are used to obtain transaxial or cross-sectional images from the multiple projection data. These methods consist of preprocessing and calibration procedures before further processing, mathematical algorithms for reconstruction from projections, and compensation methods for image degradation due to photon attenuation, scatter, and detector response.
2.4 SPECT Instrumentation Instrumentation A typical SPECT system consists of a single or multiple units of radiation detectors arranged in a specific geometric configuration and a mechanism for moving the radiation detector(s) or specially designed collimators to acquire data from different projection views. In general, SPECT instrumentation can be divided into three general categories: (1) arrays of multiple scintillation detectors, (2) one or more scintillation cameras, and (3) hybrid scintillation detectors combining the first two approaches. In addition, special collimator designs have been proposed for SPECT for specific purposes and clinica l applications. The following is a brief review of these SPECT systems and special collimators .
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2.5 Multidetector SPECT System The first fully functional SPECT imaging acquisition system was designed and constructed by Kuhl and Edwards [Kuhl and Edwards, 1963, 1964, 1968] in the 1960s, well before the conception of x-ray CT. CT. The MARK IV brain SPECT system consisted of four linear arrays of eight discrete NaI(TI) scintillation detectors assembled in a square arrangement. Projection data were obtained by rotating the square detector de tector array around the patient’s head. Although Although images from the pioneer MARK IV SPECT system were unimpressive without the use of proper reconstruction methods that were developed in later years, the multidetector design has been the theme of several other SPECT systems that were developed. An example is the Gammatom-1 developed by Cho et al. a l. [1982]. The design concept also was used in a dynamic SPECT system [Stokely et al., 1980] and commercial c ommercial multidetector SPECT systems marketed by Medimatic, A/S (Tomomatic-32). (Tomomatic-32). Recentl y, the system design was extended to a multislice SPECT system with the Tomomatic-896, Tomomatic-896, consisting of 8 layers of 96 scintillation detectors. Also, the system allows both body and brain SPECT imaging by varying the aperture size.
Examples of multidetector-based SPECT systems. (a) The MARK IV system consists of four arrays of eight individual NaI(TI) detectors arranged in a square configuration. (b) The Headtome-II system consists of a circular ring of detectors. A set of collimator vanes that swings in front of the discrete detector is used to collect projection data from different views. (c) A unique Cleon brain SPECT system consists of 12 detectors that scan both radially and tangentially.
Variations of the multiple-detectors multiple-detec tors arrangement have been proposed for SPECT system designs. shows the Headtome-II system by Shimadzu Corporation [Hirose et al., 1982], which consists of a stationary array of scintillation detectors arranged in a circular ring. Projection data are obtained by a set of collimator vanes that swings in front of the discrete detectors. A unique unique Cleon brain SPECT system , originally developed by Union Carbide Corporation in the 1970s, consists of 12 detectors that scan both radially and tangentially [Stoddart and Stoddart, 1979]. Images from the original system were unimpressive due to inadequate sampling, poor axial resolution, and a reconstruction algorithm that did not take full advantage of the unique system design and data dat a acquisition strategy strateg y. A much much improved version of the system with a new reconstruction method [Moore et al., 1984] is currently marketed by Strichman Corporation. The advantages of multidetector SPECT systems are their high sensitivity per image slice and high counting rate capability resulting from the array of multidetectors fully surrounding the patient. However, disadvantages of multidetector SPECT systems include their ability to provide only one or a few non-contiguous crosssectional image slices. 14
Also, these systems are relatively more expensive compared with camera-based SPECT systems described in the next subsection. With the advance of multicamera SPECT systems, the disadvantages of multidetector SPECT systems outweigh their advantages. As a result, they are less often found in nuclear medicine clinics.
2.6 Camera-Based SPECT Systems The most popular SPECT systems are based on single or multiple scintillation cameras mounted on a rotating gantry gantr y. The successful design was developed almost simultaneously by three separate groups [Budinger and Gullberg, 1977; Jaszczak et al., 1977; Keyes et al., 1977]. In 1981, General Electric Medical Systems offered the first commercial SPECT system based on a single rotating camera and brought SPECT to clinical use. Today Today,, there are over 10 manufacturers (e.g., ADAC, Elscint, General Electric, Hitachi, Picker, Siemens, Sopha, Toshiba, Trionix) offering an array of commercial SPECT systems in the marketplace. An advantage of camera-based SPECT systems is their use of off-the-shelf scintillation cameras that have been widely used in conventional nuclear medicine. These systems usually can be used in both conventional planar and SPECT imaging. Also, Also, camera-based SPECT systems allow truly three-dimensional (3D) imaging by providing a large set of contiguous transaxial images that cover the entire organ of interest. They are easily adaptable for SPECT imaging of the brain or body by simply changing the radius of rotation of the camera. A disadvantage of a camera-based SPECT system is its r elatively low counting rate capability. capability. The dead time of a typical state-of-the-art scintillation camera gives rise to a loss of 20% of its true counts at about 80K counts per second. A few special high-count-rate systems give the same count rate loss at about 150K counts per second. sec ond. For SPECT systems using a single scintillation camera, the sensitivity per image slice is relative low compared with a typical multidetector SPECT system.
Examples of camera-based SPECT systems. (a) Single-camera system. (b) Dual-camera system with the two cameras placed at opposing sides of patient during rotation. (c) Dual-camera system with the two cameras placed at right angles. (d) Triple-camera system. (e) Quadruple-camera system.
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2.7 Novel SPECT System Designs There are several special SPECT systems designs that do not fit into the preceding two general categories. The commercially available a vailable CERESPECT (formerly known as ASPECT) ASPECT) [Genna and Smith, 1988] is a dedicated brain SPECT system. It consists of a single fixed-annular NaI(TI) crystal that completely surrounds the patient’s head. Similar to a scintillation camera, an array of PMTs and electronics circuitry are placed behind the crystal to provide positional and energy information about photons that interact with the crystal. Projection data are obtained by rotating a segmented annular collimator with parallel holes that fits inside the stationary detector. A similar system is also being developed by Larsson et al. [1991] [ 1991] in Sweden.
Examples of novel SPECT system designs. (a) The CERESPECT brain SPECT system consists of a single fixed annular NaI(TI) crystal and a rotating segmented annular collimator. (b) The SPRINT II brain SPECT system consists of 11 detector modules and a rotating lead ring with slit opening.
2.8 Special Collimator Designs for SPECT Systems Similar to conventional nuclear medicine imaging, parallel-hole collimators are commonly used in camera-based SPECT systems. As described earlier, the tradeoff between dete ction efficiency and spatial resolution of parallel-hole collimator is a limitating factor for SPECT. A means to improve SPECT system performance is to improve the tradeoff imposed by the parallel-hole collimation. To achieve this goal, converging-hole collimator designs that increase the angle of acceptance of incoming photons without sacrificing spatial resolution have been developed. Examples are fan-beam [Jaszczak et al., 1979b; Tsui Tsui et al., 1986], cone beam [Jaszczak et al., 1987], astigmatic [Hawman and Hsieh, 1986], and more more recently varifocal collimators. As shown shown in Fig. 64.10b–d, 64.10b–d, the collimator holes converge to a line that is oriented parallel to the axis of rotation for a fan-beam collimator, to a point for a cone-beam collimator, and to various points for a varifocal collimator, respectively. The gain in detection efficiency of a typical fan-beam and cone-beam collimator is about 1.5 and 2 times of that of a parallel-hole collimator with the same spatial resolution. The anticipated gain in detection efficiency and corresponding decrease in image noise are the main reasons for the interest in applying converging- hole collimators in SPECT. SPECT. 16
Despite the advantage of increased detection efficiency effic iency,, the use of converging-hole collimators in SPECT poses special problems. The tradeoff for increase in detection efficiency as compared with parallel-hole collimators is a decrease in field of view .
Collimator designs used in camera-based SPECT systems. (a) The commonly used parallel-hole collimator. (b) The fan-beam collimator, where the collimator holes are converged to a line that is parallel to the axis of rotation. (c) The cone-beam collimator, where the collimator holes are converged to a point. (d) A varifocal collimator, collimator, where the collimator holes are converged to various focal points.
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2.9 SPECT/CT Hybrid Imagers Alogical approach to the issue of radionuclide localization is to have two scanners, one nuclear and one based on X-ray attenuation, located on attached gantries. This pair of devices shares the same patient couch. Because the distances of bed movement can be known within 1mm or less, the user can identify an uptake volume in the nuclear SPECT image with a geometrically corresponding part of the anatomy as seen via CT scan. Additionally, Additionally, attenuation corrections may be made more effectively using the CT data to improve SPECT sectional images. Some difficulties remain:(1) the breathing motion of the patient, and (2) possible changes in posture from one sequence to the other during the double imaging procedure. Complementary nature of the two images makes the interpretation of either somewhat clearer.
CT
SPECT
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SPECT/CT SPECT/C T
2.10 Final & Discussion The development of SPECT has been a combination of advances in radiopharmaceuticals, instrumentation, image processing and reconstruction methods, and clinical applications. Although substantial progress has been made during the last decade, there are many opportunities for contributions from biomedical engineering in the future. The future SPECT instrumentation will consist of more detector area to fully surround the patient for high detection efficiency and multiple contiguous transaxial slice capability. Multicamera SPECT systems will continue to dominate the commercial market. The use of new radiation detector materials and detector systems with high spatial resolution will receive increased attention. Continued research is needed to investigate special converging-hole collimator design geometries, fully 3D reconstruction algorithms, a lgorithms, and their clinical applications. To To improve image quality and to achieve quantitatively accurate SPECT images will continue to be the goals of image processing and image reconstruction r econstruction methods for SPECT. An important direction of research in analytical reconstruction methods will involve solving the inverse Radon transform, which includes the effects of attenuation, the spatially variant collimatordetector response function, and scatter. scatte r. The The development of iterative itera tive reconstruction methods will require more accurate models of the complex SPECT imaging process, faster and more stable iterative algorithms, and more powerful computers and special computational hardware. These improvements in SPECT instrumentation and image reconstruction methods, combined with newly developed radiopharmaceuticals, will bring SPECT images with increasingly higher quality and more accurate quantitation to nuclear medicine clinics for improved diagnosis and patient care.
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3
NUCLEAR MEDICINE INSTRUMENTATION
Positron Emission Tomography
3.1 Introduction
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3.2 A brief History of PET
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3.3 PET Scanner is a nuclear medicine imaging technique which produces a three-dimensional image or picture of functional processes in the body. body. The system detects pairs of gamma rays emitted indirectly by a positron-emitting radionuclide (tracer), which is introduced into the body on a biologically active molecule. Images of tracer concentration in 3-dimensional space within the body are then reconstructed by computer analysis. In modern scanners, this reconstruction is often accomplished with the aid of a CT X-ray scan performed on the patient during the same session, in the same machine. If the biologically active molecule chosen for PET is FDG, an analogue of glucose, the concentrations of tracer imaged then give tissue metabolic activity, in terms of regional glucose uptake. Although use of this tracer results in the most common type of PET scan, other tracer molecules are used in PET to image the tissue concentration of many other types of molecules of interest.
3.3.1 Positron The positron or antielectron is the antiparticle or the antimatter counterpart of the electron. The positron has an electric charge of +1, a spin of 1⁄2, and the same mass as an electron. When a low-energy positron collides with a low-energy electron, annihilation occurs, resulting in the production of two or more gamma ray photons (see electron-positron annihilation). The existence of positrons was first postulated in 1928 by Paul Dirac as a consequence of the Dirac equation. Positrons may be generated by positron emission radioactive decay (through weak interactions), or by pair production from a sufficiently energetic photon.
3.3.2 Annihilation Annihilation is defined as "total destruction" or "complete obliteration" of an object; having its root in the Latin nihil (nothing). A literal translation is "to make into nothing". In physics, the word is used to denote the process that occurs when a subatomic particle collides with its respective antiparticle. Since energy and momentum must be conserved, the particles are not actually made into nothing, but rather into new particles. Antiparticles have exactly opposite additive quantum numbers from particles, so the sums of all quantum numbers of the original pair are zero. Hence, any set of particles may be produced whose total quantum numbers are also zero as long as conservation of energy and conservation of momentum are obeyed. During a low-energy annihilation, photon production is favored, since these particles have no mass. However, high-energy particle colliders produce annihilations where a wide variety of exotic heavy particles are created.
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3.3.3 Electron–positron annihilation Electron–positron annihilation occurs when an electron and a positron (the electron's anti particle) collide. The result of the collision is the conversion of the electron and positron and the creation of gamma ray photons or, less often, other particles. The process must satisfy a number of conservation laws, including: • •
•
Conservation of charge. The net charge before and after is zero. Conservation of linear momentum and total energy. energy. This forbids the creation creati on of a single gamma ray. However, However, in quantum field theory this process is allowed Conservation of angular momentum.
As with any two charged objects, electrons and positrons may also interact with each other without annihilating, in general by elastic scattering.
Feynman Diagram of Electron-Positron Annihilation
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3.4 PET Detectors Efficient detection of the annihilation photons from positron emitters is usually provided by the combination of a crystal, which converts the high-energy photons to visible-light photons, and a photomultiplier tube that produces an amplified electric current pulse proportional to the amount of light photons interacting with the photocathode. The fact that imaging system sensitivity is proportional to the square of the detector efficiency leads to a very important requirement that the detector be nearly 100% efficient. Thus other detector systems such as plastic scintillators or gas-filled wire chambers, with typical individual efficiencies of 20% or less, would result in a coincident efficiency of only 4% or less. Most modern PET cameras are multilayered with 15 to 47 levels or transaxial layers to be reconstructed . The lead shields prevent prevent activity from the patient from causing spurious counts in the tomograph ring, while the tungsten septa reject some of the events in which one (or both) of the 511-keV photons photons suffer a Compton scatter in the patient. The sensitivity of this design is improved by collection of of data from cross-planes .
The physical basis of positron-emission tomography. Positrons emitted by “tagged” metabolically active molecules annihilate nearby electrons and give rise to a pair of high-energy photons. The photons fly off in nearly opposite directions and thus serve to pinpoint their source. The biologic activity of the tagged molecule can be used to investigate a number of physiologic functions, both normal and pathologic.
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Most modern PET cameras are multilayered with 15 to 47 levels or transaxial layers to be reconstructed. The lead shields prevent activity from the patient from causing spurious counts in the tomograph ring, while the tungsten septa reject some of the events in which one (or both) of the 511-keV photons suffer a Compton scatter in the patient. The sensitivity of this design is improved by collection of data from cross-planes.
The “individually coupled” design is capable of very high resolution, and because the design is very parallel (all the photomultiplier tubes and scintillator crystals operate independently) , it is capable of very high data throughput. The disadvantages of this type of design are the requirement for many expensive photomultiplier tubes and, additionally, that connecting round photomultiplier tubes to rectangular scintillation crystals leads to problems of packing rectangular crystals and circular phototubes of sufficiently small diameter to form a solid ring.
The arrangement of scintillators and phototubes is shown. The “individually coupled” design is capable of very high resolution, and because the design is very parallel (all the photomultiplier tubes and scintillator crystals operate independently), it is capable of very high data throughput. A block detector couples several photomultiplier tubes to a bank of scintillator crystals and uses a coding scheme to determine the crystal of interaction. In the two-layer block, five photomultiplier photomultiplier tubes are coupled to eight scintillator crystals .
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3.5 Photomultiplier tube (PMT)
Principle of operation of a photomultiplier tube (PMT). In our application the light source is a scintillation crystal. Together the BaF2 crystal and PMT form a detector for gamma rays.
3.6 Electronics 3.6.1 Discrimination of pulses with low amplitude amplitude or originating from noise As mentioned in chapter 2 Compton scattered gamma photons give a worsened resolution in PET applications. Therefore Therefore it is an advantage if one one can discriminate PM pulses pulses originating from events caused by the Compton effect. Since only a fraction of the energy of the original gamma photon is given to the electron in a Compton event, the pulses from the PM tube with low amplitude contain Compton events and also noise pulses. Such pulses are not wanted and can be taken away by a discriminator. In the discriminator a voltage level is set. Only pulses that are bigger than this level are accepted and give an output pulse from the discriminator. In our TOFPET application we use a special type of discriminator called a "constant fraction discriminator" (CFD). The CFD gives a good definition of the time of arrival of the electric pulse from a PM. Simultaneously the CFD performs an energy discrimination of pulses pulses of low amplitude. For those who are interested a brief description of a CFD is given below.
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3.6.2 Constant Fraction Discriminator In order to measure time intervals precisely, precisely, the arrival times of the different events must be derived exactly in order to achieve optimum time resolution. This is the function of the Constant Fraction Discriminator (CFD). The output pulse, from the anode of the PMT, PMT, is fed to the input of the CFD. In figure 4.1 the principle of operation of a CFD is illustrated.
Fig. 4.1. The formation of the constant-fraction signal
The CFD is designed to trigger on a certain optimum fraction of the pulse height, thus making the performance of the CFD independent of pulse amplitude. The input signal to the CFD is split into two parts. One part is attenuated attenuate d a fraction ƒ of the original amplitude V, V, the other part is delayed and inverted, see Fig. 4.2. These two signals are subsequently added to form form the constant-fraction-timing signal. The delay is chosen to make the optimum fraction point on the leading edge of the delayed pulse line up with the peak amplitude of the attenuated pulse. These two signals are subsequently added to form a bipolar pulse. The zero crossing occurs at a time after the arrival of the pulse that is independent of amplitude. The constantfraction discriminator incorporates a timing discriminator that triggers on the zero crossing and produces an output logic pulse that serves as the time marker. In addition a leading-edge discriminator provides energy selection. This energy selection constitutes the energy threshold used to suppress Compton scattered gamma photons in the PET-system. PET-system. No events with energy below the threshold will give rise to a signal from the CFD and thus will be excluded. The thresholds are set individually for all detectors by software.
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3.6.3 The PET Data Acquisition System In figure 4.2 the different parts of the PET system of detectors and electronics are illustrated. More details are found in figure 4.3. The Time to Digital Converter (TDC) converts a small time difference (ps - ns) to a digital value that can be handled numerically. The TOF-PET system system contains 48 cylindrical c ylindrical BaF2 crystals with a diameter of 15 mm and a length of 20 mm. These scintillators are optically coupled to the PMT with help of a silicon grease that transmits tra nsmits light down into ultra violet (UV) wave lengths. The PMTs are Hamamatsu R2076 and are equipped with 19 mm diameter synthetic silica windows. These windows are also transparent to the fast UV component of BaF2.
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3.6.4 Event Building Building and Software Software A dedicated C-program in the VME CPU reads the data from the TDC and builds the raw events. The 68030 VME CPU runs the operating system OS9000 suitable for real time event processing. The Data are sent to a Digital Unix workstation via a dual port RAM card. All electronics are VME based. The workstation receives the data via a PCI-based interface, and processes the raw events e vents with a C-program. This program is linked to a MATLAB MATLAB user interface, where the user can control the acquisition system. The user interface also includes a picture of the PET detector array and provides a feedback by drawing every coincidence coincidence line between the two individual coincident detectors, event by event. The event stream is also stored in a list mode file for off line processing.
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3.7 Three-Dimensional PET This technology doesn’t use any septa between planes with rings of crystals. Under this configuration, any coincidence event is possible between two small crystals in different transaxial planes. The sensitivity can be increased by a factor of ten, but the contribution of scattered photons can be around 30%.
3.8 Improvement of PET Scanners Major issues for improving the performance of PET scanners are (1) better resolution and quantitation of small lesions by correcting attenuation, scattering and partial volume effects, (2) monitoring and correcting involuntary patient motion, (3) compensating cardiac and respiratory motions by gating the register of activity activit y, (4) increasing energy e nergy resolution by using scintillators with better energy resolution (e.g., GSO) and (5) reduction of PET scanner time by using scintillators with lower light decay time (e.g.,LSO and GSO).
3.9 Hybrid PET/CT In addition to better correct photon attenuation in PET studies, the need of an anatomical frameworkfor interpreting the high uptake of 18FDG has led to the development of of hybrid PET/CT scanners. The high uptake of 18FDG by the tumor produces very high-localized intensity PET images in which it is difficult to see other organs and structures. To To evaluate the exact localization of the tumor, it is necessary that the anatomic framework be provided by the CT scan. The automatic motion of the patient table between them performs emission and transmission modalities sequentially. Emission and transmission images are registered automatically. PET/CT hybrid hybrid scanners have combined standard CT scanners with PET systems in such a way that CT can be used separately for clinical studies.162 The first hybrid PET/ CT introduced in the market was the Discovery LS (Figure 10.24) from GE Medical Systems. The instrument combines a GE Lightspeed CT scanner with the GE Advance PET scanner. The The first Discovery LS was installed at the University Hospital, Zurich in March 2001. Other commercial PET/CT scanners already introduced into the market or that will be introduced soon are the Biograph PET/CT (Siemens), which is a combination of the ECAT ECAT EXACT HR+ PET and the SOMATON SOMATON Emotion, and the Gemini PET/CT (Philips Medical Systems), which is a combination c ombination of the Allegro PET (former ADAC) ADAC) and the CT MX8000 (former Marconi). A current and very promising area of research and clinical application is the use of PET/CT images for radiation treatment planning. Modern radiotherapy treatments based on the concept of intensity modulated radiation therapy (IMRT) need to define precisely the location of the target volume, metabolic extension and heterogeneity of the tumoral tissue. By fusing the metabolic 18FDG PET image with the CT structural information, the most metabolically active areas can be identified and improve the delivery of lethal radiation doses to these regions.
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PET whole body scan and CT anatomical reference. PET and fused PET/CT images indicate uptake of 18 FDG in liver and other abdominal areas. From left to right: CT, CT, PET,and fused PET/CT coronal (top) and sagittal (bottom) views. The fourth column shows the transaxial images, from top to bottom: CT, CT, PET, and fused PET/CT images.
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3.10 Final PET is both a medical and research tool. It is used heavily in clinical oncology (medical imaging of tumors and the search for metastases), and for clinical diagnosis of certain diffuse brain diseases such as those causing various types of dementias. PET is also an important research tool to map normal human brain and heart function. PET is also used in pre-clinical studies using animals, where it allows repeated investigations into the same subjects. This is particularly valuable in cancer research, as it results in an increase in the statistical quality of the data (subjects can act as their own control) and substantially reduces the numbers of animals required for a given study. study. Alternative methods of scanning include x-ray computed tomography (CT), magnetic resonance imaging (MRI) and functional magnetic resonance imaging (fMRI), ultrasound and single photon emission computed tomography (SPECT). While some imaging scans such as CT and MRI isolate organic anatomic a natomic changes in the body, PET and SPECT are capable of detecting areas of molecular biology detail (even prior to anatomic change). PET scanning does this using radiolabelled molecular probes that have different rates of uptake depending on the type and function of tissue involved. Changing of regional blood flow in various anatomic structures (as a measure of the injected positron emitter) can be visualized and relatively quantified with a PET scan. PET imaging is best performed using a dedicated PET scanner. However, However, it is possible to acquire PET images using a conventional dual-head gamma camera fitted with a coincidence detector. The quality of gamma-camera PET is considerably lower, and acquisition is slower. slower. However, for institutions with low demand for PET, this may allow on-site imaging, instead of referring patients to another center, or relying on a visit by a mobile scanner.
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References : *Biochemical Engineering and Biotechnology - G.D. Najafpour N ajafpour *Biomedical Engineering Handbook – J.D.Bronzino J.D.Bronzino *Biomedical Information Technology - David D. Feng *CRC Press - Biomedical Photonics Handbook *CRC Press - Biomedical Technology and Devices Handbook *Kluwer - Handbook of Biomedical Image Analysis *Wiley - Encyclopedia of Medical Devices and Instrumentation **Wikipedia, the free encyclopedia ( http://en.wikipedia.org ) ** Nuclear Physics Group. www.nuclear.kth.se **ZENTRALINSTITUT FÜR ELEKTRONIK www.fz-juelich.de **imXgam http://imxgam.in2p3.fr **University of Virginia www.med-ed.virginia.edu/ **Paramedic http://paramedic.emszone.com/ **National Center for Biotechnology Information http://www.ncbi.nlm.nih.gov/
