*2.1.1.3. The photomultipliers tubes*

Their role is to convert light energy emitted by the crystal to an electrical signal that can be exploited in electronic circuits [3, 5]. This is achieved by the combination of several elements, placed in a vacuum to allow the flow of electrons. The first element, placed in contact with the crystal is the photocathode, metal foil on which the light photons are able to extract electrons. These electrons are attracted to the first dynode by the application of a high voltage between it (positively charged) and the photocathode. The electrons acceleration allows them to extract a much larger number of electrons from the dynode. Then there are several cascading dynodes, on which the same phenomenon is repeated. The successive dynodes are submitted to potentials higher and higher. From a dynode to another, we obtain a cascade of electrons more intense (amplification phenomenon), which ultimately results in a measurable electric current. This current is collected by the last element called anode and a real electrical signal is generated (Figure 4).

The use of radioactive tracers that are introduced in the living system to study its metabolism dates from 1923 when de Hevesy and Paneth studied the transport of radioactive lead in plants [6]. In 1935, de Hevesy and Chiewitz were the first to apply the method to the study of the distribution of a radiotracer (P-32) in rats [7]. The major development of scintigraphic imaging started with the invention of the gamma camera by Anger in 1956 [1]. In parallel, positron imaging was developed. Both imaging modalities are now standard in the major nuclear

Principles and Applications of Nuclear Medical Imaging: A Survey on Recent Developments

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The tracer principle, which forms the basis of nuclear imaging, is the following: a radioactive biologically active substance is chosen in such a way that its spatial and temporal distribution in the body reflects a particular body function or metabolism. In order to study the distribution without disturbing the body function, only traces of the substance are administered to the

The radiotracer decays by emitting gamma rays or positrons (followed by annihilation gamma rays).The distribution of the radioactive tracer is inferred from the detected gamma rays and

The most often used radio-nuclides are Tc-99m in 'single photon' imaging and F-18 in 'positron' imaging. Tc-99m is the decay daughter of Mo-99 which itself is a fission product of U. The half-life of Tc-99m is 6h, which is optimal for most metabolic studies but too short to allow for long time storage. Mo-99 has a half-life of 65h. This allows a Mo-99 generator (a 'cow') to be stored and Tc-99m to be 'milked' when required. Tc-99m decays to Tc-99 by emitting a gamma ray with an energy output of 14O keV. This energy is optimal for detection by scintillator detectors. Tc-99 itself has a half-life of 211100 years and is therefore a negligible

F-18 is cyclotron produced and has a half-life of 110 minutes. It decays to stable O-18 by emitting a positron. The positron loses its kinetic energy through Coulomb interactions with surrounding nuclei. When it is nearly at rest, which in tissue occurs after an average range of less than 1 mm, the probability of a collision with an electron greatly increases and becomes one. During the collision matter-antimatter annihilation occurs in which the rest mass of the electron and the positron is transformed into two gamma rays of 511 keV. The two gamma rays originate at exactly the same time (they are "coincident") and leave the point of collision

**1.** Static acquisition with a detector in a fixed position relative the patient: examination of

**2.** Scanning of the whole body: succession of static images joined: the detector move simultaneously and scan the patient's body from head to foot. The bone scan is a routine

**3.** Tomographic acquisition: The Positron Emission Single Photon (SPECT): detectors rotate around the patient to obtain in a digital representation of a 3D radioactive distribution of

medicine departments.

mapped as a function of time and/or space.

burden to the patient [8, 9].

in almost opposite directions [9].

the body: chest, pelvis, skull....

thyroid, kidney....

application.

Different modalities of scintigraphic acquisition are possible:

patient [8, 9].

**Figure 4.** PMTs disposition in a Gamma-camera. Generally a hexagonal shape of PTM is preferred then a circular be‐ cause it well cover the detection area. Additional very small PMT can also be used between principal PMT for best de‐ tection area covering (CEM, Rennes, France).

### *2.1.2. Gamma scintigraphic imaging*

Scintigraphy is a method designed to reproduce the shape or to measure the activity of an organ by administering a product which contains an element which emits radioactivity, an isotope. The radioactivity emitted by the isotope is picked up by special detectors called gamma-cameras counters described above. Generally, the dose is administered to a patient in need of scintigraphy is safe for the body (except for pregnancy). The data acquisition principle is illustrated on the diagram of Figure 5.

**Figure 5.** Illustration of data acquisition in planer gamma scintigraphy.

The use of radioactive tracers that are introduced in the living system to study its metabolism dates from 1923 when de Hevesy and Paneth studied the transport of radioactive lead in plants [6]. In 1935, de Hevesy and Chiewitz were the first to apply the method to the study of the distribution of a radiotracer (P-32) in rats [7]. The major development of scintigraphic imaging started with the invention of the gamma camera by Anger in 1956 [1]. In parallel, positron imaging was developed. Both imaging modalities are now standard in the major nuclear medicine departments.

it (positively charged) and the photocathode. The electrons acceleration allows them to extract a much larger number of electrons from the dynode. Then there are several cascading dynodes, on which the same phenomenon is repeated. The successive dynodes are submitted to potentials higher and higher. From a dynode to another, we obtain a cascade of electrons more intense (amplification phenomenon), which ultimately results in a measurable electric current. This current is collected by the last element called anode and a real electrical signal is generated

8 Imaging and Radioanalytical Techniques in Interdisciplinary Research - Fundamentals and Cutting Edge Applications

**Figure 4.** PMTs disposition in a Gamma-camera. Generally a hexagonal shape of PTM is preferred then a circular be‐ cause it well cover the detection area. Additional very small PMT can also be used between principal PMT for best de‐

Scintigraphy is a method designed to reproduce the shape or to measure the activity of an organ by administering a product which contains an element which emits radioactivity, an isotope. The radioactivity emitted by the isotope is picked up by special detectors called gamma-cameras counters described above. Generally, the dose is administered to a patient in need of scintigraphy is safe for the body (except for pregnancy). The data acquisition principle

PositionData Linearity and uniformity

corrections

Spectrometry correction and analysis

EnergyData

Gamma camera

**Figure 5.** Illustration of data acquisition in planer gamma scintigraphy.

(Figure 4).

tection area covering (CEM, Rennes, France).

*2.1.2. Gamma scintigraphic imaging*

is illustrated on the diagram of Figure 5.

The tracer principle, which forms the basis of nuclear imaging, is the following: a radioactive biologically active substance is chosen in such a way that its spatial and temporal distribution in the body reflects a particular body function or metabolism. In order to study the distribution without disturbing the body function, only traces of the substance are administered to the patient [8, 9].

The radiotracer decays by emitting gamma rays or positrons (followed by annihilation gamma rays).The distribution of the radioactive tracer is inferred from the detected gamma rays and mapped as a function of time and/or space.

The most often used radio-nuclides are Tc-99m in 'single photon' imaging and F-18 in 'positron' imaging. Tc-99m is the decay daughter of Mo-99 which itself is a fission product of U. The half-life of Tc-99m is 6h, which is optimal for most metabolic studies but too short to allow for long time storage. Mo-99 has a half-life of 65h. This allows a Mo-99 generator (a 'cow') to be stored and Tc-99m to be 'milked' when required. Tc-99m decays to Tc-99 by emitting a gamma ray with an energy output of 14O keV. This energy is optimal for detection by scintillator detectors. Tc-99 itself has a half-life of 211100 years and is therefore a negligible burden to the patient [8, 9].

F-18 is cyclotron produced and has a half-life of 110 minutes. It decays to stable O-18 by emitting a positron. The positron loses its kinetic energy through Coulomb interactions with surrounding nuclei. When it is nearly at rest, which in tissue occurs after an average range of less than 1 mm, the probability of a collision with an electron greatly increases and becomes one. During the collision matter-antimatter annihilation occurs in which the rest mass of the electron and the positron is transformed into two gamma rays of 511 keV. The two gamma rays originate at exactly the same time (they are "coincident") and leave the point of collision in almost opposite directions [9].

Different modalities of scintigraphic acquisition are possible:


### **2.2. Single photon emission computed tomography**

This medical imaging method was introduced in 1963 by Kuhl and Edwards [10]. Known by the acronym SPECT (Single Photon Emission Computerized Tomography), this imaging method is equivalent in scintigraphy to Computed Tomography (CT) in radiology. The injected radioactive tracers emit during their disintegration gamma photons which are detected by an external detector, after passing through the surrounding tissue. Because the gamma photons emission is isotropic, a collimator is placed before the detector to select the direction of the photons to be detected. Thus, if we call f(x, y, z) the distribution of radioactivity emitted point {x, y, z} per unit solid angle, the number of photons detected at the point {x',y'} of the detector is equal to (Figure 6) [11]:

$$\mathbf{N}(\mathbf{x}', \mathbf{y}') = \int\_{\mathcal{L}} \mathbf{f}(\mathbf{x}, \mathbf{y}, \mathbf{z}) d\mathbf{s} \tag{1}$$

In SPECT, the main radioactive isotopes are technetium-99m, Iodine and Thallium-201, which is used primarily for studies on the heart. At the opposite of PET system, the collimator is an indispensable component in a SPECT machine. The first collimators used were two-dimen‐ sional parallel channels (Figure 7, a). By rotating the detector & collimator assembly around the patient, two-dimensional projections are obtained, and the distribution of radioactivity may be 3D reconstructed slice by slice. These parallel collimators are used in the vast majority of SPECT systems used in Nuclear Medicine services. The resolution of these systems varied

Principles and Applications of Nuclear Medical Imaging: A Survey on Recent Developments

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11

To increase the sensitivity and resolution of SPECT systems, converging channels collimators were developed (Figure 7, b). The first proposed included a series of converging channels to a focal line which is parallel to the rotation axis of the system [12]. This system is therefore equivalent to a scanner used in X-ray fan beam tomography where 3D image is reconstructed slice by slice. For imaging small organs such as heart and brain, a converging cone collimators is used [13, 14]. This last collimator allows obtaining magnification of the object in all directions (cross and longitudinal). This kind of collimators can be used only for small field tests, so for small structures, the size of the detectors has not increased. With these systems, image data registration is completely 3D as well as in cone beam X-ray tomography, and therefore reconstruction is not performed slice by slice. In these systems, it is important to be able to shift the head of the detector relatively to the rotation axis, thereby to perform trajectories other than circular. In addition to the fact that this shift allow to complete the set of projections, such a shift is interesting to avoid obstacles, such as shoulders brain imaging. Finally, other kinds of collimators are also available for SPECT such as diverging and pinhole collimators. Diverging collimator (Figure 7, c) is reserved to large structure imaging. Pin-hole collimator (Figure 7, d) allows obtaining a mirror image with a variable magnification function of collimator depth and object to collimator distance. This collimator is suitable for small

Positron emission tomography (PET) is a medical imaging modality that measures the threedimensional distribution of a molecule labelled with a positron emitter. The acquisition is carried out by a set of detectors arranged around the patient. The detectors consist of a scintillator which is selected according to many properties, to improve the efficiency and the signal on noise. The coincidence circuit measures the two 511 keV gamma photons emitted in opposite directions resulting from the annihilation of the positron. The sections were recon‐ structed by algorithms, the same but more complex than those used for conventional CT, to accommodate the three-dimensional acquisition geometries. Correction by considering the physical phenomena provides an image representative of the distribution of the tracer. In PET scan an effective dose of the order of 8 mSv is delivered to the patient. This technique is in permanent evolution, both from the point of view of the detector and that of the used image reconstruction algorithms. A new generation of hybrid scanner "PET-CT" provides additional information for correcting the attenuation, localize lesions and to optimize therapeutic procedures. All these developments make one PET fully operational tool that has its place in

from 10 to 15 mm.

structures imaging such as thyroid and hip.

**2.3. Positron emission tomography**

medical imaging.

Where L is the line given by the direction of the channel's collimator and passing through the point (x',y'). As in CT, the various projections are obtained by rotating the detector around the object (patient).

**Figure 6.** Detection principle in SPECT imaging.

<sup>1</sup> ECG : Electrocardiogram.

In SPECT, the main radioactive isotopes are technetium-99m, Iodine and Thallium-201, which is used primarily for studies on the heart. At the opposite of PET system, the collimator is an indispensable component in a SPECT machine. The first collimators used were two-dimen‐ sional parallel channels (Figure 7, a). By rotating the detector & collimator assembly around the patient, two-dimensional projections are obtained, and the distribution of radioactivity may be 3D reconstructed slice by slice. These parallel collimators are used in the vast majority of SPECT systems used in Nuclear Medicine services. The resolution of these systems varied from 10 to 15 mm.

To increase the sensitivity and resolution of SPECT systems, converging channels collimators were developed (Figure 7, b). The first proposed included a series of converging channels to a focal line which is parallel to the rotation axis of the system [12]. This system is therefore equivalent to a scanner used in X-ray fan beam tomography where 3D image is reconstructed slice by slice. For imaging small organs such as heart and brain, a converging cone collimators is used [13, 14]. This last collimator allows obtaining magnification of the object in all directions (cross and longitudinal). This kind of collimators can be used only for small field tests, so for small structures, the size of the detectors has not increased. With these systems, image data registration is completely 3D as well as in cone beam X-ray tomography, and therefore reconstruction is not performed slice by slice. In these systems, it is important to be able to shift the head of the detector relatively to the rotation axis, thereby to perform trajectories other than circular. In addition to the fact that this shift allow to complete the set of projections, such a shift is interesting to avoid obstacles, such as shoulders brain imaging. Finally, other kinds of collimators are also available for SPECT such as diverging and pinhole collimators. Diverging collimator (Figure 7, c) is reserved to large structure imaging. Pin-hole collimator (Figure 7, d) allows obtaining a mirror image with a variable magnification function of collimator depth and object to collimator distance. This collimator is suitable for small structures imaging such as thyroid and hip.
