Nuclear Medicine Instrumentation
1. Medical background
The unprecedented development in many areas of medicine during the last 50 years is to a large degree
due to the availability of new diagnostic instruments. From the beginning of medical imaging in 1895
when Roentgen took the first X-ray picture, many methods and techniques have been developed, which
were adapted or inspired from nuclear and particle physics instrumentation. A large part of medical
imaging techniques concerns the viewing of anatomical structures of the body. X-ray computed
tomography (CT) and magnetic resonance imaging (MRI) are sophisticated techniques of this type
which yield high resolution images of physical structure parameters. However, in medical research and
in the diagnosis of many medical conditions it is often necessary to use functional information. It is thus
highly desirable to be able to acquire images of physiologic function to complement images of the
anatomy. Certain biologically active pharmaceuticals which concentrate in different organs of the
human body, when these are chemically labeled with specific radioactive materials and administered
within to a patient, the distribution and functional information can be deduced from the gamma ray
emission pattern. This class of imaging is known as nuclear medicine imaging. Nuclear medicine
techniques are applicable in the diagnosis of a wide variety of diseases, they can be used for tumor
detection, imaging of abnormal body functions and to quantitatively measure the heart or the brain
function, for example since blood flow in the brain is tightly coupled to local brain metabolism and
energy use, the 99mTc-HMPAO (hexamethylpropylene amine oxime) tracer can be used to assess brain
metabolism regionally, in an attempt to diagnose and differentiate the different causal pathologies of
dementia. Depending on the method of the radio pharmaceuticals labeling, nuclear medicine techniques
can be split into: gamma or positron emission methods
2. Gamma Camera
The first gamma camera was developed by Hal Anger in 1957, and his original design is still widely
used today. This device gives projective images, when imaging with a gamma camera the patient is
injected with a small amount of a radioactive imaging agent chosen to accumulate in specific regions.
The radioisotope disintegrates by emitting a single gamma photon with an energy typically between 80
and 350 keV. 99mTc (140keV) is the most frequently used. The components making up the Gamma
camera are the collimator, Scintillator, front end electronic and data acquisition system (Figure 1. 1.).
NaI(Tl) Scintillator
Crystal
Collimator
Light guide
Phototomultiplier Tube
Figure 1. 1. Conventional gamma camera
The collimator is made of gamma ray absorbing material (Lead or Tungsten), which acts to select a
given direction of photons incident to the camera. In parallel hole collimator only photons traveling in a
direction parallel to the collimator holes will reach the scintillator detector. The collimator defines the
geometrical field of view of the camera and determines both the spatial resolution and sensitivity of the
system. After the collimator the gamma-rays are detected by a large single thallium-activated sodium
iodide NaI(Ti)) scintillator crystal, typically about 50 cm in diameter. The interaction of the a gammaray with the crystal results in scintillation photons emitted isotropically. The emitted photons are
detected by array of photomultiplier tubes (PMTs) which are optically coupled to the surface of the
crystal detector. A PMT consist of two elements, a photo cathode coupled to an electron multiplier, the
photo cathodes ejects electrons when stimulated by light photons. Electron multiplier consists of an
arrangement of dynodes that serves both as efficient collection geometry for the photoelectrons and a
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 strong charge signal that can be
collected at the anode. At the photomultiplier output a amplifier, filters, shaper and line-driver are used
to adapt and transport the pulse to the data acquisition system were the two coordinates of the
interaction position are extracted from the amplitude distribution of the PMT signals and the total
energy is obtained from their sum. The total sum allows discrimination between different isotopes or
between scattered and direct photons. The data are then sent to a computer to process it into a readable
image of spatial distribution of the organ activity.
3. Computed Tomography
Computed tomography (CT) is a medical imaging method where digital geometry processing is used to
generate a three-dimensional image slices of an object from a large set of two-dimensional images
(projections) taken around a single axis of rotation. The mathematical principles of tomographic
reconstruction have been known for a long time, but tomographic imaging techniques had to await the
digital revolution before being implemented due to the computing power needed. The first practical
application was made in 1960’s. Then in the 1970’s there was an explosion of activity with several
techniques being developed simultaneously, most notably positron emission tomography (PET) and Xray computed tomography (CT). CT was invented in 1972 by Godfrey Hounsfield and Allan Cormack
and they were later (1979) awarded the Nobel Prize for their contributions to medicine and science. The
computed tomography technique has applications in non-medical fields as well, for example in
astronomy where images taken at different times from different angles are combined and a more
accurate composite image is constructed. The back projection algorithm was the first algorithm used for
the image reconstruction, but because it produces blurred images projected data needs to be filtered
before reconstruction. Depends on the application different filters can be applied. Filtered back
projection algorithm is one of the most used reconstruction methods.
4. Single photon emission computed tomography (SPECT)
SPECT imaging is performed by rotating a gamma camera around the object (or patient) and acquiring
projections from multiple angles (Fig.1.2.).
Figure 1. 2. A standard SPECT by rotating a single head gamma camera
A topographic reconstruction algorithm is then applied to the 2D projections to reconstruct a 3D image,
which can be used to show thin slices of the position and concentration of radionuclide distribution
along a chosen axis. The use of the collimator results in low sensitivity, in order to increase the
sensitivity and hence the scanning time, the SPECT are now being equipped with multi-detector
systems for total-body, brain and heart scanning. The systems with 2 or 3 Anger type gamma cameras
180o, 120o or 90o apart with small or large filed of view are most popular. These systems have
considerably improved the sensitivity, resolution and the scanning time comparing to the single camera
systems. The parallel collimator can be replaced by pinhole collimator which allows geometric
enlargements and obtain high resolution specialty for imaging small objects. The variable attenuation
by the body of the photons emitted from the organ of interests and the incorrect counting of scattered
photons in detector crystal can lead to significant underestimation of activity inside the body, compared
to the surface and hence an image degradation. This problem is solved by incorporating a map of
attenuation coefficients from transmission scan by X-ray CT into the reconstruction algorithm. Iterative
reconstructions algorithms [1] are alternative algorithms which are growing in importance are used to
reduce the image degradation.
Despite being efficient and widely used, SPECT with the conventional gamma camera suffer from
some limitations such us the non-uniformity, image distortion and degradation of the position towards
the edge of the camera, it is bulky and heavy. The trend is towards SPECT systems based on compact
and Dedicated gamma cameras for specific clinical applications and for small animal (molecular
imaging). The proposed devices use new semiconductor detectors (Ge, CdTe or CdZnTe) where the
gamma rays are converted directly to digital electronic signals, continuous or pixellated scintillator
crystals coupled to an array of solid state photodiodes or to a Position Sensitive Photo-Multiplier Tube.
5. Positron emission tomography (PET)
Positron emission tomography use biologically active tracers labeled with a position-emitting isotope.
Following the decay, the emitted positron slows down and annihilates with electrons in tissue
producing two back to back 511-keV photons. The detection of these two photons in coincidence
defines a line along which the emission point is known to be located. By accumulating many
annihilation events, the spatial distribution of the isotope, and therefore of the labeled
radiopharmaceutical is reconstructed. The electronic coincidence detection of the two collinear 511 keV
photons limits the position of the source to a line, there is no need of the collimator which is one of the
great advantage of PET. Figure1. 3. is a schematic of the principle of the PET technique.
Figure 1. 3. A schematic of the P ET principle
5.1. PET detectors and detector configurations
5.1.1. Detector materials
There are three important practical features affecting the PET detector performances, the attenuation
coefficient for the 511-keV photon, the light output, and the speed (decay time of the scintillator). The
attenuation coefficient determines the detector sensitivity and it should be high, the energy resolution is
improved with the high light output and a fast crystal leads to high count rate and narrow coincidence
window and thus less random count rate. The choice of the detector depends on the application and the
cost. The NaI(Tl) were used in early PET detectors, it has high light output and it is cheap, but it is
stopping power is suboptimal compared to other crystals. Due it is greatest attenuation coefficient and it
is cost BGO is widely used in current commercial PET systems. The disadvantages of BGO compared
to other crystals include its low light output compared to NaI(Tl), high decay time compared to LSO,
and it is emission at 500nm where PMTs are less sensitive. LSO is the most suitable crystal for today
PET scanners; it has short decay time, high light lead, and high density. Table 1 resumes the
characteristics of the crystals used for PET.
5.1.2. Detector configurations
Different detector configurations have been developed for commercial PET scanners. The simple
configuration is where the detectors are mounted on a rotating gantry (Fig. 6. a). For high sensitivity
cameras the detectors are mounted on circular or polygonal rings. (Fig. 6. b) and (Fig. 6. c). modern
PET systems have multiple rings configuration. Such systems have high performance at the expense of
increased cost.
Scintillator
crystals
PMT A
Y
PMT C
PMT D
X
PMT B
Figure1.1. 4 Block detector configuration for PET
Conventional detectors consists of single scintillator coupled to a single PMT,today the most common
configuration used for commercial scanners and small animal PET is the block detector configuration.
Block detector consists of 2D array of single crystals from a cut of large cubic crystal and the cuts are
filled with reflective material to optically isolate the single crystals from each other. The array is
coupled to 4 PMTs with light sharing. State-of-art PET are using PSMT for more accurate positional
and energy information. Solid state detectors offer high segmentation, hence higher spatial resolution.
They are promising candidates for future PET generations.
5.2. Two dimensional (2D) and three dimensional (3D)
PET scanners can be designed to operate in two dimensional (2D) or 3D mode. In 2D mode thin septa
of lead or tungsten separate each crystal ring and coincidences are only recorded between detectors
within the same ring or lying in closely neighboring rings (Fig. 7.a). This reduces the contributions
from scatter and random coincidences with consequent reduction in the overall sensitivity. In 3D mode,
the septa are removed, and coincidences are recorded between detectors lying in any ring combination
(Fig. 7.b). This results in a fully 3D image and in substantial increase in the sensitivity, at the expense
of an increased scatter fraction and count rate performance.
a
b
Figure 1.1. 5 PET acquisition mode a) 2D with septa b) 3D without septa
5.3. Types of coincidence events
Coincidence events in PET can be classified as: true, scattered and random (Fig. 8).True coincidences
occur when both photons from an annihilation event are detected by detectors in coincidence, neither
photon undergoes any form of interaction prior to detection, and no other event is detected within the
coincidence time-window. A scattered coincidence is when at least one of the detected photons has
undergone Compton scattering event prior to detection and because the direction of the photon is
changed the event will be assigned to wrong line of response. Scattered coincidences add a background
to the true coincidence distribution and causing the isotope concentrations to be overestimated. Random
coincidences occur when two photons not arising from the same annihilation event are incident on the
detectors within the coincidence time window, the rate of random coincidences increase roughly with
the square of the activity in the field of view (FOV). Random coincidences add statistical noise to the
data. The number of scattered and random events detected depends on the volume and attenuation
characteristics of the object being imaged, and on the geometry of the camera.
Scattered event
True event
Random event
Figure 1.1. 6. Types of coincidence events in PET
6. New detectors for PET and SPECT
6.1. Scintillators
Inorganic scintillator crystals are the most commonly used detectors for Gamma camera. The gamma
photon interact within scintillator through Rayleigh, Compton, and photoelectric effects. The photon
deposit it is energy at one location by photoelectric effect or at different position by Compton
interaction. The absorbed energy causes the crystal to make a transition to higher energy state, from
which it may undergo decay after a characteristic time by emitting lower energy photons that are
detected by photodetector. Photoelectric and Compton cross-sections are a function of the density (ρ)
and the effective atomic number (Zeff) of the crystal. A high density favors the interaction of the photon
in the crystal, whilst a higher Zeff value increases the number of photoelectric occurrences with respect
to Compton scattering. Therefore, high Zeff crystals are to be referred. The light yield and the decay time
are important physical properties of the crystal. A high light output (number of photons per Mev)
implies a better energy resolution, hence high positioning accuracy. Short decay time leads to high
counting rat. The scintillation photon wavelength has to much the properties of the photodetectors.
NaI(Tl) is the scintillator of choice for imaging with gamma camera due to it is light output
41000(Ph/Mev), which allows energy resolution of the order 9-11%(FWHM) at 140 keV. The
drawback of NaI(T)l is that it cannot be produced in shape of long and thin pillars that can be
arranged in segmented scintillator array due to its hygroscopicity. CsI(Tl) and YAP:Ce are suitable
crystals for these geometries (Fig.2). In Table 1 we list the characteristic of most of the crystals
available for gamma detection.
Scintillator
material
Density
(g/cm 3)
Zeff
Attenuation
lenght for
511-Kev
gammas
(mm)
Light output
(ph/Mev)
Decay time
(ns)
Emission
wave length
(nm)
Refractive
index
BGO
7.1
75
10.4
9000
300
480
2.15
LSO
7.4
66
11.4
30000
40
420
1.82
NaI(Tl)
3.67
51
29.1
41000
230
410
1.85
CsI(Tl)
4.51
52
22.9
66000
900
550
1.80
GSO
6.7
59
14.1
8000
60
440
1.85
LuAP
8.3
64.9
10.5
12000
18
365
1.94
YAP
5.5
33.5
21.3
17000
30
350
1.94
Table 1. 1. Physical properties of scintillator materials commonly used for SPECT and PET
6.2. Photodetectors
6.2.1. Position-sensitive photomultiplier (PSMT)
A PSPMT is position sensitive which means that photons impinging at different sites on the
photocathode would give rise to pulses of varying amplitudes at the anodes outputs. Typical The output
signal is read using independent multiple anodes where each anode is connected to an individual
independent electronic chain or by means of current or charge-dividing center-of gravity where the
anodes are connected through a resistive network and position signals are read from the two ends for
both X and Y direction. For each event, time of interaction is provided by the last dynode. PSPMT
coupled to a thin scintillator such as CsI(Tl) or NaI(Tl) are an attractive approach achieving good
spatial and energy resolution.
6.2.2. Silicon photodiodes
Photodiodes are semiconductor devices with a pn junction or pin structure where light absorbed in a
depletion region generates electrical carriers and thus a photocurrent. Such devices can be very
compact, fast, highly linear, and exhibit a high quantum efficiency. The most used for gamma ray
imaging is the avalanche photodiode. Avalanche photodiode is a semiconductor-based photodiode
which is operated with a relatively high reverse voltage (typically tens or even hundreds of volts),
sometimes just below breakdown. In this regime, carriers (electrons and holes) excited by absorbed
photons quickly get accelerated in the strong internal electric field, so that they can generate secondary
carriers, as it is known in photomultipliers. The advantage of APDs is the high siagnal to noise ration
due to high gain, and the faster response relative to PMTs and Photodiodes . APDs can be produced in
arrays and used for very compact systems.
6.2.3. Semiconductor Detectors
Semiconductor detectors are promising alternative to scintillators for gamma ray imaging due to their
superiority for energy resolution, and they can be easily pixelated and read out directly. The most
attractive semiconductor material is cadmium zinc telluride (CdZnTe). CdZnTe has very good
absorption characteristics due to its high density(5.8 g/cm3) and effective atomic number (50), the
attenuation length of 140 keV gamma-rays is being only 3 mm the photofraction is 81% , the
absorption of a 140 keV produces approximately 3*104 which is a good charge yield , hence an
excellent energy resolution . Due to the wide band gap it has a high resistivity which results in low
noise characteristic.
7. Compton cameras
In Compton camera the kinematics of Compton scatter is used as an electronic collimator to reconstruct the
radioisotope distribution. The basic concept of Compton camera consists of two detectors scatter and absorber
separated by a known distance, both with very good energy and spatial resolution. when gamma photon inter the
camera , it interacts with the first detector and scatter to the second detector (Figure 1. 6).
Detector 1
Detector 2
Figure 1.4.4: Principal of Compton camera
By using the two detectors in coincidence, the total energy and the direction of travel from the first detector to the
second can be extracted for each event. This information in conjunction with the Compton equation (bellow) can
be used to back project a cone of all possible direction of the incident gamma ray.

1
1 
cos   1  mo c 2 

E E
E 
re
 
where  is the cone semi-aperture and the Compton scatter angle; Ere is the recoil electron energy in the front
detector; E is the source energy; and m oc2 is the rest mass of an electron (511keV). A large number of
scattering events from a point source of gamma rays will define multiple cones which intersect at the
location of the source (Figure 1.4.2).
Figure 1.4.2: Cones intersection determine the interaction point
8. SPECT versus PET
The use of the collimator in SPECT results in tremendous decrease in the resolution and the efficiency
compared to PET where collimation is performed electronically which leads to high sensitivity. High
sensitivity improve signal to noise ratio which correspond to improvement in image. With PET the
whole scan is performed in short time compared to SPECT and multiple field of view can be scanned.
Although SPECT imaging resolution is many times less than PET, the high cost of PET scanners, the
need of the accelerator close the examination place, he availability of low cost SPECT pharmaceutical
and the practical and economic aspects of SPECT instrumentation makes this technique of functional
imaging attractive for a lot of clinical studies specially for the brain and the heart.