# Basic Physics of Nuclear Medicine/Gas-Filled Radiation Detectors

![](Geiger_counter.jpg "Geiger_counter.jpg")

We have learned in the last two chapters about how radiation interacts
with matter and we are now in a position to apply our understanding to
the detection of radiation.

One of the major outcomes of the interaction of radiation with matter is
the creation of ions as we saw in Chapter
5.
This outcome is exploited in gas-filled detectors as you will see in
this chapter. The detector in this case is essentially a gas, in that it
is the atoms of a gas which are ionised by the radiation. We will see in
the next
chapter
that solids can also be used as radiation detectors but for now we will
deal with gases and be introduced to detectors such as the Ionization
Chamber and the Geiger
Counter.

Before considering these specific types of gas-filled detectors we will
first of all consider the situation from a very general perspective.

## Gas-Filled Detectors

As we noted above the radiation interacts with gas atoms in this form of
detector and causes ions to be produced. On the basis of what we covered
in Chapter
5
it is easy to appreciate that it is the Photoelectric and Compton
Effects that cause the ionisations when the radiation consists of
gamma-rays with energies useful for diagnostic purposes.

There are actually two particles generated when an ion is produced - the
positive ion itself and an electron. These two particles are
collectively called an **ion pair**. The detection of the production of
ion pairs in the gas is the basis upon which gas detectors operate. The
manner in which this is done is by using an electric field to sweep the
electrons away to a positively charged electrode and the ions to a
negatively charged electrode.

Let us consider a very simple arrangement as shown in the following
figure:

```{=html}
<div class="center">
```
![](NM8_1.gif "NM8_1.gif"){width="320"}

```{=html}
</div>
```
Here we have two electrodes with the gas between them. Something like a
capacitor with a gas dielectric.

The gas which is used is typically an inert gas, for example
argon or xenon. The reason
for using an inert gas is so that chemical reactions will not occur
within the gas following the ionisations which could change the
characteristics of our detector.

A dc voltage is placed between the two electrodes. As a result when the
radiation interacts with a gas atom the electron will move towards the
positive electrode and the ion will move towards the negative electrode.
But will these charges reach their respective electrodes? The answer is
obviously dependent on the magnitude of the dc voltage. For example if
at one extreme we had a dc voltage of a microvolt (that is, one
millionth of a volt) the resultant electric field may be insufficient to
move the ion pair very far and the two particles may recombine to reform
the gas atom. At the other extreme suppose we applied a million volts
between the two electrodes. In this case we are likely to get sparks
flying between the two electrodes - a lightning bolt if you like - and
our detector might act something like a neon
sign. Somewhere in between these two extremes
though we should be able to provide a sufficient attractive force for
the ion and electron to move to their respective electrodes without
recombination or sparking occurring.

We will look at this subject in more detail below. Before we do let us
see how the concept of the simple detector illustrated above is applied
in practice. The gas-filled chamber is generally cylindrical in shape in
real detectors. This shape has been found to be more efficient than the
parallel electrode arrangement shown above.

A cross-sectional view through this cylinder is shown in the following
figure:

```{=html}
<div class="center">
```
![](NM8_2.gif "NM8_2.gif"){width="320"}

```{=html}
</div>
```
The positive electrode consists of a thin wire running through the
centre of the cylinder and the negative electrode consists of the wall
of the cylinder. In principle we could make such a detector by getting a
section of a metal pipe, mounting a wire through its centre, filling it
with an inert gas and sealing the ends of the pipe. Actual detectors are
a little bit more complex however but let us not get side-tracked at
this stage.

We apply a dc voltage via a battery or via a dc voltage supply and
connect it as shown in the figure using a resistor, R. Now, assume that
a gamma-ray enters the detector. Ion pairs will be produced in the gas -
the ions heading towards the outer wall and the electrons heading
towards the centre wire. Let us think about the electrons for a moment.
When they hit the centre wire we can simply think of them as entering
the wire and flowing through the resistor to get to the positive
terminal of the dc voltage supply. These electrons flowing through the
resistor constitute an electric current and as a result of Ohm\'s
Law a voltage is generated across the
resistor. This voltage is amplified by an amplifier and some type of
device is used to register the amplified voltage. A loud-speaker is a
fairly simple device to use for this purpose and the generation of a
voltage pulse is manifest by a click from the loud-speaker. Other
display devices include a **ratemeter** which displays the number of
voltage pulses generated per unit time - something like a speedometer in
a car - and a pulse counter (or **scaler**) which counts the number of
voltage pulses generated in a set period of time. A voltage pulse is
frequently referred to in practice as a **count** and the number of
voltage pulses generated per unit time is frequently called the **count
rate**.

## DC Voltage Dependence

If we were to build a detector and electronic circuit as shown in the
figure above we could conduct an experiment that would allow us to
explore the effect of the dc voltage on the magnitude of the voltage
pulses produced across the resistor, R. Note that the term **pulse
height** is frequently used in this field to refer to the magnitude of
voltage pulses.

Ideally, we could generate a result similar to that illustrated in the
following figure:

```{=html}
<div class="center">
```
![](NM8_3.gif "NM8_3.gif"){width="320"}

```{=html}
</div>
```
The graph illustrates the dependence of the pulse height on the dc
voltage. Note that the vertical axis representing the pulse height is on
a logarithmic scale for the sake of compressing a large linear scale
onto a reasonably-sized graph.

The experimental results can be divided into five regions as shown. We
will now consider each region in turn.

-   **Region A**     Here V~dc~ is relatively low so that recombination
    of positive ions and electrons occurs. As a result not all ion pairs
    are collected and the voltage pulse height is relatively low. It
    does increase as the dc voltage increases however as the amount of
    recombination reduces.

```{=html}
<!-- -->
```
-   **Region B**     V~dc~ is sufficiently high in this region so that
    only a negligible amount of recombination occurs. This is the region
    where a type of detector called the **Ionization Chamber** operates.

```{=html}
<!-- -->
```
-   **Region C**     V~dc~ is sufficiently high in this region so that
    electrons approaching the centre wire attain sufficient energy
    between collisions with the electrons of gas atoms to produce new
    ion pairs. Thus the number of electrons is increased so that the
    electric charge passing through the resistor, R, may be up to a
    thousand times greater than the charge produced initially by the
    radiation interaction. This is the region where a type of detector
    called the Proportional Counter
    operates.

```{=html}
<!-- -->
```
-   **Region D**     V~dc~ is so high that even a minimally-ionizing
    particle will produce a very large voltage pulse. The initial
    ionization produced by the radiation triggers a complete gas
    breakdown as an avalanche of electrons heads towards and spreads
    along the centre wire. This region is called the **Geiger-Müller
    Region**, and is exploited in the Geiger Counter.

```{=html}
<!-- -->
```
-   **Region E**     Here V~dc~ is high enough for the gas to completely
    breakdown and it cannot be used to detect radiation.

We will now consider features of the Ionisation Chamber and the Geiger
Counter in more detail.

## Ionisation Chamber

The ionisation chamber consists of a gas-filled detector energised by a
relatively low dc voltage. We will first of all make an estimate of the
voltage pulse height generated by this type of detector. We will then
consider some applications of ionisation chambers.

When a beta-particle interacts with the gas the energy required to
produce one ion pair is about 30 eV. Therefore when a beta-particle of
energy 1 MeV is completely absorbed in the gas the number of ion pairs
produced is:

```{=html}
<div class="center">
```
$n = \frac{1\ \text{MeV}}{30\ \text{eV}} = \frac{1 \cdot 10^6}{30} \approx 3 \cdot 10^4\ \text{ion pairs}$

```{=html}
</div>
```
The electric charge produced in the gas is therefore

```{=html}
<div class="center">
```
$Q = n \cdot e$

```{=html}
</div>
```
```{=html}
<div class="center">
```
$\therefore (3 \cdot 10^4\ \text{ion pairs}) \cdot (1.6 \cdot 10^{-19}\ \text{C})$

```{=html}
</div>
```
```{=html}
<div class="center">
```
$\therefore Q = 5 \cdot 10^{-15}\ \text{C}$

```{=html}
</div>
```
If the capacitance of the ionisation chamber (remember that we compared
a gas-filled detector to a capacitor above) is 100 pF then the amplitude
of the voltage pulse generated is:

```{=html}
<div class="center">
```
$V = \frac{Q}{C} = \frac{5 \cdot 10^{-15}\ \text{C}}{100 \cdot 10^{-12}\ \text{F}} = 5 \cdot 10^{-5}\ \text{V}$

```{=html}
</div>
```
```{=html}
<div class="center">
```
$\therefore V = 50\ \mu\text{V}$

```{=html}
</div>
```
Because such a small voltage is generated it is necessary to use a very
sensitive amplifier in the electronic circuitry connected to the
chamber.

We will now learn about two applications of ionisation chambers. The
first one is for the measurement of radiation exposures. You will
remember from Chapter 4
that the unit of radiation exposure (be it the SI or the traditional
unit) is defined in terms of the amount of electric charge produced in a
unit mass of a air. An ionization chamber filled with air is the natural
instrument to use for such measurements.

The second application is the measurement of radioactivity. The
ionisation chamber used here is configured in what is called a
**re-entrant** arrangement (see figure below) so that the sample of
radioactive material can be placed within the detector using a holder
and hence most of the emitted radiation can be detected. The instrument
is widely referred to as an **Isotope Calibrator** and the trickle of
electric current generated by such a detector is calibrated so that a
reading in units of radioactivity (for example MBq or mCi) can be
obtained. Most well-run Nuclear Medicine Departments will have at least
one of these devices so that doses of radioactivity can be checked prior
to administration to patients.

![](NM8_10.gif "NM8_10.gif"){width="320"}

Here are some photographs of ionisation chambers designed for various
applications:

  --------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- --------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- ---------------------------------------------------------------------------------------------------------------------
  !An exposure-area product detector used in radiography.{width="150"}                                                                                !A range of ionisation chambers of different volumes using for measuring radiation exposure.{width="150"}   !An exposure meter used in radiography.{width="150"}
  !An isotope calibrator used in nuclear medicine - the blue cylinder on the left contains the re-entrant chamber.{width="150"}   !An exposure meter used in radiography.{width="150"}                                                                                !A contemporary Geiger counter.{width="150"}
  --------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- --------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- ---------------------------------------------------------------------------------------------------------------------

## Geiger Counter

We saw earlier that the Geiger Counter operates at relatively high dc
voltages (for example 400-900 volts) and that an avalanche of electrons
is generated following the absorption of radiation in the gas. The
voltage pulses produced by this detector are relatively large since the
gas effectively acts as an amplifier of the electric charge produced.

There are four features of this detector which we will discuss. The
first is that a sensitive amplifier (as was the case with the Ionization
Chamber) is not required for this detector because of the gas
amplification noted above.

The second feature results from the fact that the generation of the
electron avalanche must be stopped in order to reform the detector. In
other words when a radiation particle/photon is absorbed by the gas a
complete gas breakdown occurs which implies that the gas is incapable of
detecting the next particle/photon which enters the detector. So in the
extreme case one minute we have a radiation detector and the following
moment we do not.

A means of stopping the electron avalanche is therefore required - a
process called **Quenching**. One means of doing this is by
electronically lowering the dc voltage following an avalanche. A more
widely used method of quenching is to add a small amount of a
**quenching gas** to the inert gas. For example the gas could be argon
with ethyl alcohol added. The ethyl
alcohol is in vapour form and since it consists of relatively large
molecules energy which would in their absence give rise to sustaining
the electron avalanche is absorbed by these molecules. The large
molecules act like a brake in effect.

Irrespective of the type of quenching used the detector is insensitive
for a small period of time following absorption of a radiation
particle/photon. This period of time is called the Dead
Time and this is the third feature of this
detector which we will consider. Dead times are relatively short but
nevertheless significant - being typically of the order of 200-400 µs.
As a result the reading obtained with this detector is less than it
should be. The true count rate, T, can be obtained using the following
equation:

```{=html}
<div class="center">
```
$T = \frac{A}{1 - \tau A}$

```{=html}
</div>
```
where A is the (actual) reading and τ is the dead time. Some instruments
perform this calculation automatically.

The fourth feature to note about this detector is the dependence of its
performance on the dc voltage. The Geiger-Müller Region of our figure
above is shown in more detail below:

![](NM8_12.gif "NM8_12.gif"){width="320"}

Notice that it contains a plateau where the count rate obtained is
independent of the dc voltage. The centre of this plateau is where most
detectors are operated. It is clear that the count rate from the
detector is not affected if the dc voltage fluctuates about the
operating voltage. This implies that a relatively straight-forward dc
voltage supply can be used. This feature coupled with the fact that a
sensitive amplifier is not needed translates in practice to a relatively
inexpensive radiation detector.

## External Links

-   Inside a smoke
    detector - about the
    ion chamber used in smoke detectors - from the How Stuff Works
    website.
-   Ionisation
    Chambers - a
    brief description from the Triumf Safety Group.
-   Radiation and
    Radioactivity -
    a self-paced lesson developed by the University of Michigan\'s
    Student Chapter of the Health Physics Society with a section on gas
    filled detectors.
-   The Geiger
    Counter - a
    brief overview from the NASA Goddard Space Flight Center, USA.


# Basic Physics of Nuclear Medicine/Scintillation Detectors

![](PET-detectorsystem.png)

The second type of radiation detector we will discuss is called the
scintillation detector.
Scintillations "wikilink") are minute
flashes of light which are produced by certain materials when they
absorb radiation. These materials are variously called fluorescent
materials, fluors, scintillators or
phosphors.

If we had a radioactive source and a scintillator in the lab we could
darken the room, move the scintillator close to the source and see the
scintillations. These small flashes of light might be green or blue or
some other colour depending on the scintillator. We could also count the
number of flashes produced to gain an estimate of the radioactivity of
the source, that is the more flashes of light seen the more radiation
present.

The scintillation detector was possibly the first radiation detector
discovered. You might have heard the story of the discovery of X-rays by
Wilhelm Roentgen in 1895. He was working
one evening in his laboratory in Wurzburg, Germany with a device which
fired a beam of electrons at a target inside an evacuated glass tube.
While working with this device he noticed that some platino-barium
cyanide crystals, which he just happened to have close by, began to glow
-- and that they stopped glowing when he switched the device off.
Roentgen had accidentally discovered a new form of radiation. He had
also accidentally discovered a scintillator detector.

Although scintillations can be seen we have a more sophisticated way of
counting and measuring them today by using some form of
photodetector.

We will learn about the construction and mode of operation of this type
of detector in this chapter. In addition, we will see how it can be used
not just for detecting the presence of ionizing radiation but also for
measuring the **energy** of that radiation.

Before we do however it is useful to note that scintillators are very
widely used in the medical radiations field. For example the X-ray
cassette used in radiography contains a
scintillator (called an **intensifying screen**) in close contact with a
photographic film. A second example is the **X-ray Image Intensifier**
used in fluoroscopy which contains
scintillators called phosphors. Scintillators are also used in some CT
Scanners and as we will see in the
next
chapter,
in the Gamma Camera and PET
Scanner. Their application
is not limited to the medical radiations field in that scintillators are
also used as screens in television sets and
computer monitors and for generating
light in fluorescent tubes -- to
mention just two common applications. What other applications can you
think of?

So scintillators are a lot more common than you might initially think
and you will therefore find the information presented here useful to you
not just for your studies of nuclear medicine.

## Fluorescent Materials

Some fluorescent materials are listed in the following table.
Thallium-activated sodium iodide, NaI(Tl)
is a crystalline material which is widely used for the detection of
gamma-rays in scintillation detectors. We will be looking at this in
more detail below.

Another crystalline material sodium-activated caesium
iodide, CsI(Na) is widely used for X-ray
detection in devices such as the X-ray image
intensifier. Another one called
calcium tungstate, CaWO~4~ has been widely used in X-ray cassettes
although this substance has been replaced by other scintillators such as
lanthanum oxybromide in many modern cassettes.

  Material                     Form
  ---------------------------- ---------
  NaI (Tl)                     crystal
  CsI (Na)                     crystal
  CaWO~4~                      crystal
  ZnS (Ag)                     powder
  p-terphenyl in toluene       liquid
  p-terphenyl in polystyrene   plastic

Notice that some scintillation materials are activated with certain
elements. What this means is that the base material has a small amount
of the activation element present. The term
doped "wikilink") is sometimes used instead
of **activated**. This activating element is used to influence the
wavelength (colour) of the light produced by the scintillator.

Silver-activated zinc sulphide is a
scintillator in powder form and p-terphenyl in toluene is a liquid
scintillator. The advantage
of such forms of scintillators is that the radioactive material can be
placed in close contact with the scintillating material. For example if
a radioactive sample happened to be in liquid form we could mix it with
a liquid scintillator so as to optimise the chances of detection of the
emitted radiation and hence have a very sensitive detector.

A final example is p-terphenyl in polystyrene which is a scintillator in
the form of a plastic. This form can be easily made into different
shapes like most plastics and is therefore useful when detectors of
particular shapes are required.

## Photomultiplier Tube

A scintillation crystal coupled to a photomultiplier
tube (PMT) is illustrated in the
following figure. The overall device is typically cylindrical in shape
and the figure shows a cross-section through this cylinder:

![](NM9_1.gif "NM9_1.gif")

The scintillation crystal, NaI(Tl) is very delicate and this is one of
the reasons it is housed in an aluminium casing. The inside wall of the
casing is designed so that any light which strikes it is reflected
downwards towards the PMT.

The PMT itself consists of a photocathode, a focussing grid, an array of
dynodes and an anode housed in an evacuated glass tube. The function of
the photocathode is to convert the light flashes produced by radiation
attenuation in the scintillation crystal into electrons. The grid
focuses these electrons onto the first dynode and the dynode array is
used for electron multiplication. We will consider this process in more
detail below. Finally the anode collects the electrons produced by the
array of dynodes.

The electrical circuitry which is typically attached to a PMT is shown
in the next figure:

![](NM9_2.gif "NM9_2.gif")

It consists of a high voltage supply, a resistor divider chain and a
load resistor, R~L~. The high voltage supply generates a dc voltage,
V~dc~ which can be up to 1,000 volts. It is applied to the resistor
divider chain which consists of an array of resistors, each of which has
the same resistance, R. The function of this chain of resistors is to
divide up V~dc~ into equal voltages which are supplied to the dynodes.
As a result voltages which increase in equal steps are applied to the
array of dynodes. The load resistor is used so that an output voltage,
V~out~ can be generated.

Finally the operation of the device is illustrated in the figure below:

![](NM9_3.gif "NM9_3.gif")

The ionizing radiation produces flashes of light in the scintillation
crystal. This light strikes the photocathode and is converted into
electrons. The electrons are directed by the grid onto the first dynode.

Dynodes are made from certain alloys which emit electrons when their
surface is struck by electrons with the advantage that more electrons
are emitted than are absorbed. A dynode used in a PMT typically emits
between two and five electrons for each electron which strikes it.

So when an electron from the photocathode strikes the first dynode
between two and five electrons are emitted and are directed towards the
second dynode in the array (three are illustrated in the figure). This
electron multiplication process is repeated at the second dynode so that
we end up with nine electrons for example heading towards the third
dynode. An electron avalanche therefore develops so that a sizeable
number of electrons eventually hits the anode at the bottom of the
dynode chain.

These electrons flow through the load resistor, R~L~ and constitute an
electric current which according to Ohm\'s Law generates a voltage,
V~out~ which is measured by electronic circuitry (which we will describe
later).

A number of photographs of devices based on scintillation detection are
shown below:

  ------------------------------------------------------------------------------------------------------------------------------------------------- -------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- -------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------
  !An X-ray cassette used in radiography.{width="150"}   !A NaI crystal coupled to a photomultiplier tube.{width="150"}                 !A number of identical photomultiplier tubes from a Gamma Camera.{width="150"}
  !A single channel analyser.{width="150"}                                !A single channel analyser with a crystal-PMT assembly.{width="150"}   !A well detector.{width="150"}
  ------------------------------------------------------------------------------------------------------------------------------------------------- -------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- -------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------

The **important feature** of the scintillation detector is that this
output voltage, V~out~ is directly proportional to the energy deposited
by the radiation in the crystal. We will see what a useful feature this
is below. Before we do so we will briefly analyze the operation of this
device.

## Mathematical Model

A simple mathematical model will be presented below which will help us
get a better handle on the performance of a scintillation detector. We
will do this by quantifying the performance of the scintillator, the
photocathode and the dynodes.

Let\'s use the following symbols to characterize each stage of the
detection process:

-   *m*: number of light photons produced in crystal
-   *k*: optical efficiency of the crystal, that is the efficiency with
    which the crystal transmits light
-   *l*: quantum efficiency of the
    photocathode, that is the efficiency with which the photocathode
    converts light photons to electrons
-   *n*: number of dynodes
-   *R*: dynode multiplication factor, that is the number of secondary
    electrons emitted by a dynode per primary electron absorbed.

Therefore the charge collected at the anode is given by the following
equation:

```{=html}
<div class="center">
```
$Q = m k l R^n e\,\!$

```{=html}
</div>
```
where *e*: the electronic charge.

For example supposing a 100 keV gamma-ray is absorbed in the crystal.
The number of light photons produced, m, might be about 1,000 for a
typical scintillation crystal. A typical crystal might have an optical
efficiency, k, of 0.5 -- in other words 50% of the light produced
reaches the photocathode which might have a quantum efficiency of 0.15.
A typical PMT has ten dynodes and let us assume that the dynode
multiplication factor is 4.5.

Therefore

```{=html}
<div class="center">
```
$Q = 1000(0.5)(0.15)(4.5^{10})(1.6 \cdot 10^{-19})\ \text{C}$

```{=html}
</div>
```
```{=html}
<div class="center">
```
$\therefore Q = 41 \cdot 10^{-12}\ \text{C}$

```{=html}
</div>
```
```{=html}
<div class="center">
```
$\therefore Q \approx 40\ \text{pC}$

```{=html}
</div>
```
This amount of charge is very small. Even though we have used a
sophisticated photodetector like a PMT we still end up with quite a
small electrical signal.

A very sensitive amplifier is therefore needed to amplify this signal.
This type of amplifier is generally called a **pre-amplifier** and we
will refer to it again later.

## Output Voltage

We noted above that the voltage measured across the resistor, R~L~, is
proportional to the energy deposited in the scintillation crystal by the
radiation. Let us consider how the radiation might deposit its energy in
the crystal.

Let us consider a situation where gamma-rays are detected by the
crystal. We learnt in Chapter
5
that there were two interaction mechanisms involved in gamma-ray
attenuation -- the Photoelectric Effect and the Compton Effect. You will
remember that the Photoelectric Effect involves the total absorption of
the energy of a gamma-ray, while the Compton Effect involves just
partial absorption of this energy. Since the output voltage of a
scintillation detector is proportional to the energy deposited by the
gamma-rays it is reasonable to expect that Photoelectric Effects in the
crystal will generate distinct and relatively large output voltages and
that Compton Effects will result in lower output voltages.

The usual way of presenting this information is by plotting a graph of
the count rate versus the output voltage pulse height as shown in the
following figure:

![](NM9_8.gif "NM9_8.gif")

This plot illustrates what is obtained for a monoenergetic
gamma-emitting radioisotope, for example ^99m^Tc -- which, as we have
noted before emits a single gamma-ray with an energy of 140 keV.

Before we look at it in detail remember that we noted above that the
output voltage from this detector is proportional to the energy
deposited by the radiation in the crystal. The horizontal axis can
therefore be used to represent the output voltage or the gamma-ray
energy. Both of these quantities are shown in the figure to help with
this discussion. In addition note that this plot is often called a
**Gamma-Ray Energy Spectrum**.

The figure above contains two regions. One called the **Photopeak** and
the other called the **Compton Smear**. The Photopeak results because of
Photoelectric absorption of the gamma-rays from the radioactive source
-- remember that we are dealing with a monoenergetic emitter in this
example. It consists of a peak representing the gamma-ray energy (140
keV in our example). If our radioisotope emitted gamma-rays of two
energies we would have two photopeaks in our spectrum and so on.

Notice that the peak has a statistical spread. This has to do with how
good our detector is and we will not get into any detail about it here
other than to note that the extent of this spread is a measure of the
quality of our detector. A high quality (and more expensive!) detector
will have a narrower statistical spread in the photopeaks which it
measures.

The other component of our spectrum is the Compton Smear. It represents
a range of output voltages which are lower than that for the Photopeak.
It is therefore indicative of the partial absorption of the energy of
gamma-rays in the crystal. In some Compton Effects a substantial
scattering with a valence electron can occur which gives rise to
relatively large voltage pulses. In other Compton Effects the gamma-ray
just grazes off a valence electron with minimal energy transfer and
hence a relatively small voltage pulse is generated. In between these
two extremes are a range of scattering events involving a range of
energy transfers and hence a range of voltage pulse heights. A \'smear\'
therefore manifests itself on the gamma-ray energy spectrum.

It is important to note that the spectrum illustrated in the figure is
simplified for the sake of this introductory discussion and that actual
spectra are a little more complex -- see figure below for an example:

![](NM9_4a.gif "NM9_4a.gif"){width="320"}

You will find though that your understanding of actual spectra can
easily develop on the basis of the simple picture we have painted here.

It is also important to appreciate the additional information which this
type of radiation detector provides relative to a gas-filled detector.
In essence gas-filled detectors can be used to tell us if any radiation
is present as well as the amount of that radiation. Scintillation
detectors also give us this information but they tell us about the
**energy** of this radiation as well. This additional information can be
used for many diverse applications such as the identification of unknown
radioisotopes and the production of nuclear medicine images. Let us stay
a little bit longer though with the fundamental features of how
scintillation detectors work.

The photopeak of the Gamma-Ray Energy Spectrum is generally of interest
in nuclear medicine. This peak is the main signature of the radioisotope
being used and its isolation from the Compton Smear is normally achieved
using a technique called **Pulse Height Analysis**.

## Pulse Height Analysis

This is an electronic technique which allows a spectrum to be acquired
using two types of circuitry. One circuit is called a **Lower Level
Discriminator** which only allows voltages pulses through it which are
higher than its setting. The other is called an **Upper Level
Discriminator** which only allows voltage pulses though which are lower
than its setting.

The result of using both these circuits in combination is a
variable-width window which can be placed anywhere along a spectrum. For
example if we wished to obtain information from the photopeak only of
our simplified spectrum we would place the discrimination controls as
shown in the following figure:

![](NM9_9.gif "NM9_9.gif")

A final point to note here is that since the scintillation detector is
widely used to obtain information about the energies of the radiation
emitted from a radioactive source it is frequently referred to as a
**Scintillation Spectrometer**.

## Scintillation Spectrometer

Types of scintillation spectrometer fall into two basic categories --
the relatively straight-forward **Single Channel Analyser** and the more
sophisticated **Multi-Channel Analyser**.

The Single Channel Analyser is the type of instrument we have been
describing so far in this discussion. A block diagram of the instrument
is shown below:

![](NM9_10.gif "NM9_10.gif"){width="320"}

It consists of a scintillation crystal coupled to a photomultiplier tube
which is powered by a high voltage circuit (H.V.). The output voltages
are initially amplified by a sensitive pre-amplifier (**Pre-Amp**) as we
noted above before being amplified further and conditioned by the
amplifier (**Amp**).

The voltage pulses are then in a suitable form for the pulse height
analyser (**P.H.A.**) -- the output pulses from which can be fed to a
**Scaler** and a **Ratemeter** for display of the information about the
portion of the spectrum we have allowed to pass through the PHA. The
Ratemeter is a display device just like the speedometer in a car and
indicates the number of pulses generated per unit time. The Scaler on
the other hand usually consists of a digital display which shows the
number of voltage pulses produced in a specified period of time.

We can illustrate the operation of this circuitry by considering how it
might be used to generate a Gamma-Ray Energy Spectrum. What we would do
is set up the LLD and ULD so as to define a narrow window and place this
to pass the lowest voltage pulses produced by the detector through to
the Scaler and Ratemeter. In other words we would place a narrow window
at the extreme left of the spectrum and acquire information about the
lowest energy gamma-ray interactions in the crystal. We would then
adjust the LLD and ULD settings to acquire information about the
interactions of the next highest energy. We would proceed in this
fashion to scan the whole spectrum.

A more sophisticated detector circuit is illustrated in the following
figure:

![](NM9_11.gif "NM9_11.gif"){width="320"}

It is quite similar to that in the previous figure with the exception
that the PHA, Scaler and Ratemeter are replaced by a Multi-Channel
Analyser and a computer. The Multi-Channel Analyser (**MCA**) is a
circuit which is capable of setting up a large number of individual
windows to look at a complete spectrum in one go. The MCA might consist
of 1024 individual windows for example and the computer might consist of
a personal computer which can acquire information simultaneously from
each window and display it as an energy spectrum. The computer generally
contains software which allows us to manipulate the resultant
information in a variety of ways. Indeed the ^137^Cs spectrum shown
above was generated using this approach.

## External links


# Basic Physics of Nuclear Medicine/Nuclear Medicine Imaging Systems

!`<span style="color:white;">`{=html}`</span>`{=html}

Topics we have covered in this wikibook have included radioactivity, the
interaction of gamma-rays with matter and radiation detection. The main
reason for following this pathway was to bring us to the subject of this
chapter: nuclear medicine imaging systems. These are devices which
produce pictures of the distribution of radioactive material following
administration to a patient.

The radioactivity is generally administered to the patient in the form
of a radiopharmaceutical -- the term
**radiotracer** is also used. This follows some physiological pathway to
accumulate for a short period of time in some part of the body. A good
example is ^99m^Tc-tin colloid which following intravenous injection
accumulates mainly in the patient\'s liver. The substance emits
gamma-rays while it is in the patient\'s liver and we can produce an
image of its distribution using a nuclear medicine imaging system. This
image can tell us whether the function of the liver is normal or
abnormal or if sections of it are damaged from some form of disease.

Different radiopharmaceuticals are used to produce images from almost
every region of the body:

  Part of the Body     Example Radiotracer
  -------------------- -------------------------------------------------------------------
  Brain                ^99m^Tc-HMPAO_exametazime "wikilink")
  Thyroid              Na^99m^TcO~4~
  Lung (Ventilation)   ^133^Xe gas
  Lung (Perfusion)     ^99m^Tc-MAA_albumin_aggregated "wikilink")
  Liver                ^99m^Tc-Tin (or Sulphur) Colloid
  Spleen               ^99m^Tc-Damaged Red Blood Cells
  Pancreas             ^75^Se-Selenomethionine
  Kidneys              ^99m^Tc-DMSA

Note that the form of information obtained using this imaging method is
mainly related to the physiological functioning of an organ as opposed
to the mainly anatomical information which is obtained using X-ray
imaging systems. Nuclear medicine therefore provides a different
perspective on a disease condition and generates additional information
to that obtained from X-ray images. Our purpose here is to concentrate
on the imaging systems used to produce the images.

Early forms of imaging system used in this field consisted of a
radiation detector (a scintillation detector for example) which was
scanned slowly over a region of the patient in order to measure the
radiation intensity emitted from individual points within the region.
One such device was called the Rectilinear
Scanner. Such imaging systems have
been replaced since the 1970s by more sophisticated devices which
produce images much more rapidly. The most common of these modern
devices is called the Gamma Camera and we
will consider its construction and mode of operation below. A review of
recent developments in this technology for cardiac applications can be
found in Slomka et al (2009)[^1].

## Gamma Camera

The basic design of the most common type of gamma camera used today was
developed by an American physicist, Hal Anger
and is therefore sometimes called the Anger Camera. It consists of a
large diameter NaI(Tl) scintillation crystal which is viewed by a large
number of photomultiplier tubes.

A block diagram of the basic components of a gamma camera is shown
below:

!Block diagram of a gamma
camera

The crystal and PM Tubes are housed in a cylindrical shaped housing
commonly called the **camera head** and a cross-sectional view of this
is shown in the figure. The crystal can be between about 25 cm and 40 cm
in diameter and about 1 cm thick. The diameter is dependent on the
application of the device. For example a 25 cm diameter crystal might be
used for a camera designed for cardiac applications while a larger 40 cm
crystal would be used for producing images of the lungs. The thickness
of the crystal is chosen so that it provides good detection for the 140
keV gamma-rays emitted from ^99m^Tc -- which is the most common
radioisotope used today.

Scintillations produced in the crystal are detected by a large number of
PM tubes which are arranged in a two-dimensional array. There are
typically between 37 and 91 PM tubes in modern gamma cameras. The output
voltages generated by these PM tubes are fed to a position circuit which
produces four output signals called ±X and ±Y. These position signals
contain information about where the scintillations were produced within
the crystal. In the most basic gamma camera design they are fed to a
cathode ray oscilloscope (**CRO**). We will
describe the operation of the CRO in more detail below.

Before we do so we should note that the position signals also contain
information about the intensity of each scintillation. This intensity
information can be derived from the position signals by feeding them to
a summation circuit (marked ∑ in the figure) which adds up the four
position signals to generate a voltage pulse which represents the
intensity of a scintillation. This voltage pulse is commonly called the
**Z-pulse** which, following pulse height analysis, (**PHA**) is fed as
the **unblank pulse** to the CRO.

So we end up with four position signals and an unblank pulse sent to the
CRO. Let us briefly review the operation of a CRO before we continue.
The core of a CRO consists of an evacuated tube with an electron gun at
one end and a phosphor-coated screen at the other end. The electron gun
generates an electron beam which is directed at the screen and the
screen emits light at those points struck by the electron beam. The
position of the electron beam can be controlled by vertical and
horizontal deflection plates and with the appropriate voltages fed to
these plates the electron beam can be positioned at any point on the
screen. The normal mode of operation of an oscilloscope is for the
electron beam to remain switched on. In the case of the gamma camera the
electron beam of the CRO is normally switched off -- it is said to be
**blanked**.

When an unblank pulse is generated by the PHA circuit the electron beam
of the CRO is switched on for a brief period of time so as to display a
flash of light on the screen. In other words the voltage pulse from the
PHA circuit is used to unblank the electron beam of the CRO.

So where does this flash of light occur on the screen of the CRO? The
position of the flash of light is dictated by the ±X and ±Y signals
generated by the position circuit. These signals as you might have
guessed are fed to the deflection plates of the CRO so as to cause the
unblanked electron beam to strike the screen at a point related to where
the scintillation was originally produced in the NaI(Tl) crystal.
Simple!

The gamma camera can therefore be considered to be a sophisticated
arrangement of electronic circuits used to translate the position of a
flash of light in a scintillation crystal to a flash of light at a
related point on the screen of an oscilloscope. In addition the use of a
pulse height analyser in the circuitry allows us to translate the
scintillations related only to photoelectric events in the crystal by
rejecting all voltage pulses except those occurring within the photopeak
of the gamma-ray energy spectrum.

Let us **summarise** where we have got to before we proceed. A
radiopharmaceutical is administered to the patient and it accumulates in
the organ of interest. Gamma-rays are emitted in all directions from the
organ and those heading in the direction of the gamma camera enter the
crystal and produce scintillations (note that there is a device in front
of the crystal called a **collimator** which we will discuss later). The
scintillations are detected by an array of PM tubes whose outputs are
fed to a position circuit which generates four voltage pulses related to
the position of a scintillation within the crystal. These voltage pulses
are fed to the deflection circuitry of the CRO. They are also fed to a
summation circuit whose output (the Z-pulse) is fed to the PHA and the
output of the PHA is used to switch on (that is, unblank) the electron
beam of the CRO. A flash of light appears on the screen of the CRO at a
point related to where the scintillation occurred within the NaI(Tl)
crystal. An image of the distribution of the radiopharmaceutical within
the organ is therefore formed on the screen of the CRO when the
gamma-rays emitted from the organ are detected by the crystal.

What we have described above is the operation of a fairly traditional
gamma camera. Modern designs are a good deal more complex but the basic
design has remained much the same as has been described. One area where
major design improvements have occurred is the area of image formation
and display. The most basic approach to image formation is to photograph
the screen of the CRO over a period of time to allow integration of the
light flashes to form an image on photographic
film. A stage up from this is to use a
storage oscilloscope which allows each
flash of light to remain on the screen for a reasonable period of time.

The most modern approach is to feed the position and energy signals into
the memory circuitry of a computer for storage. The memory contents can
therefore be displayed on a computer monitor and can also be manipulated
(that is **processed**) in many ways. For example various colours can be
used to represent different concentrations of a radiopharmaceutical
within an organ.

The use of digital image
processing is now widespread in
nuclear medicine in that it can be used to rapidly and conveniently
control image acquisition and display as well as to analyse an image or
sequences of images, to annotate images with the patient\'s name and
examination details, to store the images for subsequent retrieval and to
communicate the image data to other computers over a
network.

The essential elements of a modern gamma camera are shown in the next
figure. Gamma rays emitted by the patient pass through the collimator
and are detected within the camera head, which generates data related to
the location of scintillations in the crystal as well as to the energy
of the gamma rays. This data is then processed on-the-fly by electronic
hardware which corrects for technical factors such as spatial linearity,
PM tube drift and energy response so as to produce an imaging system
with a spatially-uniform sensitivity and distortion-free performance.

A multichannel analyzer (MCA) is used to display the energy spectrum of
gamma rays which interact inside the crystal. Since these gamma rays
originate from within the patient, some of them will have an energy
lower than the photopeak as a result of being scattered as they travel
through the patient\'s tissues -- and by other components such as the
patient table and structures of the imaging system. Some of these
scattering events may involve just glancing interactions with free
electrons, so that the gamma rays lose only a small amount of energy.
These gamma rays may have an energy just below that of the photopeak so
that their spectrum merges with the photopeak. The photopeak for a gamma
camera imaging a patient therefore contains information from
spatially-correlated, unattenuated gamma rays (which is the information
we want) and from spatially-uncorrelated, scattered gamma rays. The
scattered gamma rays act like a variable background within the true
photopeak data and the effect is that of a background haze in gamma
camera images.

!Essential elements of a modern gamma camera. MCA: Multi-Channel
Analyzer{width="400"}

While scatter may not be a significant problem in planar scintigraphy,
it has a strong bearing on the fidelity of quantitative information
derived from gamma camera images and is a vital consideration for
accurate image reconstruction in emission tomography. It is the
unattenuated gamma rays (also called the
`<b>`{=html}primary`</b>`{=html} radiation) that contain the desired
information, because of their direct dependence on radioactivity.

The scatter situation is illustrated in more detail in the figure below,
which shows estimates of the primary and scatter spectra for ^99m^Tc in
patient imaging conditions. Such spectral estimates can be generated
using Monte Carlo methods. It is seen
in the figure that the energy of the scattered radiation forms a broad
band, similar to the Compton
Smear
described previously, which merges into and contributes substantially to
the detected photopeak. The detected photopeak is therefore an
overestimate of the primary radiation. The extent of this overestimate
is likely to be dependent on the specific imaging situation because of
the different thicknesses of tissues involved. It is clear however that
the scatter contribution within the detected photopeak needs to be
accounted for if an accurate measure of radioactivity is required.

!Detected gamma ray energy spectrum for ^99m^Tc (green) with estimates
of the scatter (blue) and primary (red)
components. with estimates of the scatter (blue) and primary (red) components."){width="400"}

One method of compensating for the scatter contribution is illustrated
in the figure below and involves using data from a lower energy window
as an estimate for subtraction from the photopeak, i.e.

```{=html}
<div class="center">
```
Primary Counts =Photopeak Window Counts- k (Scatter Window Counts)

```{=html}
</div>
```
where k is a scaling factor to account for the extent of the scatter
contribution. This approach to scatter compensation is referred to as
the `<b>`{=html}Dual-Energy Window`</b>`{=html} (DEW) method. It can be
implemented in practice by acquiring two images, one for each energy
window, and subtracting a fraction (k) of the scatter image from the
photopeak image.\
\
For the spectrum shown above, it can be seen that the scaling factor, k,
is about 0.5, but it should be appreciated that its exact value is
dependent on the scattering conditions. Gamma cameras which use the DEW
method therefore generally provide the capability of adjusting k for
different imaging situations. Some systems use a narrower scatter window
than that illustrated, e.g. 114-126 keV, with a consequent increase in k
to about 1.0, for instance.\
\
A host of other methods of scatter compensation have also been
developed. These include more complex forms of energy analysis such as
the `<b>`{=html}Dual-Photopeak`</b>`{=html} and the
`<b>`{=html}Triple-Energy Window`</b>`{=html} techniques, as well as
approaches based on deconvolution and models of photon attenuation. An
excellent review of these developments is provided in Zaidi & Koral
(2004).

!Gamma ray energy spectrum for ^99m^Tc, with energy discrimination
settings of 92-126 keV for scatter estimation (blue) and of 126-154 keV,
centred on 140 keV, for the photopeak
(red). and of 126-154 keV, centred on 140 keV, for the photopeak (red)."){width="400"}

Some photographs of gamma cameras and related devices are shown below:

  ----------------------------------------------------------------------------------------------------------------------------------------------------------- ----------------------------------------------------------------------------------------------------------------- ---------------------------------------------------------------------------------------------------------------------------------------- -------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------
  !A single-headed gamma camera.{width="160"}                                                              !Another single-headed gamma camera.{width="160"}   !The NaI crystal of a gamma camera.{width="160"}                          !The cathode ray oscilloscope (CRO) of a gamma camera. of a gamma camera."){width="160"}
  !The image processing system of a gamma camera.{width="160"}   !A dual-headed gamma camera.{width="160"}            !Another view of a dual-headed gamma camera.{width="160"}   !The image acquisition and processing console of a dual-headed gamma camera.{width="160"}
  ----------------------------------------------------------------------------------------------------------------------------------------------------------- ----------------------------------------------------------------------------------------------------------------- ---------------------------------------------------------------------------------------------------------------------------------------- -------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------

We will continue with our description of the gamma camera by considering
the construction and purpose of the collimator.

## Collimation

The collimator is a device which is attached
to the front of the gamma camera head. It functions something like a
lens used in a photographic camera but this analogy is not quite correct
because it is rather difficult to focus gamma-rays. Nevertheless in its
simplest form it is used to block out all gamma rays which are heading
towards the crystal except those which are travelling at right angles to
the plane of the crystal:

!Diagram of parallel-hole collimator attached to a crystal of a gamma
camera. Obliquely incident gamma-rays are absorbed by the
septa.{width="360"}

The figure illustrates a magnified view of a **parallel-hole
collimator** attached to a crystal. The collimator simply consists of a
large number of small holes drilled in a lead plate. Notice that
gamma-rays entering at an angle to the crystal get absorbed by the lead
and that only those entering along the direction of the holes get
through to cause scintillations in the crystal. If the collimator was
not in place these obliquely incident gamma-rays would blur the images
produced by the gamma camera. In other words the images would not be
very clear.

Most gamma cameras have a number of collimators which can be fitted
depending on the examination. The basic design of these collimators is
the same except that they vary in terms of the diameter of each hole,
the depth of each hole and the thickness of lead between each hole
(commonly called the **septum thickness**). The choice of a specific
collimator is dependent on the amount of radiation absorption that
occurs (which influences the **sensitivity** of the gamma camera), and
the clarity of images (that is the **spatial resolution**) it produces.
Unfortunately these two factors are inversely related in that the use of
a collimator which produces images of good spatial resolution generally
implies that the instrument is not very sensitive to radiation.

Other collimator designs beside the parallel hole type are also in use.
For example a **diverging** hole collimator produces a minified image
and **converging** hole and **pin-hole** collimators produce a magnified
image. The pin-hole collimator is illustrated in the following figure:

!Diagram of a pin-hole collimator illustrating the inversion of
acquired
images.{width="360"}

It is typically a cone-shaped device with its walls made from lead. A
cross-section through this cone is shown in the figure. It operates in a
similar fashion to a pin-hole photographic
camera and produces an inverted image of
an object -- an arrow is used in the figure to illustrate this
inversion. This type of collimator has been found useful for imaging
small objects such as the thyroid gland.

## Example Images

A representative selection of nuclear medicine images is shown below:

  ----------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- ---------------------------------------------------------------------------------------------------------------------------------- ------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- -------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- ------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------
  !A SPECT slice of the distribution of ^99m^Tc Ceretec within a patient\'s brain.{width="120"}                                                                                !A SPECT slice through a patient\'s liver.{width="120"}   !Images from a patient\'s bone scan.{width="120"}                                                                                !A PET slice of a patient\'s brain, with a region of interest drawn to indicate the skin surface.{width="120"}   !Images from a ventilation (V) and perfusion (Q) scan of a patient\'s lungs. and perfusion (Q) scan of a patient's lungs."){width="120"}
  !A series of planar images acquired every 10 seconds during a renogram of a patient with a stone blocking their right kidney.{width="120"}   !Selected images from a renogram series.{width="120"}      !A graphical display showing the number of counts in each kidney versus time for a renogram.{width="120"}   !A SPECT slice of a patient\'s heart.{width="120"}                                                                                !A blood pool study covering the whole body of a patient.{width="120"}
  !A series from a SPECT study of a patient\'s brain.{width="120"}                                                                                !Images from a SPECT study of a patient\'s heart.{width="120"}    !A thyroid uptake study.{width="120"}                                                                                !A gastric-emptying study evaluating a patient\'s digestive system.{width="120"}                                                                  !A ^201^Tl study of the whole body of a patient.{width="120"}
  ----------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- ---------------------------------------------------------------------------------------------------------------------------------- ------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- -------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------- ------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------------

## Emission Tomography

The form of imaging which we have been describing is called **Planar
Imaging**. It produces a two-dimensional image of a three-dimensional
object. As a result images contain no depth information and some details
can be superimposed on top of each other and obscured or partially
obscured as a result. Note that this is also a feature of conventional
X-ray imaging.

The usual way of trying to overcome this limitation is to take at least
two views of the patient, one from the front and one from the side for
example. So in chest radiography a posterio-anterior (PA) and a lateral
view can be taken. And in a nuclear medicine liver scan an
antero-posterior (AP) and lateral scan are acquired.

This limitation of planar X-ray imaging was overcome by the development
of the CAT Scanner about 1970 or thereabouts.
CAT stands for Computerized Axial Tomography or Computer Assisted
Tomography and today the term is often shortened to Computed Tomography
or CT scanning (the term tomography comes from the Greek word
`<i>`{=html}tomos`</i>`{=html} meaning **slice**). Irrespective of its
exact name the technique allows images of slices through the body to be
produced using a computer. It does this in essence by taking X-ray
images at a number of angles around the patient. These slice images show
the third dimension which is missing from planar images and thus
eliminate the problem of superimposed details. Furthermore images of a
number of successive slices through a region of the patient can be
stacked on top of each other using the computer to produce a
three-dimensional
image_reconstruction "wikilink").
Clearly CT scanning is a very powerful imaging technique relative to
planar imaging.

The equivalent nuclear medicine imaging technique is called **Emission
Computed Tomography**. We will consider two implementations of this
technique below.

#### Single Photon Emission Computed Tomography (SPECT)

:   This
    SPECT
    technique uses a gamma camera to record images at a series of angles
    around the patient. These images are then subjected to a form of
    digital image processing called Image
    Reconstruction in order to
    compute images of slices through the patient.

```{=html}
<!-- -->
```

:   The Back Projection
    reconstruction process is illustrated below. Let us assume for
    simplicity that the slice through the patient actually consists of a
    2x2 voxel array with the radioactivity in each voxel given by
    A1\...A4:

!Illustration of the acquisition of four projections around the
patient,
P1\...P4{width="320"}

:   The first projection, P1, is imaged from the right and the second
    projection, P2, from the right oblique and so on. The back
    projection process involves firstly adding the projections to each
    other as shown below:

!Illustration of the back projection computation
process.{width="320"}

:   and then normalising the summed (or superimposed) projections to
    generate an estimate of the radioactivity in each voxel. Since this
    process can generate streaking artefacts in reconstructed images,
    the projections are generally filtered prior to back projection, as
    described in a later
    chapter,
    with the overall process referred to as Filtered Back Projection
    (FBP):

!Illustration of the filtered back projection computation
process.{width="320"}

:   An alternative image reconstruction technique is called Iterative
    Reconstruction, a successive
    approximation technique. The Maximum-Likelihood
    Expectation-Maximisation (ML-EM) algorithm is widely applied where a
    division process is used to compare the actual and estimated
    projections, as shown below:

!Illustration of the Maximum-Likelihood Expectation-Maximisation
(ML-EM)
algorithm. algorithm."){width="480"}

:   One cycle of data through this processing chain is referred to as
    one `<b>`{=html}iteration`</b>`{=html}. Sixteen or more iterations
    can be required in order to generate an adequate reconstruction and,
    as a result, computation times can be rather long. The
    Ordered-Subsets
    Expectation-Maximisation
    (OS-EM) algorithm can be used to substantially reduce the
    computation time by utilising a limited number of projections
    (called `<b>`{=html}subsets`</b>`{=html}) in a sequential fashion
    within the iterative process. Noise generated during the
    reconstruction process can be reduced, for example, using a Gaussian
    filter built into the reconstruction calculations or applied as a
    post-filter:

!Illustration of an Iterative Reconstruction
process.{width="480"}

:   Images generated using different iterations, subsets and filtration
    settings can be found in an online
    book.

```{=html}
<!-- -->
```

:   A comparison of these image reconstruction techniques is shown below
    for a slice through a ventilation scan of a patient\'s lungs:

![](fbpiRComp.jpg "fbpiRComp.jpg"){width="480"}

:   The gamma camera is typiclly rotated around the patient in order to
    acquire the images. Modern gamma cameras which are designed
    specifically for SPECT scanning can consist of two camera heads
    mounted parallel to each other with the patient in between. The time
    required to produce images is therefore reduced by a factor of about
    two. In addition some SPECT gamma cameras designed for brain
    scanning have three camera heads mounted in a triangular
    arrangement.

```{=html}
<!-- -->
```

:   A wide variety of strategies can be used for the acquisition and
    processing of SPECT images.

#### Positron Emission Tomography (PET)

:   You will remember from chapter
    2
    that positrons can be emitted from radioactive nuclei which have too
    many neutrons for stability. You will also remember that positrons
    do not last for very long in matter since they will quickly
    encounter an electron and a process called **annihilation** results.
    In the process the positron and electron vanish and their energy is
    converted into two gamma-rays which are emitted at roughly 180^o^
    degrees to each other. The emission is often referred to as two
    **back-to-back** gamma-rays and they each have a discrete energy of
    0.51 MeV.

```{=html}
<!-- -->
```

:   So if we administer a positron-emitting radiopharmaceutical to a
    patient an emitted positrons can annihilate with a nearby electron
    and two gamma-rays will be emitted in opposite directions. These
    gamma-rays can be detected using a ring of radiation detectors
    encircling the patient and tomographic images can be generated using
    a computer system. The detectors are typically specialised
    scintillation devices which are optimised for detection of the 0.51
    MeV gamma-rays. This ring of detectors, associated apparatus and
    computer system are called a PET
    Scanner:

![](petDiag2.jpg "petDiag2.jpg"){width="320"}

:   The locations of positron decays within the patient are highlighted
    by the solid circles in the above diagram. In addition only a few
    detectors are shown in the diagram for reasons of clarity. Each
    detector around the ring is operated in coincidence with a bank of
    opposing detectors and the annihilation gamma-rays thus detected are
    used to build up a single profile.

```{=html}
<!-- -->
```

:   It has also been found that gamma cameras fitted with thick crystals
    and special collimators can be used for PET scanning.

```{=html}
<!-- -->
```

:   The radioisotopes used for PET scanning include ^11^C, ^13^N, ^15^O
    and ^18^F. These isotopes are usually produced using an instrument
    called a cyclotron. In addition these
    isotopes have relatively short half lives. PET scanning therefore
    needs a cyclotron and associated radiopharmaceutical production
    facilities located close by. We will consider cyclotrons in the
    next
    chapter
    of this wikibook.

```{=html}
<!-- -->
```

:   Standardized Uptake Value (**SUV**) is a semi-quantitative index
    used in PET to express the uptake of a radiopharmaceutical in a
    region of interest of a patient\'s scan. Its typically calculated as
    the ratio of the radioactivity in the region to the injected dose,
    corrected for body weight. It should be noted that the SUV is
    influenced by several major sources of variability and it therefore
    should not be used as a quantitative measure.

```{=html}
<!-- -->
```

:   A number of photographs of a PET scanner are shown below:

  -------------------------------------------------------------------------------------------------------------------------------------------- ----------------------------------------------------------------------------------------------------------------------------------------------------------------------
  !The detectors and associated electronic circuitry.{width="200"}   !The scanner itself -- the detectors are under the covering panel.{width="200"}
  !Another view of the detectors.{width="200"}                                           !The image processing computer.{width="200"}
  -------------------------------------------------------------------------------------------------------------------------------------------- ----------------------------------------------------------------------------------------------------------------------------------------------------------------------

Images reconstructed using different settings of subsets and iterations
for iterative reconstruction[^2] are shown below:
!centre\|frame\|Images reconstructed using different subsets/iterations
settings within an iterative reconstruction algorithm (VPFXS). GF:
Gaussian
filtered. GF: Gaussian filtered")

## References

```{=html}
<references />
```
## External links

-   Centre for Positron Emission
    Tomography at the Austin &
    Repatriation Medical Centre, Melbourne with sections on what PET is,
    current facilities, projects & research and a PET image library

[^1]: Slomka PJ, Patton JA, Berman DS & Germano G, 2009. Advances in
    technical aspects of myocardial perfusion SPECT
    imaging. Journal of
    Nuclear Cardiology, 16(2), 255--76.

[^2]: Maher KP, 2016. Iterations, Subsets & Gaussian Filtration, 3rd
    Edition
    (Bookemon.com\]