In nuclear medicine the region under study becomes an active source of radiation. The imaging
problem is then to determine the three dimensional source. Radioactive materials are administered
selectively to create the source. The source of radioactivity must be less intense than in
radiography using x-rays. This is because in radiography the patient is only exposed during the
time the external source is active, which is a fraction of a second in the case of an x-ray shadow
graph to a few seconds in the case of a CT scan. In nuclear medicine the patient is irradiated from
the time the dose of radioactive material is administered until it is eliminated from the body. A
second reason for keeping the source at a low intensity is due to the collimation necessary and the
type of detectors used. While it is the source that is being imaged and not the attenuation
coefficients of intervening tissues that are being measured, it would seem that a high energy source
would be ideal. However it is then difficult to collimate the beam and to detect it as the photons
pass straight through the collimator and the detector. Photons must be absorbed by the detector to
Radioactive materials used in nuclear medicine are gamma ray emitters. I131 was the earliest
material used as it was selectively absorbed by the thyroid. Tc99, an isotope of technicium, is now
the most widely used source. It can be made easily without requiring a cyclotron. It has a short
half life of six hours, so that it does not irradiate the patient for a long period. Its gamma ray
emission energy of 140 kev is a good compromise between body attenuation and source imaging.
It can be attached to a sufficiently wide variety of compounds that can be selectively taken up by
different body organs. The most commonly used radionuclides in nuclear medicine are shown in
Radionuclides Commonly Used in Nuclear Medicine
Radionuclide Half-life Transition Eγ (MeV) Production
Carbon-11 20.38 min β+ 0.511 Cyclotron
Fluorine-18 109.77 min β+ 0.511 Cyclotron
Phosphorus-32 14.29 days β -
Chromium-51 27.70 days EC 0.320 Reactor
Cobalt-57 270.90 days EC - Cyclotron
Gallium-67 78.26 hr EC - Cyclotron
Molybdenum-99 66.00 hr β +
Technetium-99m 6.02 hr IT 0.141 Generator
Indium-111 2.83 days EC - Cyclotron
Indium-113 100.00 min IT 0.393 Cyclotron
Iodine-123 13.20 hr EC - Cyclotron
Iodine-125 60.14 days EC 0.027 Reactor
Iodine-131 8.04 days β- 0.364 Reactor
Xenon-133 5.24 days β -
Thallium-201 3.04 days EC - Cyclotron
Earliest detectors consisted of a single crystal of a scintillating material, like sodium iodide, at one
end of a long narrow diameter hole drilled in a lead block (Figure 228). The lead block collimated
the gamma ray beam. Behind the detecting crystal was a photomultiplier tube which picked up the
light generated in the crystal by bombardment with the gamma rays. The detector was scanned
over the area of interest to produce an image.
Figure 228 Scanning gamma ray detector.
Depending on the organ being imaged, different degrees of gamma ray intensity have different
implications. In the thyroid gland in the neck, areas of high radioactivity also are areas of
abnormally high tissue activity and vice versa. Tumours are usually inactive and so form "cold
nodules" in the image. Brain tumours tend to take up high levels of radioactive material. Liver will
take up material, but tumours inside the liver will not take up radioactive substances.
Scanning with a single detector is slow. The image is affected by patient movement and it is not
possible to do any dynamic studies where the change in function of an organ is measured with
time. Gamma cameras have been developed which can view an entire region simultaneously. The
speed with which an image can be formed is governed by the efficiency of the camera and the
maximum radioactive dose that can be safely given to the patient.
The dose of radiation is the Curie (Ci), which is defined as 3.7 x 1010 disintegrations per second. In
most procedures a dose of 5 Ci is given. This is usually much greater than the dose given by an x-
ray radiographic procedure.
Gamma Ray Cameras
Gamma ray cameras allow all points of a two dimensional image to be formed simultaneously
without requiring a mechanical scan. They are the most commonly used instruments in nuclear
medicine. A basic camera is shown in Figure 229. The camera consists of a collimator of some
type to convert the essentially random ray distribution into a two dimensional image. The
collimator has a similar function to the lens on a light camera. Behind the collimator is a detector
for detecting the position of each gamma ray photon. The output of the detector goes to a
recorder. The detector material, as in radiography, needs to be sufficiently thick to stop high
energy photons. Typically 10 to 12 mm thick sodium iodide crystal are used. The recorder is often
an array of photomultipliers. More modern machines use diode arrays as recorders or specially
doped arrays of diodes which act as a combined detector and recorder. The output of the diode
arrays or photomultipliers are digitised and displayed or stored.
Figure 229 Basic gamma camera.
One problem is that many of the gamma ray photons do not come directly from the organ to be
imaged but are subject to Compton scattering en route. These lead to distortions in the image.
They have lower energy than the direct photons so can be removed by thresholding the output of
the recorders. A pulse height analyser, which forms a histogram of the intensities of the photons,
can be used to set the threshold level.
This camera, named after its inventor, allows a higher resolution than the spacing between detector
elements, as applies with the ordinary gamma camera. Light from a gamma ray photon is emitted
in many directions and can strike a number of recorders simultaneously (Figure 230). By
interpolating between all the photodetectors that are recording simultaneously according to their
output amplitudes, the position of the original gamma photon in the detecting crystal can be found.
Thus in a system consisting of two photodetectors, the position of an event (gamma ray photon
strike) x is:
x = I1 x1 + I2 x2 (81)
I1 + I2
If two or more gamma rays strike the detector crystal simultaneously the event is misinterpreted.
Fortunately this is a rare occurrence. Resolution is typically 2 mm for an array size of 400 mm (ie
about 40,000 pixels) using 25 photomultipliers.
Figure 230 Anger camera principle.
A typical system is shown in Figure 231. A single detecting crystal is used. If the output pulses
from the photomultipliers do not fall within the energy range for a direct gamma ray photon for the
particular isotope used, the screen is not illuminated at that point. Centroid calculations are used to
determine the x and y positions by a generalisation of equation (81).
Figure 231 Components of an Anger camera.
The simplest collimator is a simple pinhole aperture which works in the same way as a light
pinhole camera (Figure 232). The image is magnified by M:
M = id (810)
This has the same problems as the equivalent light camera. The intensity of photons available to
form the image is small. The magnification varies inversely as the distance of the object from the
camera. If the organ imaged is a known distance away there is no problem. If it is at an unknown
depth it is difficult to determine the size of lesions. A pinhole system and thyroid image obtained is
shown in Figure 233.
Figure 233 Thyroid image and pinhole collimator scanner.
The parallel hole collimator shown in Figure 234 is placed as closely as possible to the organ to
be imaged. Each hole in the array allows the portion of the source near it to illuminate the
corresponding part of the image detector. This produces a 1:1 source to image magnification
regardless of the distance of the source from the camera. There is some distortion of the image due
to photons travelling obliquely through the holes. Although the signal to noise ratio is higher than
with the pinhole collimator, it is still too low in many applications.
Coded aperture systems are used to overcome the low signal to noise of the pinhole and parallel
hole collimators. An array of holes, known as a coded aperture plate, allows a greater number of
photons to reach the detectors (Figure 235). The resultant image is the convolution of the source
with the aperture plate. The original source can be reconstructed by a suitable cross correlation.
Figure 235 Imaging with a coded aperture.
An example of a scanner and the output of a whole body scan is shown in Figure 236.
Figure 236 Commercial scanner and typical whole body image.
Positron Emission Tomography (PET) Scanners
Some radioactive substances emit positrons which immediately interact with electrons, annihilating
both particles and generating two gamma rays, each having energies of 510 kev and travelling in
almost opposite directions. A simple imaging system can be constructed as shown in Figure 237. If
the image depth z is known, the x and y coordinates (at right angles to the plane of the paper and
also in the plane of the paper at right angles to the z direction) can be calculated:
x = x1 z2 + x2 z1
z1 + z2 z1 + z2
y = y1 z2 + y2 z1
z1 + z2 z1 + z2
The position indicating detectors only record rays if both are picked up simultaneously. This
avoids errors due to external radiation or Compton scattering. The detectors can be set to
selectively respond to photons of around 510 kev. The disadvantage of this approach is that if
there is radioactive material in other xy planes (at different z values) there will be blurring of the
Figure 237 Positron emission imaging system.
An alternate approach is similar to projection measurements made in x-ray CT scans. A ring of
detectors surrounds the patient as shown in Figure 238. Only coincident events are recorded
(Figure 239). If the number of coincident events reaching a pair of detector locations x1, y1 and x2,
y2 is summed, the result is the line integral of all sources along the line joining the two detectors.
When the line integrals are found over all angles and all positions, the source distribution can be
reconstructed in the same way as in x-ray CT scans.
There are two limitations to positron emission tomography. The positron can move a significant
distance from its point of creation to its point of annihilation. The two photons
may not be 1800 apart, depending on the momentum of the positron. Some attempt to improve the
resolution of the image is being made by measuring the small difference in times of "coincident"
events. As in ultrasonic imaging, this can be used to measure the depth at which the annihilation
took place. Due to the pico second accuracy needed in the timing measurements only a window
can be placed on the annihilation position. Nevertheless some improvement in image resolution can
be gained. Resolution is approximately 1 cm. A practical difficulty is that many of the required
isotopes require an on-site cyclotron to produce them.
Although technetium is the most commonly used radionuclide, a number of others are used in
many common applications (Table 2). The choice of a radionuclide for medical image depends on
four main factors:
1. Emission of gamma rays or positrons
2. Chemical compounds must be biologically convenient for administration
3. Chemical compounds must be physiologically selective and relevant
4. Radiation dose should be as short as possible, ie short half-life and/or rapid excretion.
Although short half-life reduces the radiation dose, it can be inconvenient in that it does not allow
enough time to transport the radionuclide from a reactor to the site where it is being used. To
overcome this problem some specialist hospitals have an on-site cyclotron to produce the
radionuclides. Cyclotrons are smaller, cheaper and safer than a reactor.
Frequently Used Procedures in Nuclear Medicine
Procedure Radionuclide Chemical form
Plasma/blood volume I, I Iodinated serum albumin
Red cell mass & life Cr Haemoglobin
Vitamin B12 absorption Co Cyanocobalamin (B12)
Thyroid function I, I Sodium iodide
Tumour & abscess Ga Citrated complex
Thrombus In Oxine labelled platelets
Cerebro-spinal fluid In DTPA chelated complex
Myocardium Tl Thallous chloride
Cardiac blood pool Tc Blood
Brain Tc Pertechnetate
Kidney Tc DTPA chelated complex
Bone Tc Methylene diphosphonate
Liver/spleen Tc Colloid
Liver/gall bladder Tc Iminodiacetic acid
Lung perfusion Tc, O Albumin, gas
Lung ventilation Xe Gas
Thyroid therapy I Sodium iodide
Polycythaemia rubra vera tmt P Sodium phosphate
Figure 238 Positron ring detector.
Figure 239 Coincidence counter in a positron imaging system.