PET Imaging

PET is a form of tomographic nuclear imaging.

However, PET relies on simultaneous detection of the pair of gamma photons that are released from annihilation of a positron and an electron.

Annihilation

1. Positron decay

A positron (represented as e+, β+ or e) is released:

• Antiparticle of the electron (e-)
• Has the same mass and magnitude of charge except that the charge is positive.

Radionuclides:

• Decay via positron emission
• Proton-rich (have larger number of protons > neutrons)
• Produced in a cyclotron.

The energy difference between the parent and daughter nuclei must exceed 1.022 MeV (2 x 0.511 MeV) for positron decay to occur.

2. Positron travel through matter

1. As it travels, it collides with atoms losing energy and causing ionisation (main method of radiation dose deposition in patient)

2. As it collides, it is deflected - tortuous path with length depends upon the number of collisions and the starting energy of the positron.

This means that the distance between where the positron is emitted and where it annihilates is variable

3. Annihilation

• Shortly after positron production (travels about 0.6 - 2mm) - it annihilate with an electron -converted into two photons with energy of 511 keV travel in opposite directions (180o from each other) (final momentum is zero)
• Only occurs when the emitted positron has lost all its kinetic energy (ie has come to rest)

PET radiopharmaceuticals

• Most commonly used radionuclide - fluorine-18 (18F) … half‐life 110 min
• Most common pharmaceutical label - fluoro-deoxy-glucose (FDG)
• FDG is a tracer for glucose metabolism

NB: Other useful radionuclides are 68Ga (68min) and 82Rh (1 min), and these two are produced by radionuclide generators

PET scanner

Blocks of scintillation crystals (detector blocks) are arranged in a circle mounted on a gantry in one or two rows (10000–20000 detectors for each PET scanner).

Scintillation crystal

• NAI scintillation crystal used in SPECT and planar imaging not suitable for PET as

LAC not enough for the annihilation photons which have a higher energy of 511 keV

• Material used:

- Bismuth Germanate (BGO) - Most commonly used scintillator in PET imaging, but the light output and light decay time are inferior to NaI

- Newer materials (lutetium oxyorthosilicate (LSO) / gadolinium oxyorthosillicate (GSO))

• Each scintillation detector block viewed by 4 photomultilipler tubes
• Block of crystal subdivided by cutting smaller blocks into the scintillation crystal (“called detector elements”) and placing a reflective material in the slits to prevent cross-talk of the light photons between the elements

Ideal properties of PET scintillation detectors:

• High effective atomic number and physical density.
• High ratio of photoelectric to Compton interactions
• High detection efficiency for the 511 keV photons
• High relative light output (i.e. number of light photons produced per gamma photon absorbed)
• Good energy resolution
• Very short scintillation decay time

Forming an Image

• As annihilation produces two gamma photons that travel in opposite directions; this is used to determine which photons should be used to form the image.
• Two opposite detector elements must simultaneously detect a gamma photon (to within 1 nanosecond) for those photons to contribute to the image.
• The simultaneous gamma photon detection by opposite detector elements is called coincidence and the line between the two detector elements is called the line of response.
• The detector elements also encode the total energy deposited by the gamma photons.
• Any pulses that do not coincide in time are ignored by the electronics, as are any single photons of background radiation

N.B. this is how collimation is achieved in PET. The lead collimator grids that are used in planar imaging and SPECT are not required.

1) Data acquisition

2D vs 3D acquisition:

• In 2D acquisition only coincidences confined to a single slice of the patient are used to form the image.
• This is achieved by using a collimator ring made of tungsten (highly attenuating) to reject photons that reach the detector at an angle and, therefore, are likely to have originated from a different slice of the patient.

• In 3D acquisition no collimator is used and coincidences from a much greater volume of tissue are accepted.
• This method enables a higher total count rate due to more coincidences by allowed to reach the detector.
• It is useful with relative little scatter / administered radiation such as in brain or paediatric imaging.

NB:

• Some scanners can retract the septa and operate in either modes
• Both 2D & 3D data acquisition modes can be used to produce 3D images
• PET effective dose to the patient is the same as routine gamma imaging (short t½ of positron emitters compensate for the beta energy deposition)

Unwanted coincidence rejection:

Unwanted coincidences cause artefactual lines of response to be calculated which do not correspond to the true location of the annihilations.

Cause:

Increased scatter coincidence, which occurs with:

• 3D acquisition
• More material to travel through (e.g. body vs brain imaging)
• Solution: Energy discrimination. However, the photopeak window is wide due to poor energy resolution of the scintillators so scatter coincidence not eliminated

Increased random coincidence, which occurs with:

• 3D acquisition
• Increased administered radioactivity
• Increased duration of coincidence window

How to reduce random and scatter events

Narrow lead or tungsten septa (1 mm thick , 10 mm deep) are used between each ring of detectors (in Z plane) i.e. act as antiscatter grid

2) Data correction

Attenuation correction

Attenuation is greater in PET > SPECT due to longer path the photon must travel through the patient.

Correction:

1. Assume cross-sectional shape and uniform LAC of tissue at 511 keV

2. Measure the transmission of 511 keV photons through the patient for each line of response. A radioactive rod source (gallium-68) that gives rise to annihilation radiation is rotated around inside the detector gantry without the patient and then with the patient. This allows a calculation to be made correcting for the attenuation by the patient.

Normalisation

Individual detector elements differ in dimensions and fraction of scintillation light photons that reach the PMTs -Same radiation source may not produce same response in every detector element.

Solution:

Rotating rod source used without object in the scanner to calculate the correction factor required for differences in the individual detector elements

Correction factor

= measured counts for line of response / average counts for all lines

Dead time (Temporal resolution)

• Following the detection of a photon, a detector element cannot detect another photon for a period of time (dead time)= 1 μs
• Results in loss of counts especially in 3D scanning

NB: → if two flashes of light overlap and they are treated as one.

Solution:

Dead time measured and mathematical algorithms that take into account detector behaviour applied to extrapolate from measured counts

Radioactive decay

• Radioactivity decays as the scanner moves down the patient.
• The longer the delay from start to finish the more the radioactivity will have decayed

Solution:

Counts corrected for radioactive decay

3) Data reconstruction

2D acquisitions are reconstructed using filtered back projection or iterative reconstruction

PET-CT scanning:

• Integrated PET‐CT systems are used.
• To be able to locate and visualize this information within the patient's anatomy.
• Both detection systems are mounted on the same support, adjacent to each other, so that the single patient table moves along the central axis. Once the CT scan is complete, the patient table moves into position for the PET data collection; almost perfect matching of the images is obtained.

Advantage:

• The combined PET‐CT images are particularly useful in oncology, both for diagnosis and for accurate tumor localisation and follow‐up.

• Carefully gating image acquisition to the cardiac cycle can also produce useful fused images in cardiology.

Difficulties:

• Adjustment of matrix size • Matching the transverse planes for co-registration

• Solutions:

• Fusion software (complicated by different patient’s position and movement)

• Integrated PET CT scanners (mounted adjacent to each other)

– Perfect matching is produced

– Patient table move to acquire the CT scan, then the table return to position to acquire PET data

– CT scan can be also used for PET attenuation correction → whole scan in <30 min

– When used in cardiology: cardiac gating may be mandatory

NB: Use of gamma camera as PET scanners

Dual headed conventional gamma camera can be used as PET scanner

Conditions:

– Integration of coincidence circuity

– Rotation of the cameras around the patient without the collimators

– Thicker crystals of NaI must be used

Disadvantages:

– Higher background noise (decrease contrast)

– Poorer spatial resolution