[Cat 1] Describe and contrast the various geometries used for CT scanning.
[Cat 1] Discuss differences between single slice and multi slice CT, sequential vs helical.
[Cat 1] Define the CT-numbers.
[Cat 1] Discuss the quality of CT images in terms of resolution and noise, highlighting factors that affect each.
[Cat 1] Describe the origin and appearance of common artifacts in CT images.
[Cat 1] Discuss radiation dose features unique to CT scanning.
[Cat 1] Explain in generic terms how tube current modulation works and its impact on patient dose.
[Cat 1] Explain the impact of multi detector CT vs. single slice CT on patient dose.
[Cat 1] Distinguish between collimated X-Ray beam width and imaged sliced width.
Introduction
Computed tomography (CT) is a tomographic imaging technique that generates two-dimensional cross-sectional images in the axial plane.
Attenuation
CT images are maps of the relative linear attenuation values of tissues. The relative attenuation coefficient (μ) is normally expressed in Hounsfield units (HU), which are also known as CT numbers.
CT number represents the average linear attenuation coefficient in the voxel and is therefore dependent on:
- Tissue heterogeneity
- Variation in the attenuation coefficient of each tissue relative to water
- kV
- Filtration of the X-ray beam
It is calculated using the formula
CTn = 1000 × (μt - μw) /μw
where μt = attention coefficients of tissue μw are the attention coefficients of
tissue and water, respectively
μ is a constant for a homogeneous material and monochromatic x‐rays and is a measure of the total interaction probability per unit length of material
By definition, the HU value for water is always zero.
-HU value for air is −1,000 (μair is negligible compared with μwater).
-Table 8.1 lists typical HU values for a range of tissues.
-Because (μx and μwater are dependent on photon energy (keV), HU values depend
on the kilovolt peak and filtration.
-Therefore, HU values generated by a CT scanner are only approximate and related
to the tube voltage used to generate the image.
Chest CT : 8 mSv
Normal Chest X‐ray : 0.02 mSv
X‐ray images, in general, are obtained by
measuring the intensity of an X‐ray beam that has
been attenuated by the patient
• Attenuation is a reduction in beam intensity
– In X‐ray imaging, this is due to the interaction of the
X‐ray photons with the object (i.e. person) it passes
through
Attenuation
Remember μ is constant for monoenergetic X-rays
A typical average CT x-ray energy is around 60 keV. At ~ 75 keV, Compton interactions account for the majority of attenuation. CT contrast is therefore mainly from physical properties that influence Compton scatter:
- Physical density (dominant role)
- Electron density
CT Number
Components
X-ray tube
The CT X-ray tube requires:
- high voltage potential (80 – 140 kV)
- high anode heat capacity and dissipation
Filters
- Flat filtration: leads to higher peripheral doses, unequal detector does (central rays would be noiser
- Bow-tie filter: additional filtrates attenuating off-axis x-rays to achieve a more uniform signal at detectors
Collimator
Pre-patient collimator – reduces amount of tissue exposed
Post-patient collimator – reduces scatter reaching detector
In plane post-patient collimatory – further reduces scatter reaching detector
Detectors
- measures incident photon intensity
- Charge Integration –
- Photon counting – potential for higher spatial resoltuion
Requirements
High dynamic range (50-500 mA, 0.5 – 2 s rotations, 0.5 – 5 mm slice width -> intensity varies by
a factor of 400
– High efficiency (quantum, geometric)
– Low electronic noise
Scintillation
• Converts X-rays into light then light into electric signal (typically voltage level)
– Scintillator Crystal (e.g. Ceramic) coupled to a photodetector (e.g. Semi conductor/photomultiplier tube)
– High efficiency (~100%), can be easily made into arrays
Direct Conversion (photon counting)
• Converts X-rays directly into electric signal
– Individual photons and their energy can be detected
– Susceptible to pulse pile-up (solve with smaller pixels?)
Detectors are arranged in an arc around the
patient (to minimise inverse square law)
• Many detector channels in a row, typically 1
mm centre to centre
• Fan angle (typically ~60 degrees) defines the
field of view (FOV)
Smaller fan angle gives a smaller FOV – relevant
for dual source scanners
X-ray beam width is defined by collimation
Imaged slice width is defined by detector width/binning