Mammography
Physical Principles and Instrumentation
Lecture Notes III
IT Garbett
This section of work considers the instrumentation/engineering aspects
of mammography.
• X Ray tube Design
– Anode & focal spot sizes
– cathode
– tube port/filtration
• X Ray Generator & AEC
• Image Receptors
– Film/Screen
– Digital
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X Ray Tube Design
(i)Anode
Anode Material:
This has already been well covered earlier.
Mo most common
(ii) Anode Design
• Rotating anode.
• Anode angles from 10o to −9o .
Tube tilt is used to get enough field coverage for very small angles.
Tilt is in cathode-anode axis with cathode nearest to chest wall.
Use this orientation because anode-heel effect =⇒ anode ‘end’ of beam
has smallest intensity.
• Metal Insert with earthed anode
Use metal insert around portion of tube instead of glass.
Enables higher ratings because deposited cathode filament atoms on
a glass wall cause a third electrode for stray electrons from the beam.
However glass susceptible to damage from the kinetic energy of these
electrons. Metal more robust.
In addition by setting anode at the same potential (0V) as the metal
insert, electrons rebounding from focal area of target (=anode) are as
likely to go to the metal insert as back to anode. This reduces so called
‘extra focal X radiation’.
• Focal Spot sizes.
The effective (projected) focal spot size is a function of the distance
along the cathode-anode axis.
Largest spot sizes occur nearest cathode.
Actual focal spot iszes are 0.3 to 0.4 mm for contact imaging. 0.1 to
0.15 mm for magnification imaging.
More image detail occurs at the end of the image where the effective
focal spot is smallest. (Where is this then?)
(ii)Cathode
The cathode has a single element electrically isolated from the cathode cup.
The cathode cup is held at negative potential with respect to the filament
when the small focal spot size is selected (eg -400 V applied).
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Figure 1: mammography system design
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Figure 2: X ray tube tilt. Source: Bushberg et al:Essential Physics of
medical Imaging
Figure 3: Mammography tube. Source: Thompson et al:Principles of Imaging Science & Protection
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Figure 4: focal spot size. Source: Bushberg et al: Essential Physics of
medical Imaging
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Figure 5: Cathode cup bias. Source: IT Garbett
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Figure 6: space charge limited operation. Source: ITGarbett
The cathode filament is much closer to the anode in a mammography
tube than in a regular X ray tube =⇒ obtain a given mA value at much
lower filament temperatures.
=⇒ less deposition of tungsten (filament metal) onto the inner surface
of the tube insert =⇒ less attenuation of the beam.
This is important because mammography spectrum effective energy is
<< than regular X ray spectrum, and beam attenuation would be too great.
Compensation circuits are required to adjust the filament current. This
is because the mammography tube operates at low kV and so operates in
the space charge limited region.
In order to acquire the ‘set’ tube current the filament current must be
altered (by the ‘compenstaion circuits’) accordingly.
Thus tube current is a function of filament current in mammography. (Tube
current is essentially independent of filament current in normal radiography.)
Typical values 100mA for large (0.3mm) f spot and 25mA for small (0.1mm)
f spot.
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Figure 7: Tube diaphragm and Be window Source: Forster Equipment for
Diagnostic Radiography
(iii) Tube port/filtration
Already considered filtration material choice with target material.
Typical HVT1 value for 30kVp beam is ∼ 0.3 to 0.4 mm Al.
This value dpeends also on filtration used and the composition of the
compression paddle.
Radation dose requires accurate assessement of the HVT value.
Worn out tubes will ↑ or ↓ the HVT value ? (Choose one)
A beryllium window is used < attenuating than glass.
An additional metal diaphragm is used inside the tube that restricts
emission of any extra-focal X radiation.
X Ray Generator & AEC system
(i) The Generator:
1φ, 3φ or high frequency.
High frequency becoming standard (enables 1φ power supply, low ripple,
fast response time and stability in use.
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Figure 8: AEC mammography system Source:Source: Bushberg et al: Essential Physics of medical Imaging
(ii) AEC system:
Also called photo-timer. AEC system comprised of:
• Detector
Single ionization chamber or array of three semiconductor-diodes. These
detect the X radiation transmitted through the breast, grid (if used),
and image receptor.
• Amplifier
This amplifies the current from the detector that is then used to charge
up the ...
• Capacitor.
When the capacitor charges up the voltage across it builds. This voltage signal goes to the...
• Comparator. This produces an output signal when the capacitor voltage reaches the value of the other signal provided to the comparatorthe reference voltage.
The output from the comparator causes the ...
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• Termination circuit to switch off the power to the X ray tube.
The reference voltage used by the comparator is set by density adjustment at the control panel. This provides 5 to 10 steps to increase or
decrease exposure. Generally ∼ 10 to 15 % of exposure per step.
A backup timer ends exposure if insufficient exposure occurs beyond a
certain time.
If this occurs adjustment of the kVp and not the AEC density control
must be effected.
Why?
If the AEC density control is DEcreased then this MAY cause termination
of the exposure without the backup timer operating...however the density
of the radiograph will be underexposed.
If the AEC density control is INcreased then the required exposure for
termination is set even higher, and yet again the back-up timer will operate.
By increasing the kVp the beam penetration increases and allows enough
ionization at the detector site and so enough optical density on a film/screen
radiograph.
Image Detectors
(i) Film/Screen
Cassettes low attenuation carbon fibre.
Gadolinium oxysulphide (Gd2 O2 S) is mostly used for the fluorescent
screen.
Green light produced =⇒ green sensitive film.(green dye)
Conventional screen/film system for mammography uses a single emulsion film with a single screen.
In this case the film is situated proximal to the X ray tube than the
screen with the emulsion layer adjacent the fluorescent screen.
Most interactions (absorptions) of X photons occur at the top of
the screen-near the emulsion.
This means that MOST of the interactions are near the active film emulsion and so the generated visible light photons have less travel path to the
film.
A smaller path of travel =⇒ ↓ lateral spread of light and ↑ spatial resolution.
Double screen and double emulsion film systems exist also (see
figure).
Back screen thickness is same as the single screen system.
Front screen thickness is 12 thickness of the back screen.
Using this design, both screens contribute the same amount of light to
their adjacent emulsions.
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Figure 9: Film/screen system Source:Source: Bushberg et al: Essential
Physics of medical Imaging
Also ∼ equal light blurring (from lateral spreading) occurs in both films.
There is an anti-crossover barrier (film base material) used to stop visible
light crossing to the opposite side film region. (Why do this ?)
The QDE for this 2-screen 2-film system is about double that of the
single screen single film system.
The trade off is worse resolution. (But becomes acceptable in magnification imaging).
There is a clear trade off between spatial resolution and sensitivity (aka
‘speed’), when using fluoresent screens.
In addition the non-linear response of film to (visible) photon flux also
Figure 10: Film/screen system resolution. Source: Smith, A.P Fundamentals Of Digital mammography in Radiology Management Sep/Oct 2003
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Figure 11: Film/screen dynamic range. Source: Smith, A.P Fundamentals
Of Digital mammography in Radiology Management Sep/Oct 2003
means that there is a very small tolerance to the exposure values that can
be used to show contrast. (‘small exposure latitude’)
(ii) Digital
Full field digital mammography (FFDM) employs an area beam detector.
There are basically two designs of the ‘flat panel’ detectors used in
FFDM.
• Indirect Conversion
• Direct Conversion
(a) Indirect Conversion (IC)
Earliest design.
A two-step process.
step 1: X photons absorbed within fluorescent material CsI(Tl).
=⇒ visible light photons generated.
step 2: v light photons interact with photo-diodes.
These diodes arranged in an array.
Each diode produces a brief flow of current in proportion to v light energy
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Figure 12: Indirect Conversion CsI/TFT & Resolution. Source: Smith, A.P
Fundamentals Of Digital mammography in Radiology Management Sep/Oct
2003
incident upon the diode⇒ transistor ⇒ current amplification ⇒ charge built
up and then stored (eg capacitor on the output side of transistor circuit).
Effectively there are two overlying layers of arrays:
The photo-diode array & the thin film transistor (TFT) array.
These are both deposited onto a glass substrate.
A further layer of transistors involved with ‘read-out’ is also used in
addition to the ‘charge collector’ array of transistors.
Unfortunately, IC designs ⇒ light spread problem (like film/screen systems). Thicker fluorescent screens =⇒ ⇓ spatial resolution
(To some degree the use of CsI (long thin needle crystals) reduces this
effect)
However, a problem occurs compared to screen/film:
The photo-diode & transistor arrays are NOT transparent to X photons,
unlike film, =⇒ cannot put the ‘array layer’ proximal to the X ray tube.
Instead it must be situated at exit surface of the fluorescent layer.=⇒ ⇓
spatial resolution compared to film/screen system.
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Figure 13: Direct Conversion. Source: Smith, A.P Fundamentals Of Digital
mammography in Radiology Management Sep/Oct 2003
(b) Direct Conversion (DC)
Direct Conversion (DC) employs a single-step process in production of
the image signals.
Uses a photo-conductive layer.
This ‘photoconductor’ produces electron-hole pairs when X-photons are
absorbed by the photoconductor material.
An Electric field is applied across the photoconductor.
Electrons and holes then move in direction of the field.
→
−
Electrons or holes (depends on direction of the applied E field) move
toward an array electrode (each electrode ≡ a pixel). This charge then
stored on an electrode for that pixel area.
→
−
NOTE: Due to movement along the axis of the E field, there is no lateral
spreading of the electron or hole in its trajectory
=⇒ very narrow point spread function (PSF ' 1 micrometre).
Because of the action of the applied electric field, the thickness of the
photoconductor can be freely chosen.
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Figure 14: Direct Conversion and resolution. Source: Smith, A.P Fundamentals Of Digital mammography in Radiology Management Sep/Oct 2003
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µ
plot Source: Smith, A.P Fundamentals Of Digital
ρ
mammography in Radiology Management Sep/Oct 2003
Figure 15: Selenium
In practise, thickness is chosen to stop 95% of the incident X ray beam
energy.
The photoconductor used is
amorphous SELENIUM.
The virtue of this material (apart from a well established manufacturing
history in zeroradiography) is its attenuation characteristics.
The K-absorption edge is well placed in relation to the photon energies
used in mammographic spectra.
The TFT array for DC detector systems does not require the addition
of the photo-ddiode array used in the IC detectors.
This enables the same manufacturing process well established in large
area TFT/LCD arrays, commonly used in computer displays.
( It should be noted that an alternative to TFT arrays for DC detectors
is the use of charge coupled devices (CCDs), however these will not be
discussed here and also suffer from light spread problems).
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Pixel Size and Resolution
Pixel size and IC detectors:
The smaller the pixel size (≡ array detector element) then the greater the
ability of the image to display items that are separate as actually separate
(ie the greater the spatial resolution).
However, IF the information presented TO the detector array itself does
not contain two separate signals, then there is no way that the image will
display separation, even if it actually exists within the irradiated structure.
This will be the case no matter how SMALL the pixel size.
It is because of the light scattering in the fluorescent screen that the IC
detectors
are ‘scintillator’ limited in their resolution.
=⇒ there is no reason to reduce the pixel size below a certain value for IC
detectors.
This limiting value ' 100 micrometres.
This is just at the limit of ‘use’ in mammography. (Calcifications of 100
micrometres need to be resolved).
Pixel size and DC detectors:
Because of the lack of lateral spread of the generated electron-holes in
their trajectory to the detector panel, the resolution for DC detectors is
limited by the size of the pixel.
If pixel size is reduced to very small sizes then eventually a size is reached
that may be inefficient in regard to the incident photon flux (ie a large percentage of the pixel elements may have no incident photons ! ). (Quantum
limited)
At least one manufacturer(Hologic) uses a pixel size of 70 micrometres.
There is, inevitably, a practical limit (as well as a quantum limit) on the
use of very small pixel sizes. If one halves the pixel size (side of a square)
then the number of individual stored numbers increases by fourfold. Both
data storage requirements and network bandwidth needed for image distribution becomes more costly.
Intrinsic Efficiency
Intrinsic efficiency is the ratio of
energy absorbed in the detector
energy incident upon the detector
Screens used in film/screen mammography =⇒ 50% to 70%.
IC detectors using CsI(Tl) =⇒ 50% to 80%.
As discussed earlier the use of Se photoconductor material =⇒ 95%
intrinsic detector efficiency for DC detectors.
Unlike conventional screen/film mammography, the optimum effective
photon energy has not been fully established for digital mammography.
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Figure 16: absorption percentage. Source: Smith, A.P Fundamentals Of
Digital mammography in Radiology Management Sep/Oct 2003
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Detectors with a high intrinsic efficiency have potential to allow the use
of higher effective photon energy (why?)
This potential arises in both IC and DC detector designs.(see percentage
absorption graph)
DQE in digital mammography
The DQE is a measure of the transfer of the signal to noise ratio within
the system.
Recall (from study of fluoroscopy) that:
DQE =
2
SN Rout
2
SN Rin
In fact the DQE varies with the spatial frequency. (ie DQE is a function
of the spatial frequency).
This arises because in fact both the output noise and the output signal
are dependent upon the spatial frequency
(Recall spatial frequency is the rate of change of a structure with distance,
thus 2 line pairs per mm is a frequency of 2 mm−1 and yields:
low exposure value → high exposure value →low exposure value → high
exposure value
across one millimetre).
If this ’variation’ in radiation distribution is not reproduced well at the
output, there is a consequent low signal and (supposing for simplicity the
noise is constant) this means a low SN R out value.
This would indeed be the case where the signal cannot be discerned (eg
from two edges of two structures) because of light blurring (as happens in
both film/screen and also IC digital detectors).
Consequently we expect the DQE variation with spatial frequency to be
poorer in the case of both screen film and also IC detectors.
Assessment of Image Quality
A standard quality control procedure for accreditation of mamography
centres employs the use of a standard imaging phantom.
The resultatnt images need to match minimum values for certain characteristics.
One such phantom is the ACR (American College of Radiologists) phantom.
This perspex phantom is made equivalent to a 4.2 cm compressed breast
consisting of a 50% glandular 50% adipose mix of tissue.
It contains fine and coarse structures (fibres, specks and masses) that range
from visible through to invisible on a standard mamographic film image.
Spatial resolution checks are routinely made using a line pair test object.
For an ideal digital system the pixel size would determine the spatial
resolution.
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Figure 17: DQE vs spatial frequency. Source: Smith, A.P Fundamentals Of
Digital mammography in Radiology Management Sep/Oct 2003
Figure 18: ACR Mammography Phantom. Source: Smith, A.P Fundamentals Of Digital mammography in Radiology Management Sep/Oct 2003
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Figure 19: ‘Line Pair’ Test Object. Source: Smith, A.P Fundamentals Of
Digital mammography in Radiology Management Sep/Oct 2003
For example a pixel size (side of a square) of 100 micrometres means that
effectively one lead strip of 100 micrometres can occupy one pixel width, and
the gap of 100 micrometres (before the next strip) can occupy fully the adjacent pixel. This means 1 line pair occupies 200 micrometres.
Thus the limit of resolution for this system would be 5 line pairs per millimtre.
However, as previously mentioned, IC digital detectors have a spatial
resolution imposed ultimately by the light spread in the fluorescent material. Pixel size is only a good predictor of spatial resolution before this limit
is reached (100 micrometres).
Detector Size and Field of View.
For film/screen there are two cassette sizes used: 18cm x 24 cm and
24cm x 30 cm.
Larger size can be used with larger breast.
Use of larger cassette with smaller breast wasteful of film and processing
materials.
Use of smaller cassette size with larger breast =⇒ need two overlapping
irradiations to ensure coverage of the breast contents. =⇒ ⇑ absorbed dose,
multiple compressions, and is time consuming.
For digital receptors, these flat panel detectors are very costly and also
heavy and unwieldy to change =⇒ use just one detector.
=⇒ use a large detector.
The large digital detector does not have same problems as film/cassette
because the image can be cropped and there are no processing materials.
There is also a bonus in using the larger detector area for the smaller
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breast if magnification mode is used:
This means that a pixel on the magnified image represents a smaller
actual breast dimension.
Suppose a breast that ‘fits’ a 18cm x24 cm using contact imaging, is
projected onto a larger 24cm x 30 cm detector using magnification technique.
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30 × 24
=
Then area magnification, am =
18 × 24
3
3
This means that a detector element represents now = 0.6 of the object
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area it represented with no magnification.
Suppose a detector element area (pixel area) is actually
70 micrometres x 70 micrometres.
Then the magnified pixel represents
70 × 70 × 0.6 = 2940 (square micrometres) of the breast.
This
√ is equivalent to a pixel that is of size (along one side of its square
area): 2940 = 54.221767µm ' 54.2µm
Therefore imaging a ‘regular size’ breast using magnification with a large
area detector is in fact equivalent to using a detector in contact mode that
has pixel dimensions of 54 µm.
This represents a benefit in image resolution for ' 80% of women.
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