Università degli Studi di Napoli Federico II
Scuola Politecnica e delle Scienze di Base
Dipartimento di Fisica
Corso di Laurea Triennale in Fisica
TESI DI LAUREA SPERIMENTALE
Breast computed tomography with monochromatic and
polychromatic X-ray beams: phantom studies
Relatori Candidata
Prof. Paolo Russo Veronica Corvino
Prof. Giovanni Mettivier matr. 567/533
Anno Accademico 2011/2012
1
Università degli Studi di Napoli Federico II
Scuola Politecnica e delle Scienze di Base
Dipartimento di Fisica
Corso di Laurea Triennale in Fisica
TESI DI LAUREA SPERIMENTALE
Tomografia computerizzata dedicata al seno con fasci di
raggi X monocromatici e policromatici: studi in fantoccio
Relatori Candidata
Prof. Paolo Russo Veronica Corvino
Prof. Giovanni Mettivier matr. 567/533
Anno Accademico 2011/2012
2
Dedicated to my father and my grandparents,
who cares about me and protect me
from above...
3
Contents
List of figures……………………………………………………………………………..5
List of tables……………………………………………………………………………..10
Abstract…………………………………………………………………………………..12
Introduction……………………………………………………………………………...14
Chapter 1………………………………………………………………………………...16
Screening and diagnosis for breast cancer: from 2D to 3D……………………………...16
1.1. Breast cancer…………………………………………………………………...16
1.1.1. Breast anatomy………………………………………………………..17
1.2. Breast cancer imaging: from digital mammography to digital tomosynthesis…18
1.2.1. Digital mammography………………………………………………..22
1.2.2. Digital breast tomosynthesis………………………………………….26
1.3. Cone-beam breast computed tomography (CBBCT)…………………………..30
1.3.1. The prototype scanner at University of California, Davis……………33
1.3.2. The prototype scanner at University of Rochester……………………44
1.3.3. The design of the prototype scanner at University of Erlangen,
Germany……………………………………………………………………48
1.3.4. The prototype scanner at University and INFN Napoli………………50
1.4. Breast computed tomography with synchrotron radiation……………………..57
1.4.1. The SYRMEP beamline at the ELETTRA, Trieste…………………..59
Chapter 2………………………………………………………………………………...62
Measurements with a CBBCT prototype and with a SR beam………………………….62
2.1 The experimental setup of breast CT with synchrotron radiation at Elettra,
Trieste……………………………………………………………………………62
2.1.1 CT Imaging measurements………………………………………………65
2.1.2 Image processing………………………………………………………...66
2.1.3 Dose distribution into the phantom……………………………………...83
2.2 The experimental setup for CBBCT prototype at the University of Napoli…….92
4
2.2.1 Imaging measurements…………………………………………………..95
2.2.2 Dose distribution into the phantom…………………………………….103
2.3 Comparing between SR based and CBBCT based results……………………..106
Conclusions…………………………………………………………………………….110
References……………………………………………………………………………...112
Acknowledgements…………………………………………………………………….117
5
List of figures
Fig. 1.1: Scheme of the human female breast in sagittal section: 1. Chest-wall 2. Pectorals muscles 3.
Lobules 4. Nipple 5. Areola 6. Milk duct 7. Fatty tissue 8. Skin…………………....................18
Fig. 1.2: Ten principal cancer types for the estimated new cancer cases by females, United States, 2011.
The breast cancer should represent 30% of all new cancer cases. [1]……………………........19
Fig. 1.3: Temporal trend in age-adjusted cancer death rates* among females for some chosen cancers,
United States, 1930 to 2007. [1] *Rates are age adjusted to the 2000 US standard population
†Uterus indicates uterine cervix and uterine corpus………………………………..………….20
Fig. 1.4: Relationship between 15-year-survival rate (%) and tumor size. [5]…………………………......20
Fig. 1.5: On the left, scheme of principal components of a mammography system: x-ray tube, filter,
collimator and compression paddle. On the right a real mammography machine………...…...22
Fig. 1.6: Linear attenuation coefficient of the breast tissue as a function of energy. It underlines attenuation
coefficient’s differences between the fat, glandular and cancerous tissues, infiltrating ductal
carcinoma. [14]………………………………………………………………………………...23
Fig. 1.7: FoM curve for three different tissue thickness (2.5 cm, 4.5 cm and 7.0 cm) in function of energy.
[15]…………………………………………………...…………...……………………………24
Fig. 1.8: Photo of a compressed breast during a mammographic exam………………………………….…25
Fig. 1.9: On the left, scheme of CC and MLO view. On the right, projection of a breast with digital
mammography: CC and MLO view……………………………...............................................25
Fig. 1.10: The figure on the left shows normal breast tissue while on the right, the while area circled in blue
in the tissue indicates a cancer, [From the National Cancer Institute.
http://history.nih.gov/exhibits/genetics/sect2.htm].....................................................................26
Fig. 1.11: 3D tomosynthesis with MAMMOMAT Inspiration, Siemens 2009. The X-ray tube of
MAMMOMAT Inspiration moves in a 50° arc around the breast while 25 low-dose images are
taken during the exam with a frame rate up to 2 images per second.
[www.siemens.com/healthcare]………………………………………………………………..27
Fig. 1.12: Tomosynthesis resolves overlap tissue through the shift-and-add technique. Slices perpendicular
to the z-axis can be reconstructed shifting the single projection views according to the height
and rotating the X-ray tube of different tilting [16]……………………………………………28
Fig. 1.13: The fixed detector acquired 25 images as short pulses during a continuous scan, of ± 25° relative
to the 0° position with an angle increment of 2° per image in about 20 seconds. The distance
between rotation center and detector surface is 4.7 cm and between the X-ray source and the
detector surface is about 66 cm. [16]…………………………………………………………..29
Fig. 1.14: Breast of a 76-year-old woman with a 0.9 mm intraductal infiltrating carcinoma: on the left tomo
slice and on the right 2D DM. Tumor evaluation is better with BT than DM. [16]…………...30
Fig. 1.15: A dedicated breast CT scanner: the patient is prone on a table with one pendant breast in the hole
on the table. Under the table, an x-ray tube and a flat panel detector rotate around the breast,
acquiring cone-beam projection images……………………………………………………….33
Fig. 1.16: The Albion scanner with its main components: X-ray tube, Flat Panel detector, rotanting gantry,
X-ray shutter system, chain-link cable conduit system and BCT panels. [39]………………34
Fig. 1.17: Albion scanner assembly. For radiation shielding, the bCT is surrounded by panels with an
internal layer of lead and an external aluminum support………………………………………34
Fig. 1.18: On the left, dedicated breast CT geometry – patient prone / pendant breast. On the right, half
cone-beam CT geometry view…………………………………………………………………35
Fig. 1.19: The Varian PAXSCAN 4030CB flat panel detector with a CsI scintillator. It has a native pixel
dimension of 194 µm in a 2048x1536 array but can be used in a 2x2 binning mode which
results in effective detector element size of 388 µm in a 1024x768 pixels. So it is possible 30
frames per second………………………………………………………………………….......35
Fig. 1.20: Comet X-ray tube: 640 Watt and 80 kVp to 8 mA. It has a 0.4 mm x 0.4 mm focus and it is
positioned to 47 mm from the top of the tube housing and used a window to turn on and off the
X-ray beam during the CT scan acquisition. Also has a W anode and used a water cooled
anode………………………………...........................................................................................36
Fig. 1.21: A model is shown positioned on the breast CT scanner, with her right breast positioned in the
pendant geometry and in the scanning position………………………………………………..37
6
Fig. 1.22: The Kollmorgen Servo Motor, Housed Direct Drive Rotary (DDR) D081M. It is a motion
control system which consists of the high precision bearing, angle encoder, motor and 13 ft-lb
continuous torque………………………………………………………………………………38
Fig. 1.23: At the top, Internal and external of Bodega system. Bodega’s components are similar to that
Albion scanner but this system is higher than the earlier prototype and is equipped with stairs
that allow to the technologist and the patient respectively to access the table more easily. At the
bottom, the PET hardware installed into the Bodega scanner. These PET heads consist of a 36
by 36 array of 3 mm × 3 mm × 20 mm LSO crystals, coupled to arrays of position sensitive
photomultiplier tubes. The two PET heads rotate 180° around the breast on a separate gantry
system, which in turn sits on top of the CT gantry. [39]………………………………………39
Fig. 1.24: Kollmorgen Servo Motor of Bodega scanner, Housed Direct Drive Rotary (DDR) D103M 100
ft-lb continuous torque…………………………………………………………………………39
Fig. 1.25: The geometric calibration of the scanner. The geometric calibration of the scanner is performed
by imaging a phantom consisting of a vertical row of Pb ball bearings (BB’s), in the scanner
field of view. The position of each BB is tracked over a 2π acquisition of images. The
trajectory of each BB follows an elliptical path…………………….………………………….41
Fig. 1.26: The process calibration, named “flat field correction”, makes use of the gain image and an offset
image (with no x-rays incident on the detector) acquired just prior to the acquisition. [39]…..41
Fig. 1.27: The figure shows the original image, on the left, and after HU correction, on the right………...42
Fig. 1.28: (A) The preprocessed projection image; (B) The back projection reconstruction process.
[39]..................................................................................................... .........................................43
Fig. 1.29: The spatial resolution of the breast CT scanner, as characterized by the Modulation Transfer
Function (MTF), from the center (black line) to the edge (blue line) of the scanner field of view
is reduced due to the interplay between the x-ray tube rotation around the breast and the
detector frame time of 33 ms. [39]…………………………………………………………......43
Fig. 1.30: On the left, relationship between diameter breast and compressed breast thickness. On the right,
relationship between two-view mean dose and compressed breast thickness…………………44
Fig. 1.31: This figure shows a series of breast CT images from different women, with non-contrast. These
images are all coronal sections through the breast. Is noted the difference in the characteristic
parenchyma pattern for each women. A large spiculated mass is seen in the upper left image,
with associated microcalcifications. The breast CT image on the lower right has a large field of
microcalcifications. [39]…………………………………………………………………….…45
Fig. 1.32: Cone-beam breast CT scanner: a Varian’s Rad 71SP X-ray tube and a Varian’s PaxScan 4030CB
flat panel detector mounted on a rotating assembly. Above this rotating assembly is placed a
patient table. [50]………………………………………………………………………………46
Fig. 1.33: On the left, the slice of the medium breast phantom and on the right, the slice of the large breast
phantom. This images clearly shown calcifications and tumors of different sizes. [50]………47
Fig. 1.34: Results performed on a patient show clearly two adjacent tumors. On the left a tumor of 0.27 mm
thick and on the right of 5.5 mm thick. [50]…………………………………………………...47
Fig. 1.35: Prototype scanner of Erlanger’s University. The figure shows the gantry, on which mounted X-
ray tube and detector, which can move up and down allowing spiral acquisition and more
comfortable patient’s access. Above the system placed a patient table with breast aperture.
[51]……………………………………………………………………………………………48
Fig. 1.36: On the left, water cylindrical phantoms with a tungsten wire of 10 µm diameter. On the right
homogeneous phantom with soft-tissue lesions and microcalcifications and relative
magnification. [51]……………………………………………………………………………..49
Fig. 1.37: This photo shows principal components of University and INFN prototype CBBCT. (1)
Microfocus X-ray tube; (2) High resolution Flat panel detector; (3) Rotating gantry…………50
Fig. 1.38: First European prototype (5 in USA) for Cone-Beam Breast CT/SPECT for laboratory
investigations, composed by: X-ray tube (1); flat panel detector (2); rotating gantry (3); pinhole
compact gamma camera (4); PMMA breast phantom (5). [55]………………………………..51
Fig. 1.39: Microfocus X-ray Source (Hamamatsu model L8121-03)…………………................................51
Fig. 1.40: CMOS Flat Panel Sensor model C7942CA-02 (Hamamatsu, Japan)………................................52
Fig. 1.41: On the left final version of prototype scanner. On the right shown adopted geometry: the X.ray
tube and detector, mounted on rotating gantry, rotates around the breast during
acquisition……………………………………………………………………………………...53
Fig. 1.42: CT scanner geometry: 3D coordinates system (X, Y, Z) on the scanner isocenter and 2D
coordinate system on the detector plane. [56]……………........................................................54
7
Fig. 1.43: Top and side view of the half cone-beam geometry. α1 and (α2+α3) are the fan and cone angle,
respectively. [56]……………………………………………………………………………….55
Fig. 1.44: Drawing of the breast phantom: hemi-ellipsoid of rotation on a cylindrical base, with six cavities
in its mid-plane to locate TLDs. In this figure units are in cm………………………………...56
Fig. 1.45: two halves of a breast phantom, hemi-ellipsoid on a cylindrical base, with six disk cavities to
locate six TLDs. [57]…………………………………………………………………………56
Fig. 1.46: Schematic representation of a synchrotron with the following main elements: 1) detector; 2)
injector; 3) the focusing magnet (quadrupole); 4) bending magnet (dipole); 5) cavity to radio
frequency……………………………….....................................................................................58
Fig. 1.47: Basic diagram of a synchrotron for the production of radiation...………………………….……58
Fig. 1.48: Patient bed and scanning system used at the SYRMEP beamline at Elettra, for breast
mammography and tomography with synchrotron radiation. [58]…………………………….60
Fig. 1.49: Principal components of the SYRMEP beamline at Elettra with relative distances. [58]……….61
Fig. 2.1: The experimental setup for breast CT with Synchrotron radiation……………………………….62
Fig. 2.2: Photo of the phantom 2. On the left, phantom 2 closed. You see inserts for measurements of
spatial resolution. On the right, phantom opened. You see six disk cavities for the positioning
of TLDs.………………………………………………………………………………………63
Fig. 2.3: Scheme of the phantom 2 with the size of the holes. On the left, you see six cavities for housing
TLDs: 3 places along the axis of rotation (Axtop
, Axmid
, Axbot
) and 3 along the edge (PERtop
,
PERmid
, PERbot). On the right, you see inserts for spatial resolution…………………………63
Fig. 2.4: Photo of the phantom 3 with the cylindrical inserts contained in it………………………………63
Fig. 2.5: On the left, axial scheme of the phantom 3 in which we see, in blue, the holes 8mm, 4mm, 2mm,
1mm and 0.5mm diameter. On the right scheme is a 3D plot of the phantom………………...64
Fig. 2.6: Scheme of the insert of the phantom 3 with sizes holes…………………………………………65
Fig. 2.7: Scheme of the acquisition geometry for imaging……………………………................................65
Fig. 2.8: On the left, ImageJ software available on the NIH (National Institute of Health) website,
www.nih.gov. On the right, Feldkamp’s filtered back projection reconstruction software
(COBRA by EXXIM Computing Corp. Pleasanton, CA, USA)…………………………….67
Fig. 2.9: View of the screen of Cobra main parameters……...……………………………………………..67
Fig. 2.10: Reconstructed slices of the insert B of the phantom 3 with the incident beam energy of 28 keV.
Are observed holes filled with air, egg shells fragments, olive oil and animal fat…………….70
Fig. 2.11: On the left, axial view of the insert B of the phantom 3 at 28 keV and on the right, magnified
view of the details in which egg shells fragments, animal fat and air are evident……………..71
Fig. 2.12: On the left, density linear profile along the selected line of the third slices, on the right, which
passes through the phantom holes filled with animal fat, egg shells fragments, olive oil and air
(28 keV)……………………….……………………………………………………………….71
Fig. 2.13: On the left, density linear profile along the selected line of the third slices, on the right, which
passes through the phantom holes filled with air (28 keV)………………………………….71
Fig. 2.14: Reconstructed slices of the insert B of the phantom 3 with the incident beam of 24 keV. Are
observed holes filled with air, egg shells fragments and animal fat. The hole containing animal
fat and egg shells fragments shows the so-called streaks artifacts, due to the heterogeneity of
the objects contained in it and to the difference of absorption………………………………72
Fig. 2.15: On the left, axial view of the insert B with details, of the phantom 3 at 24 keV. On the right,
magnified view of the details in which egg shells fragments, olive oil, animal fat and air are
evident………………………………………………………………………………………….73
Fig. 2.16: On the left, density linear profile along the selected line of the slice, on the right, which passes
through the phantom holes filled with animal fat, egg shells fragments, olive oil and air (24
keV)…………………………………………………………………………………………….73
Fig. 2.17: On the left, linear profile along the selected line of the slice, on the right, which passes through
the phantom holes filled with air (24 keV)………………………………………………….....73
Fig. 2.18: Reconstructed slices of the insert A of the phantom 3 with the incident beam of 20 keV. Holes
are filled with CaCO3, egg shell fragments, nylon wires and air. This slice presents streaks
artifacts very pronounced………………………………………………………………………74
Fig. 2.19: On the left, axial view of the insert A of the phantom 3 at 20 keV and on the right, magnified
view of the details in which nylon wires are evident………………………………………….74
Fig. 2.20: On the left, density linear profile along the selected line of the slice, on the right, which passes
through the phantom holes filled with egg shells fragments, CaCO3, nylon wires (20 keV)….74
8
Fig. 2.21: Trend of the values of the CNR as a function of material density: CNR increases with material
density. Points represent the nylon wire, animal fat, olive oil, CaCO3 and eggshell fragments
values, respectively…………………………………………………………………………..75
Fig. 2.22: 3D plot of the insert B of the phantom 3. It shows the different materials structure, in particular
can be observe the eggshell fragments structure and the saturated air………………………...77
Fig. 2.23: Magnified axial view of the processed image with the incident beam at 34 keV: are clearly seen
the microcalcifications inside a phantom hole…………………………………………………78
Fig. 2.24: Axial views of the processed images with the incident beam at: a) 32 keV; b) 30 keV; c) 28 keV
and d) 24 keV respectively…………………………………………………………………….78
Fig. 2.25: Sagittal views of the processed images at: a) 34 keV; b) 32 keV; c) 30 keV and d) 28 keV. In all
cases are visible the five microcalcifications inside phantom holes…………………………...79
Fig 2.26: On the left, density linear profile along a diagonal, shown on the right, containing three
microcalcifications, at 28 keV…………………………………………………………………79
Fig. 2.27: Volume viewers of the processed images at: a) 34 keV; b) 30 keV and c) 24 keV. The phantom
shape is different from the actual because the stacks corresponding to the upper part of the
phantom have not been acquired……………………………………………………………….80
Fig. 2.28: On the left, 3D graph of the intensities of pixels in a pseudo color images (non-RGB images) of
the selected ROI in the picture on the right, at 32 keV………………………………………81
Fig. 2.29: On the left, 3D graph of the intensities of pixels in a pseudo color images (non-RGB images) of
the selected ROI in the picture on the right, at 28 keV………………………………………...81
Fig. 2.30: Scheme for the measurement of absorbed dose in the phantom 2……………………………….83
Fig. 2.31: Scheme of the measurement of TLDs in air to calculate the air kerma………………………….84
Fig. 2.32: Graph of the normalized dose ratio values as a function of TLDs distance from the edge of the
phantom, in cm, for the measures of the first shift (28, 24 e 20 keV)…………………………86
Fig. 2.33: Graph of the normalized dose ratio values as a function of TLDs distance from the edge of the
phantom, in cm, for the two measures to 24 keV with a step of vertical translation of 2mm and
3mm……………………………………………………………………………………………87
Fig. 2.34: Histogram of the normalized dose ratio values as a function of energy (28, 24 and 20 keV), for
beamwidth of 2mm, for each position of the TLDs (PERtop
, PERmid
, PERbot
, Axtop
, Axmid
,
Axbot
)…………………………………………………………………………………………87
Fig. 2.35: Graph of the normalized dose ratio values as a function of TLDs distance from the edge of the
phantom, in cm, for the measures of the second shift (34, 32, 30, 28 and 24 keV)……………90
Fig. 2.36: Histogram of the normalized dose ratio values as a function of energy (34, 30, 32, 28 and 24
keV), for beamwidth of 2mm, for each position of the TLDs (PERtop
, PERmid
, PERbot
, Axtop
,
Axmid
, Axbot
)……………………………………………………………………………………91
Fig. 2.37: The plot number 1 shows the beam profile, the plot number 2 shows the same beam profile
shifted of 2mm and plot number 1+2 shows the sum of two profiles. As shown, since the beam
profile is approximately Gaussian, there is an area that is irradiated twice, i.e. a double
absorbed radiation dose………………………………………………………………………91
Fig. 2.38: The low energy setup. The X-ray tube and the flat panel detector were in a fixed position while
the breast phantom is rotated during the acquisition………………………………………..…92
Fig. 2.39: Experimental setup for CBBCT at the University of Napoli. Shown: X-ray tube (1); flat panel
detector (2); rotating gantry (3); pinhole compact gamma camera (4); PMMA breast phantom
(5). [55]………………………………………...........................................................................93
Fig. 2.40: Main screen of the software to control the x-ray tube (on the left), flat panel detector (in the
center) and the motor. [59]…………………………………......................................................93
Fig. 2.41: Scheme of the insert A and the insert B of the phantom 1, with holes of different sizes filled with
various substances……………………………………………………………………………94
Fig. 2.42: Scheme of the phantom 2 with six cylindrical cavities (12mm diameter x 1mm depth) for
housing TLDs and two sets of details…………………………….............................................95
Fig. 2.43: On the left axial and sagittal views of the phantom 2 acquired at 80 kVp; on the right linear
profile along the diameter of a 13cm axial slice of the same phantom………………………...97
Fig. 2.44: On the left axial and sagittal views of the phantom 2 acquired at 70 kVp; on the right linear
profile along the diameter of a 13cm axial slice of the same phantom………………………...97
Fig. 2.45: On the left axial and sagittal views of the phantom 2 acquired at 60 kVp; on the right linear
profile along the diameter of a 13cm axial slice of the same phantom………………………...98
Fig. 2.46: On the left axial and sagittal views of the phantom 2 acquired at 50 kVp; on the right linear
profile along the diameter of a 13cm axial slice of the same phantom………………………...98
9
Fig. 2.47: In the upper right is shown the linear profile of the insert A, along the region of interest 1,
containing details, and the bottom right the linear profile along the region of interest 2, as
shown in the image in the top left corner, at 80 kVp…………………………………………..99
Fig. 2.48: In the upper right is shown the linear profile of the insert B, along the region of interest 1,
containing details, and the bottom right the linear profile along the region of interest 2, as
shown in the image in the top left corner, at 80 kVp…………………………………………100
Fig. 2.49: On the right, axial view of the insert A of the phantom 1 at 80 kVp and on the left, magnification
of the details containing animal fat and three eggshell fragments…………………………100
Fig. 2.50: A) Coronal view of central hole containing five microcalcifications, acquired at 200 µm flat
panel pixel size, at 80 kVp and at an air kerma of 5.0 mGy. B) The same slice shown in A)
processed using a FFT band pass filter. C) Linear profile along a diagonal containing three
microcalcifications (B, A and D). The microcalcifications FWHM is also indicated………102
Fig. 2.51: A) Coronal view of central hole containing five microcalcifications, acquired at 200 µm flat
panel pixel size, at 80 kVp and at an air kerma of 7.5 mGy. B) The same slice shown in A)
processed using a FFT band pass filter. C) Linear profile along a diagonal containing three
microcalcifications (B, A and D). The microcalcifications FWHM is also indicated………102
Fig. 2.52: A) Coronal view of central hole containing five microcalcifications, acquired at 50 µm flat panel
pixel size, at 80 kVp and at an air kerma of 9.0 mGy. B) The same slice shown in A) processed
using a FFT band pass filter. C) Linear profile along a diagonal containing three
microcalcifications (B, A and D). The microcalcifications FWHM is also indicated………103
Fig. 2.53: Plot of tube output (air Kerma per mAs) as a function of the tube voltage at the isocenter of low,
on the left, and high, on the right, energy setup……………………………………………....104
Fig. 2.54: Normalized dose ratio values respect to the intermost position, Axbot
. a) For high energy setup. b)
For low energy setup…………………………………………………………….……..……..106
10
List of tables
Table 1.1: X-ray tube key specifications…………………………………………….……………………...52
Table 1.2: General ratings……………………………………………………………………………….….53
Table 2.1: Size and content of the holes of the inserts A and B of the phantom 3…………………………64
Table 2.2: Scheme of acquired and reconstructed phantoms at various energies. At 34, 32 and 30 keV has
been acquired the phantom2 with microcalcifications; at 28 and 24 keV have been acquired the
insert B of the phantom3 and the phantom2 with microcalcifications; finally at 20 keV has
been acquired the insert A of the phantom3…………………………………………………...66
Table 2.3: Scheme of internal Aluminum filters which have been used for each acquired phantom at
various energies, corresponding to the table 2.2……………………………………………….66
Table 2.4: Current variation during acquisition for phantom 3 at different energies……………………….70
Table 2.5: The FWHM (mm) and the relative error values calculated using a Gaussian fit from profiles at
28 keV………………………………………………………………………………………….72
Table 2.6: Dose mean values (expressed as air kerma, AK), in mGy, measured at various energies for
phantom 3………………………………………………………………………………………75
Table 2.7: Density mean value, standard deviation, COV (%), CNR, CNRD and SNR evaluated on CT
slices for insert B and A of the phantom 3. It is also shown the expected density value for the
different materials……………………………………………………………………………...76
Table 2.8: Dose values measured at various energies for phantom 2…………………................................82
Table 2.9: Density mean value, standard deviation, COV (%), CNR, CNRD and SNR evaluated on CT
slices for the phantom 2. The CaCO3 expected value is 2.93 g/cm3…………………………...82
Table 2.10: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the
inner position (Axbot
), of six TLDs exposed at 28 keV………...................................................84
Table 2.11: Air kerma, photon fluence, collected charge dose and normalized dose ratio, respect to the
inner position (Axbot
), of six TLDs exposed at 24 keV with aperture beam of 2mm.................85
Table 2.12: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the
inner position (Axbot
), of six TLDs exposed at 24 keV with aperture beam of 3mm………….85
Table 2.13: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the
inner position (Axbot), of six TLDs exposed at 20 keV………...................................................85
Table 2.14: TLD distances measured from the edge of the phantom…………………...............................86
Table 2.15: Normalized dose ratio values in percent, respect to Axbot
, as a function of energy (with
beamwidth of 2mm), for different TLDs positions (PERtop
, PERmid
, PERbot
, Axtop
, Axmid
,
Axbot)…………………………………………………………………………………………...87
Table 2.16: Dose, photon fluence, collected charge and normalized dose ratio, respect to the inner position
(Axbot), of six TLDs exposed at 34 keV……………………………………………………......88
Table 2.17: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the
inner position (Axbot), of six TLDs exposed at 32 keV………...................................................88
Table 2.18: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the
inner position (Axbot
), of six TLDs exposed at 30 keV………...................................................89
Table 2.19: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the
inner position (Axbot), of six TLDs exposed at 28 keV………...................................................89
Table 2.20: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the
inner position (Axbot), of six TLDs exposed at 24 keV………...................................................89
Table 2.21: Normalized dose ratio values in percent, respect to Axbot
, as a function of energy (with
beamwidth of 2mm), for different TLDs positions (PERtop
, PERmid
, PERbot
, Axtop
, Axmid
,
Axbot)…………………………………………………………………………………………..90
Table 2.22: HVL values and corresponding effective energy measured at different tube……………….…96
Table 2.23: Values of air kerma, tube load, calculated DgN and calculated MGD at various tube
voltages………………………………………………………………………………………...96
Table 2.24: FWHM values for each detail of diameter Φ, at different tube voltages, obtained by the
Gaussian fit to the line profiles of the air-filled details in fig. 2.43-2.46. There is little variation
in the detail resolution at different tube voltages………………………………………………99
Table 2.25: Detail contrast, CNR and CNRD evaluated on CT coronal slices for phantom 1……………101
Table 2.26: SNR and CNR values evaluated on the acquired images at different air kerma and with
different flat panel pixel size………………………………………………………………….101
Table 2.27: HVL values and corresponding effective energy measured at different tube voltages for high
energy setup…………………………………………………………………………………..104
11
Table 2.28: HVL values and corresponding effective energy measured at different tube voltages for low
energy setup…………………………………………………………………………………..104
Table 2.29: TLDs charge values (Q in nC), charge per air Kerma (Q/AK) and dose (QxFC) measured in the
Axbot
position for high energy setup……………………………………………………….…105
Table 2.30: TLDs charge values (Q in nC), charge per air Kerma (Q/AK) and dose (QxFC) measured in the
Axbot
position for low energy setup…………………………………………………………..105
Table 2.31: CNR data calculated with polychromatic beam (CNRCBBCT) and monochromatic beam
(CNRSR). It is evident that the details in the case of monochromatic beam have higher contrast
so they are more visible……………………………………………………………………....107
Table 2.32: Normalized dose ratio values in percent, respect to Axbot
, as a function of energy for
monochromatic beam……………………………………………………………...………….109
Table 2.33: Normalized dose ratio values in percent, respect to Axbot
, as a function of energy, for
polychromatic beam…………………………………………………………………………..109
12
Abstract
Lo scopo di questa tesi è confrontare misure di imaging su fantocci (che simulano
l’attenuazione dei tessuti e la forma del seno non compresso) ottenute mediante
tomografia computerizzata (CT, Computed Tomography) dedicata al seno con fasci di
raggi X monocromatici (radiazione di sincrotrone) e policromatici al fine di verificare la
qualità dell’immagine e la distribuzione interna di dose. L’interesse per questo lavoro
nasce dai potenziali vantaggi, in termini di qualità di immagine e di uniformità di dose
di radiazione somministrata, della diagnosi precoce del tumore con una tecnologia CT
dedicata. Infatti la CT dedicata al seno (breast CT) permette di eliminare nell’immagine
la sovrapposizione dei tessuti, che si verifica invece con la tecnica mammografica,
mediante l’acquisizione di un numero elevato di proiezioni e di una ricostruzione
tridimensionale dell’immagine, utilizzando una dose assorbita, per ogni seno,
comparabile a quelle delle due viste mammografiche.
In particolare in questo lavoro sono state elaborate misure su fantocci, per valutare la
qualità delle immagini e la distribuzione di dose, acquisite presso il SYRMEP beamline
di Trieste dal gruppo di fisica medica di questo Dipartimento. Le misure sono state
acquisite in due turni alla facility di luce di sincrotrone ELETTRA: il primo turno a
luglio 2008 ed il secondo turno a novembre 2009.
Queste misure sono state poi confrontate con le misure su fantoccio, acquisite mediante
un prototipo CT con fasci di raggi X policromatici, acquisite ed elaborate presso il
laboratorio di fisica medica dipartimentale.
La breast CT con radiazione di sincrotrone fornisce un fascio laminare, monocromatico e
regolabile in energia. Queste caratteristiche permettono di ridurre la radiazione diffusa,
aumentando la qualità dell’immagine, di rimuovere gli effetti di “beam hardening” e di
selezionare l’energia più adatta a seconda dello spessore e della composizione del seno in
esame, riducendo la dose somministrata. Tuttavia essendo il fascio laminare (dimensioni
trasverse di 120mm x 4mm), per eseguire misure di imaging e di dose il campione deve
essere traslato verticalmente in vari passi per coprire tutta la sua dimensione. Per quanto
riguarda la dosimetria sono state esaminate le letture di dosimetri TLD disposti nei
fantocci, acquisite con la stessa modalità e gli stessi parametri scelti per l’imaging.
Il prototipo breast CT con sorgente di raggi X policromatici, assemblato presso il
laboratorio di fisica medica è costituito da un tubo a raggi X e un rivelatore CsI:Tl
13
CMOS flat panel montati su un sistema gantry rotante motorizzato. La geometria usata in
questo caso per misure di fantoccio è una geometria cone-beam che permette di irradiare
completamente le dimensioni del fantoccio.
Nella tesi sono mostrate le caratteristiche e le peculiarità di entrambe le tecnologie e il
confronto in termini di qualità dell'immagine, visibilità delle microcalcificazioni, artefatti
di imaging, risoluzione spaziale, geometria di acquisizione e dose di radiazione assorbita.
14
Introduction
The aim of this thesis is to analyze and compare previous measurements on phantoms,
which simulate tissue attenuation and shape of the uncompressed breast, in a breast
computed tomography scanner with monochromatic (synchrotron radiation) and
polychromatic (radiographic tube) X-ray beams, in terms of image quality and dose
distribution. There are potential advantages of cancer early diagnosis with a CT
technology, dedicated to the scan of the breast, in terms of image quality and delivered
radiation dose. In fact, breast CT with dedicated scanners would allow to reduce the
overlap of the tissues in the image, which occurs with mammography, through the
acquisition of a large number of projections and an image three-dimensional
reconstruction. In such an imaging procedure, still at the experimental stage, the
absorbed radiation dose, for each breast, is comparable to those of the two views
mammography. The first breast CT prototype scanner has been designed at the
University of California, Davis, from the academic group led by Prof. J. Boone. Other
prototypes have been developed at the University of Rochester, Duke University,
University of Texas MD Anderson Cancer Center, and at the University of
Massachusetts, Worcester. In the European Union there are two main groups involved in
this research: the University of Erlangen, Germany, and the University of Naples and
INFN. Particularly in this work I have processed phantoms measurements to evaluate the
image quality and radiation dose distribution, acquired at the SYRMEP beamline of the
ELETTRA synchrotron radiation facilities in Trieste, by the medical physics group at the
University of Naples. The measurements were acquired in two shifts: the first in July
2008 and the second in November 2009. These measurements have been compared with
the phantoms measurements using a CT prototype with polychromatic X-ray beams,
previously acquired and processed at the Medical Physics Laboratory, University of
Naples by the same medical physics group. The breast CT with synchrotron radiation
provides a laminar, monochromatic and tunable beam. These characteristics allow to
reduce the scattered radiation, to increase the image quality, to remove the beam
hardening effects and to select the most suitable X-ray energy depending on the thickness
and composition of the breast under examination, and a reduction of the dose delivered.
However, since the beam is laminar (transverse dimensions of 120mm x 4mm), for
imaging and radiation dose measurements the phantom must be translated vertically in
15
various steps to cover all its size. As regards the dosimetry will be examined reading
from thermoluminescent dosimeters (TLD) acquired using the same method and the
same parameters chosen for imaging.
The breast CT prototype with polychromatic X-ray source, assembled at the Laboratory
of Medical Physics, University of Naples, consists essentially of an X-ray tube (W anode,
35-80 kVp, 0.25 mA, 50µm focal spot size) and a CsI: Tl CMOS flat panel detector (12 x
12 cm2 area, 50μm pitch, up to 9 fps at 4 x 4 binning), mounted on a motorized
translating and rotating gantry. The geometry used for phantom measurements is a cone-
beam geometry which allows to completely irradiate the phantom size.
This thesis is organized as follows:
- In the first chapter there is an overview of the state of the art of mammography,
tomosynthesis and in particular of the various prototypes CT for screening and
diagnosis of breast cancer developed by various academic groups.
- In the second chapter are described the two experimental setups, Synchrotron
Radiation Breast CT (SRBCT) and Cone-Beam Breast CT (CBBCT), and their
imaging and radiation dose measurements in phantoms. Also are shown the
processed data and the comparison between SR based and CBBCT based results.
16
Chapter 1
Screening and diagnosis for breast cancer: from 2D to 3D
1.1 Breast cancer
The breast cancer is the most common form of cancer among women worldwide. It is
estimated that approximately one in eight women will develop breast cancer during her
lifetime. The principal means to reduce the mortality rate are screening and diagnostic
tools. The technique presently used for detecting breast cancer is digital mammography
(DM), described in the paragraph 1.2.1. It is typically used for two purposes: for medical
controls of apparently healthy women, without symptoms (called screening
mammography) to detect any cancer in its early stage, and to aid in the diagnosis of a
woman who has symptoms, palpable lesions or suspicious finding identified by screening
mammography (called diagnostic mammography). The marketing of the first system of
digital mammography was approved by the FDA (Food and Drug Administration) in the
U.S.A. in 2000 and during time it replaced the previous tool, screen-film mammography,
allowing manipulation of the digital image, and electronic archiving. Digital
mammography is a two-dimensional (2D) X-ray imaging of the compressed breast which
represents a two-dimensional projection of a three-dimensional structure, the breast
(described in paragraph 1.1.1). Hence geometrically, tissues belonging to different planes
in the breast volume result superimposed in the radiographic image so obscuring
suspected lesions. To reduce the overlap of the various anatomical structures of the
breast new radiographic techniques have been studied providing an increasing number of
mammographic views. The first technique introduced to reduce the overlap of the breast
tissue is the digital breast tomosynthesis (DBT) (described in paragraph 1.2.2). DBT is a
tool which allows to study the breast as “layered” since the X-ray tube moves around the
compressed breast over a limited angular range. The breast is viewed from many angled
images which then are superimposed: from these projections several layer (about 50) of
the breast are reconstructed digitally. Hence DBT is not a fully three-dimensional (3D)
imaging technique. A three-dimensional technique is the breast computed tomography
(BCT), with dedicated scanners in which the X-ray tube rotates 360° around the breast
while acquiring a large number of views producing a tomographic image. The term
tomography refers to a picture (graph) of a slice (tomo) of the sample. BCT thanks to its
17
ability to produce three-dimensional slices virtually, eliminates the problem of the
overlap of anatomic structures of the breast. In paragraph 1.3 I will describe one of the
chosen approaches for BCT, cone-beam breast computed tomography (CBBCT), and in
its subparagraphs the various prototypes realized by some academic groups.
1.1.1 Breast anatomy
The human breast is an organ placed in the anterior region of the chest wall and rests on
two muscular structures: a more external, pectorals major muscle, and a deep, pectorals
minor muscle. The breast is an inhomogeneous anatomic structure composed of layers of
different types of tissue, among which predominate two types, adipose tissue and
glandular tissue. Its main elements are: the gland, located in the adipose tissue, the skin
and the nipple-areola complex (NAC) (fig. 1.1). The breast has the rough shape of a cone
with the base at the chest wall, and the apex at the nipple, the center of the nipple-areola
complex. The superficial tissue layer is separated from the skin by 0.5–2.5 cm of adipose
tissue. The Cooper’s ligaments are fibrous-tissue prolongations that radiate from the
superficial fascia to the skin envelope. Each mammary gland consists of about 15-20
glandular separated units, the lobes, which terminate with a duct at the level of the
nipple. The milk comes from the lobules to the nipple through small tubes called
lactiferous ducts. The blood vessels and lymphatic vessels are found in the stroma, the
tissue which forms the support structure, surrounding the lobules and ducts. About 75%
of the lymph travels from the breast to the axillary lymph nodes, which include the
pectoral (chest-wall), subscapular (under the scapula), and humeral (humerus-bone area),
while 25% of the lymph travels to the parasternal nodes (beside the sternum bone), to the
other breast, and to the abdominal lymph nodes. The lymphatic drainage of the breasts is
very important to oncology, because cancer cells can secede from a tumor and by means
of the lymphatic system go to other parts of the woman’s body.
Cancer refers to a set of diseases in which cells, in a part of the human body, grow in an
abnormal way. The common factor for different types of cancers is the growth of the
cells out of control. To this initial stage follows a progression, in which the abnormal
cells are able of multiplying and to move away from the original cell population. It is
thus possible to form masses and cell aggregates which interfere with the organ and the
apparatus in which they reside, even migrating to distant organs threatening the life of
the whole organism. Breast cancer can be of two types: non-invasive and invasive.
18
The non-invasive forms are:
Ductal carcinoma in situ (DCIS): it is an initial form of breast cancer limited to
the cells that form the wall of the ducts. It can become invasive if not treated.
Lobular carcinoma in situ (LCIS): it is an early form of breast cancer limited to
the area lobular.
The invasive forms are:
Infiltrating ductal carcinoma (IDC): when the tumor exceeds the wall of the duct.
It represents about 80% of all forms of breast cancers.
Infiltrating lobular carcinoma (ILC): when the tumor exceeds the wall of the
lobule. It represents about 15% of all breast cancers.
Fig. 1.1: Scheme of the human female breast in sagittal section: 1. Chest-wall 2. Pectorals muscles 3. Lobules 4. Nipple
5. Areola 6. Milk duct 7. Fatty tissue 8. Skin.
1.2 Breast cancer imaging: from digital mammography to digital
tomosynthesis
Each year, the American Cancer Society estimates the numbers of new cancer cases and
deaths expected in the United States in the current year and compiles the most recent data
on cancer incidence, mortality, and survival based on incidence data from the National
Cancer Institute (NCI), the Centers for Disease Control and Prevention, and the North
American Association of Central Cancer Registries and mortality data from the National
Center for Health Statistics (NCHS). Jemal et al. [1] reported the most common cancers
expected to occur in men and women in 2011. In particular, the three most commonly
19
diagnosed types of cancer among women in 2011 were in the breast, lung and bronchus,
and colon and rectum, representing about 53% of estimated cancer cases in women. As
shown in fig. 1.2 breast cancer alone should represent 30% (230,480) of all new cancer
cases among women in 2011. In Italy, instead, it is estimated that in 2011 the number of
new cases of breast cancer diagnosed is 29% (45,000) of all cases of cancer among
women (AIRTUM, Italian Association of Cancer Registries). Fig. 1.3 shows the annual
cancer death rates among women in USA for some types of cancers, from 1930 to 2006.
Since in this work we are interested in breast cancer, we observe that the same figure
shows the comparison between the deaths from breast cancer in the 20 years before
screening was introduced (1958-1977) with those of breast cancer diagnosed in the 20
years after the introduction of screening (1978-1997). It was found that in the past 15
years, the mortality rate for this type of cancer among women was reduced by 30%-50%
with respect to the previous period, thanks to improvements in early detection and
development of new therapy modalities [2, 3].
Fig. 1.2: Ten principal cancer types for the estimated new cancer cases by females, United States, 2011. The breast
cancer should represent 30% of all new cancer cases. [1]
Although X-ray screening mammography has saved many lives and is the most
commonly used technique for early detection of breast cancer, it has limitations.
Screening mammography has a limited sensitivity to detect breast cancer, especially in
women with “dense” breast tissue (breast with a high fibroglandular tissue content, as is
typical in younger women), where sensitivity is related to a test's ability to identify
positive results. It is also limited with respect to the tumor size that can be detected (less
than several millimeters). A breast imaging screening device must be able to detect
20
tumors at the earliest possible stage in their development: the earlier the tumor is
detected, the higher are the chances of survival after years from its detection and therapy.
Michaelson et al. [4] suggested that the relationship between the 15-year-survival
probability for patient with breast carcinoma and tumor size can be fit to a simple quite
exponential equation.
Fig. 1.3: Temporal trend in age-adjusted cancer death rates* among females for some chosen cancers, United States,
1930 to 2007. [1] *Rates are age adjusted to the 2000 US standard population †Uterus indicates uterine cervix and uterine corpus.
Fig. 1.4 [5] shows relationship based on the study reported by Tabar et al. [6], and
underlines the importance of screening devices which can detect smaller tumors. For
example, the average-size breast tumor detected by conventional mammography is on the
order of 10-12 mm in diameter [7].
Fig. 1.4: Relationship between 15-year-survival rate (%) and tumor size. [5]
21
By the exponential relationship in fig. 1.4, a new imaging technique which could detect
breast tumors of 5 mm in diameter would increase survival rates by 8%-10%. That is to
say, if considering the estimated new breast cancer cases by women, (230,480) as
reported above, the number of women with diagnosed breast cancer will decrease by
about 20,743 units annually. This has led many researchers to investigate a number of
alternative technologies for breast lesion analysis [8]. Some of these include X-ray
computed tomography (CT), positron emission tomography (PET), single-photon
emission tomography (SPECT). This work will be devoted to the study of computed
tomography dedicated to breast imaging.
The present technique used to detect the breast cancer is mammography. One of the main
problems with conventional mammography is that the recorded image represents the
projection of a three-dimensional object, the breast, on a two-dimensional plane, which
results in two projection images: a view from head to foot (cranio caudal, CC) and an
angled side view (medio lateral oblique, MLO). This produces overlapping of normal
tissue, soft tissue masses and calcifications, which makes visualization and detection of
suspected lesions difficult. In the last 10-12 years, there has been a steady trend in the
replacement of film-screen mammography with digital flat-panel detectors. Digital
mammography has been demonstrated to be more accurate than film-screen
mammography for women with denser breast tissue [9]. However despite the
improvement in screening accuracy with the use of digital mammography compared with
screen-film mammography, the overlap of breast tissue can still obscure a breast lesion
and hinder detection and/or diagnosis. A technique that has been proposed for improving
visualization of breast tissue is tomosynthesis, or limited-angle tomographic
mammography, which allows a three-dimensional image view. One of the advantages of
breast tomosynthesis is that it can be realized by an upgrade of a conventional digital
mammography system. The principles of tomosynthesis were first discussed by Ziedses
des Plantes in 1932 [10] and have been studied for use with film-screen based
radiography systems [11, 12]. Niklason et al. [13] was one of the first reports to present
promising results of imaging breast phantoms and breast specimens using a
tomosynthesis technique with a stationary, amorphous silicon flat-panel detector.
However, due to a very large blurring along the direction orthogonal to the detector
plane, tomosynthesis is not really a 3-D imaging system, as will be described in
22
paragraph 1.2.2. Another important factor as well as the imaging, is the delivered
radiation dose to the patient: since the breast is one of the most radiosensitive organs
(mainly due to the radio-sensitivity of fibroglandular tissue), the risk of cancer induced
by X-ray exposure has to be minimized. However, for all mammographic techniques a
good image is a result of a compromise between high image quality defined in terms of
low noise and artifacts, and low delivered dose to the breast.
1.2.1 Digital mammography
Digital mammography is a two-dimensional imaging technique for screening and
diagnosis of breast cancer, which uses X-ray radiation, resulting by an upgrade of screen
film mammography (fig. 1.5). Indeed in digital mammography the radiographic film is
replaced by a digital detector which absorbs X-rays transmitted through the breast and
converts their deposited energy into electronic signals.
Fig. 1.5: On the left, scheme of principal components of a mammography system: x-ray tube, filter, collimator and
compression paddle. On the right a real mammography machine.
The principal components of a mammography machine are: the X-ray tube, filter/target
combination and the compression paddle. The target material, usually molybdenum (Mo
with Z=42), rhodium (Rh with Z=45) or tungsten (W with Z=74), used to achieve
characteristic X-rays of the desired energy while filters are used to reduce the low energy
X-rays so they do not affect the patient exposure and radiation risk.
Digital mammography permits the detection of low contrast and small size details such
as masses and clusters of microcalcifications, which are possible indicators of early
23
breast cancer. Microcalcifications are small (typically from 50 µm to less than 500 μm in
diameter) calcium deposits sometimes associated with breast cancer. The visibility of
details, for example of microcalcifications, which are set in a tissue background, can be
described by the signal to noise ratio (SNRΔs). The SNRΔs, considering the Poisson
statistics of the detected photons, it is defined as:
where N1 is the average photon fluence (photons/mm2) on the image in the background
region and N2 is the average fluence in the area of the detail. The linear attenuation
coefficient of normal and cancerous tissue in the breast decreases with the energy, as
shown in fig. 1.6, so the SNR is higher at lower energies.
Fig. 1.6: Linear attenuation coefficient of the breast tissue as a function of energy. It underlines attenuation
coefficient’s differences between the fat, glandular and cancerous tissues, infiltrating ductal carcinoma. [14]
But lower energy means higher absorbed dose (at constant photon fluence) in the organ
therefore a parameter which describe the image quality versus the energy is the Figure of
Merit (FoM) defined as the signal-to-noise ratio normalized to the mean glandular dose
(MDG):
where MGD (Mean Glandular Dose, mGy) is a quantity representing the average dose
delivered to the glandular breast tissue. It is used as a parameter to evaluate the risk of
breast cancer for ionizing radiation. It is a complex quantity which cannot be directly
24
measured because it depends on various parameters such as the kVp of the X-ray tube;
target/filter combination; composition and thickness breast. However the upper limit of
the MGD delivered by a two-view mammography, for each breast, is defined in
protocols. In particular in USA this limit is 6 mGy while in Europe it is of 5 mGy for an
average compressed breast (about 4.5cm thickness) consisting of 50% glandular and 50%
adipose tissue. MGD is calculated through the following equation:
MGD = DgN x ESAK
where ESAK is the entrance skin air Kerma, (kerma is an acronym for Kinetic Energy
Released in Matter and it defined as the kinetic energy transferred to charged particles by
ionizing radiation per unit mass), in mGy and DgN (normalized glandular dose) is an air
kerma-average glandular dose conversion factor (mGy/mGy). The DgN values also vary
depending on the target/filter combination used in mammography unit, breast
composition and thickness, kVp and half-value layer, HVL, (half-value layer is the
thickness of a given material through which 50% of the incident energy is attenuated; it
is photon energy dependent and inversely proportional to the effective attenuation
coefficient). Fig. 1.7 [15] shows the maximum of the FoM curves corresponding to a
definite energy for various tissue thicknesses: at increasing thickness, the maximum
moves to higher energies and at that energy the image quality is maximized for the same
radiation dose.
Fig. 1.7: FoM curve for three different tissue thickness (2.5 cm, 4.5 cm and 7.0 cm) in function of energy. [15]
During the mammographic procedure, one breast is placed between two parallel flat
plates and compressed using a compression paddle. In addition one of these two plates
25
moves to produce a force on the breast as shown in fig. 1.8. The compression allows
uniforming the breast tissue to increase the image quality, since the reduction of the
thickness of the tissue that X-rays must penetrate decreases the amount of scattered
radiation, which produces a reduction of the image contrast. This also includes the
decrease of the required dose of radiation and motion artifacts
Fig. 1.8: Photo of a compressed breast during a mammographic exam.
Fig. 1.9: On the left, scheme of CC and MLO view. On the right, projection of a breast with digital mammography: CC
and MLO view.
26
In screening mammography, two projections are made of each breast: cranio-caudal, CC,
and medio-lateral oblique, MLO (fig. 1.9). Diagnostic mammography may include these
and other projections, including those magnified for the study and the deepening of
details. The image in digital mammography is displayed on a monitor in real time with
the reading of radiological picture via computer and can then be appropriately modified
by varying some parameters (contrast, brightness, magnification) with the result of
obtaining a correct view of each area of the breast (fig. 1.10).
Fig. 1.10: The figure on the left shows normal breast tissue while on the right, the while area circled in blue in the
tissue indicates a cancer, [From the National Cancer Institute. http://history.nih.gov/exhibits/genetics/sect2.htm].
1.2.2 Digital breast tomosynthesis
Digital breast tomosynthesis (DBT) is a three-dimensional imaging technique which
allows to reconstruct three-dimensional images of the breast from a finite number of low
dose two-dimensional projections, obtained with different tilting of the X-ray tube
assembly. The radio-geometrical principle of tomosynthesis is similar to that applied in
the old stratigraphic technique. Stratigraphy is a radiographic technique invented in the
30s by an Italian radiologist, Alessandro Vallebona, which permits visualization of only a
desired layer, according to the principles of projective geometry, with the removal of
confounding surrounding structures. However the fundamental difference between the
two techniques is that, while the stratigraphy required the acquisition of multiple
exposures for each layer that it wants to "focus", the digital tomosynthesis enables to
reconstruct an arbitrary number of planes from the same sequence of two-dimensional
projections. This is made possible by the separation between the process of acquisition
and visualization mode allowed by the use of digital detectors for which the same raw
projections can be processed to reconstruct different planes. The three-dimensional
27
reconstruction, in principle, allows to overcome one of the main limitations of two-
dimensional imaging, that is the masking of lesions (in the case of breast: masses,
microcalcifications, etc.), caused by the superposition of normal structures. So digital
tomosynthesis permits a substantial improvement in detection and in analysis of breast
lesions. As an example of DBT setups, Siemens Healthcare (one of the world's largest
player in the healthcare industry and in the medical imaging) produces a 3D
tomosynthesis system for breast imaging, Mammomat Inspiration shown in fig. 1.11,
which requires a similar radiation dose as a normal digital mammography but better
diagnostic power. The X-ray tube of MAMMOMAT Inspiration moves in a 50° arc
around the breast while 25 low-dose images are taken during the exam with a frame rate
up to 2 images per second.
The principle used in digital tomosynthesis for solving the problem of overlapping
images is projection of the same object after rotating the X-ray tube. Tomosynthesis
allows to reconstruct any plane of the object in exam, shifting the X-ray tube in the
height (long Z-axis) and to acquire one discrete set of X-ray projections through rotation
of the X-ray tube. The algorithm used to reconstruct topographic image, called shift-and-
add, is shown in fig. 1.12.
Fig. 1.11: 3D tomosynthesis with MAMMOMAT Inspiration, Siemens (2009). The X-ray tube of MAMMOMAT
Inspiration moves in a 50° arc around the breast while 25 low-dose images are taken during the exam with a frame rate
up to 2 images per second. [www.siemens.com/healthcare].
28
Fig. 1.12: Tomosynthesis resolves overlap tissue through the shift-and-add technique. Slices perpendicular to the z-axis
can be reconstructed shifting the single projection views according to the height and rotating the X-ray tube of different
tilting [16].
The projection images taken at different tilting of the X-ray tube are electronically shifted
and added. In this way the focused image plane at a certain depth under surface allows a
better visualization of the characteristic of that plane.
The parameters that influence image quality are: quality of each projection, number of
projections, angular range of the projections and also the image reconstruction algorithm.
The quality of each projection is determined by the radiation dose and the detector used.
The number of projections is limited by the performance of the detector, because the dose
applied to the patient has to be limited; in addition the number of projection views at a
given tilt should not be great to avoid the presence of streak artifacts. However to reduce
the artifacts and noise typical for breast tomosynthesis it is used a dedicated filtered back
projection (FBP) reconstruction algorithm. The filter parameter can be tuned for the
specific imaging function. The complete filtering process includes three filters: a ramp
type filter, a spectral filter and a so-called “slice-thickness filter”. The first compensates
for the blurring introduced by the back projection. The second reduces high-frequency
noise and the third filter guarantees a constant depth resolution to a certain degree.
In the particular case of Siemens’ 3D breast tomosynthesis, parameters and the exam
procedure are similar to that of a digital mammography. The breast is compressed on a
dedicated device containing a full-field digital mammography detector with the following
characteristics: high DQE direct-converting amorphous selenium (a-Se) flat panel with
29
an array of 2816 x 3584 pixels, a 85 μm pixel pitch rendering an active area of 23.9 cm x
30.5 cm and high-speed, low-noise digital images. DQE (Detective Quantum Efficiency)
is a parameter which gives an indication of detector quality evaluating the noise
introduced by the system and its spatial resolution. It defined as the ratio of the SNR2
output from the system to the SNR2
of the signal input into the system:
DQE =
The read out time of the fixed detector is optimized for digital breast tomosynthesis
imaging: 25 projections on an angular range of 50° can be acquired with full detector
resolution in about 20 seconds. Twenty-five images acquired by the detector as short
pulses during a continuous scan, of ± 25° relative to the 0° position with a angle
increment of 2° per image. The distance between rotation center and detector surface is
4.7 cm and between the X-ray source and the detector surface is about 66 cm (fig. 1.13).
These image acquisition parameters provide the optimal compromise between image
quality, dose, and field-of-view. The acquired images are displayed on a monitor and can
be elaborated to improve the detection of important markers such as speculated masses,
micro-calcifications, etc.
Fig. 1.13: The fixed detector acquired 25 images as short pulses during a continuous scan, of ± 25° relative to the 0°
position with an angle increment of 2° per image in about 20 seconds. The distance between rotation center and
detector surface is 4.7 cm and between the X-ray source and the detector surface is about 66 cm. [16]
30
Investigators [16] have found that breast tomosynthesis improves the detection of
cancerous tissue and enables a better classification of lesions with an increase of
absorbed radiation dose absolutely negligible, respect to digital mammography, as shown
in fig. 1.14. An additional advantageous feature of breast tomosynthesis is the lower
compression pressure of the breast than the procedure of digital mammography which is
cause of pain for patients.
Fig. 1.14: Breast of a 76-year-old woman with a 0.9 mm intraductal infiltrating carcinoma: on the left tomo slice and
on the right 2D DM. Tumor evaluation is better with BT than DM. [16]
1.3 Cone-beam breast Computed Tomography (CBBCT)
The first investigation on the use of breast CT is been reported by Reese et al. [17] in
1976. After promising studies performed, the General Electric (GE) Company has
constructed a prototype, dedicated breast CT scanner, called CT\M (mammography). In
October 1975 two CT\M scanner have been installed, one at the Mayo Clinic in
Rochester and another at the University of Kansas College of Health Sciences. The GE
CT\M scanner used fan-beam geometry to acquire 1 cm thick CT slices in about 10
seconds. The system composed of a GE Maxiray 75 x-ray tube and an array of 127 high-
pressure xenon gas detectors. Over this system there was a canvas table with a hole for
the breast. Women were imaged in the prone position with a breast in the opening, by
submerging the breast into a container filled with running warm water, before and after
the intravenous administration of 300 ml iodine contrast agent. The typical parameters of
imaging were: 120 kVp, 20 mA, 10s for 360° and the mid-breast dose for a six-slice (6
cm) was 1.75 mGy. Chang et al. [18] and Gisvold et al. [19] have suggested that breast
CT imaging has high sensitivity for the detection of malignant lesions but lower
specificity with false-positive results, attributed to the poor spatial resolution of the
31
scanner. Problems related to poor spatial resolution and the concern about radiation dose
led GE not to commercialize these scanners and in the late 1970s, the two prototypes, at
the Mayo Clinic and University of Kansas, were dismissed. In the next decade other
studies have been conducted to evaluate the use of a conventional whole-body CT
scanner as a diagnostic device to the breast lesions and have been achieved significant
improvements, especially in terms of spatial resolution and contrast resolution [20, 21,
22, and 23]. However the breast imaging with a whole-body CT scanner has many
problems. One is the relatively high radiation dose to the breast compared to
conventional mammography [20, 21, and 24]. For instance, in a study by Miyake et al.
[21], the radiation dose was measured at 23.5 mGy, which was almost 10 times that of
mammography. Also the x-ray travel through the entire thorax, and thus a great quantity
of non-breast tissue is exposed to radiation.
With the advent of digital flat panel detectors for mammography and other applications,
researchers have developed CT systems dedicated exclusively to breast imaging.
Dedicated breast CT systems have a number of advantages: they reduce the radiation
exposure to the non-breast tissue and improve detection of microcalcifications and tumor
margins. Another important advantage of the patient geometry proposed for dedicated
CT breast imaging is the imaging without compression. Breast compression is performed
in conventional mammography to maximize soft tissue contrast. Dullum et al. [25] have
reported that about 50% of women undergoing a mammography scan feel moderate or
greater pain. Reducing the pain level of breast imaging many women would be
encouraged to undergo regularly breast screening. One of the approaches chosen in breast
CT to ensure an excellent sensitivity and specificity is the cone-beam imaging system.
The cone-beam breast computed tomography is an X-ray imaging technique for breast
cancer diagnosis. It is a young research field that grows up in the last years since the
article by Boone et al. in 2001 [26]. The need for a dedicated CT devise arises from the
need to reduce the radiation dose to the non-target chest tissue and the cardiac and
respiratory motion artifacts that contributes to poor image quality. The added advantage
of CT imaging with respect to conventional mammography is a tomographic (3D) view
of breast lesions that allow visualizing tissue lesions separated from overlying normal
tissue structures. The overlying structures, due to the breast compression, make the
detection of small carcinomas (size a few millimeters) difficult because of occultation in
32
dense areas, causing a high rate of false-positive, which reaches percentages of at least
30-40% of all mammography. These cases mainly concern patients with breast so-called
“dense” that is breast with high fibroglandular tissue content, as is typical in younger
women. The 3D imaging would make possible to obtain three-dimensional images of the
breast, providing a more accurate diagnosis of structures and patterns of very small
lesions and eliminating the compressions also causes pain to the patients. This type of
imaging allows separating information from structures located at different depths in the
organ; in this so it would be possible to distinguish clearly different characteristics and
also remove the bottom is not uniform, which can disturb the detection of details of small
size or low contrast.
One of the significant questions in the dedicated CT breast systems is the radiation dose
provided to the breast. In order to be proposed as a dose-comparable technique to
mammography in breast cancer screening, CBBCT has to deliver a mean glandular dose
(MGD) not higher that the MGD for two-view mammography for the equivalent average
breast. The Mean Dose to the radiosensitive Glandular tissue (MGD) is the quantity
recommended by many international protocols [27, 28] as a control on the delivered dose
to the organ. The European guidelines for quality assurance in digital mammography
[28] indicated a maximum value of 2.5 mGy while the American College Radiology
practice guideline [27] for the performance of screening and diagnostic mammography
defines a maximum limit of 3 mGy, for one-view for an average compressed breast of
about 5 cm, consisting of 50% glandular tissue and 50% adipose tissue. The average
uncompressed breast diameter (to be imaged in CBBCT) as measured at the chest wall,
was found to be 14 cm [29] corresponding to 5 cm for the same compressed breast as
imaged in conventional mammography [30]. With an uncompressed breast the
requirement of a MGD of about 5 mGy combined with good image quality, can be
realized using W-anode X-ray tubes operated at typically 80 kVp tube voltage, with X-
ray tube currents in the order of a few mA, flat panel digital detectors operated at 30
frames per second (fps), scanning times in the range 10-20s [26] and optimal
magnification factors in the range 1.4-2.2 [31]. This scan time is much longer than in
mammography but short enough for operating in a breath-hold condition. The basic
design is shown in fig. 1.15, where the patient is prone on a table with one pendant breast
33
in the hole made especially on the table. Under the table, an x-ray tube and a flat panel
detector rotate around the breast, acquiring cone-beam projection images.
Fig. 1.15: A dedicated breast CT scanner: the patient is prone on a table with one pendant breast in the hole on the
table. Under the table, an x-ray tube and a flat panel detector rotate around the breast, acquiring cone-beam projection
images.
Some academic groups and new enterprises are investigating dedicated breast CT
imaging using similar systems. One of the leaders in this field is the group at University
of California (UC) at Davis, led by John Boone. Another important academic group
developing similar prototypes is at the University of Rochester, NY, led by Prof. Ruola
Ning [32]. Other academic groups include the Duke University led by Prof. Martin
Tornai [33]; University of Texas (UT) M.D. Anderson Cancer Center, led by Dr. Chris
Shaw [34]; and the University of Massachusetts School (UMASS), Worcester, led by Dr.
Stephen Glick [35]. In the European Union, an FP7 Project (“Dedicated CT of the female
breast”) led by University of Erlangen, Germany, started in January 2008 and it ended in
2010 [36], but no dedicated scanner development has been reported yet. Another
academic group is the Medical Physics group at University of Napoli and INFN where
the realization of a cone-beam breast CT started in 2007. The first clinical trial was
started by the UC Davis group in 2004 and reports have been presented recently by this
group [37] and by the University of Rochester group [38]. Some of these prototypes are
described below.
1.3.1 The prototype scanner at University of California, Davis
The UC Davis group has developed and built three prototype dedicated breast CT
systems that are been also undergoing to clinical trial, and a fourth has been designed. In
34
this paragraph are described the first two prototypes scanner, Albion and Bodega. The
third, not yet reported in the literature, is derived from an upgrade of the first two.
Design and fabrication
The first prototype scanner, named Albion (fig. 1.16-1.17), has been fabricated in 2001,
and has been tested on phantoms and then in clinical test on patients, on November
2004.
Fig. 1.16: The Albion scanner with its main components: X-ray tube, Flat Panel detector, rotanting gantry, X-ray
shutter system, chain-link cable conduit system and BCT panels. [39]
Fig. 1.17: Albion scanner assembly. For radiation shielding, the bCT is surrounded by panels with an internal layer of
lead and an external aluminum support.
35
Fig. 1.18: On the left, dedicated breast CT geometry – patient prone / pendant breast. On the right, half cone-beam CT
geometry view.
The basic geometry of the Albion scanner is a half cone beam, which requires only the
rotation of the gantry around the breast to acquire the data for reconstructing the entire
breast volume.
This system uses a Varian PAXSCAN 4030CB flat panel detector (fig. 1.19), which has
a native detector element size of 0.194 mm in a 2048 × 1536 array, resulting in a 400 mm
× 300 mm field of view in the detector plane. Since at full resolution, the frame rate of
this detector is 7.5 frames per second, which is too slow for breath hold breast CT, the
flat panel is used in a 2 × 2 binning mode, which results in effective detector element size
of 0.388 mm and reduces the matrix to 1024 × 768 pixels. In this way, a frame rate of 30
frames per second is possible. The PAXSCAN flat panel detector is an indirect system
which employs a CsI scintillator, and at the detector plane the Nyquist frequency for the
0.388 mm detector element pitch is 1.28 mm-1
.
Fig. 1.19: The Varian PAXSCAN 4030CB flat panel detector with a CsI scintillator. It has a native pixel dimension of
194 µm in a 2048x1536 array but can be used in a 2x2 binning mode which results in effective detector element size of
388 µm in a 1024x768 pixels. So it is possible 30 frames per second.
36
The Albion scanner uses a Comet industrial (Comet AG, Flamatt, Switzerland) 640 Watt
X-ray tube, which means that at 80 kVp it can run continuously at 8.0 mA, and at 64 kVp
it could run at 10 mA (fig. 1.20) and is placed about 47 cm away from isocenter of the
system. This tube has a 0.4 mm focal spot which is positioned to 47 mm from the top of
the tube housing. The form factor of the Comet X-ray tube is a stationary tungsten anode
system and therefore uses a water cooled anode, thereby requiring the coupling of water
hoses to and from the X-ray tube on the rotating frame. In addition, this X-ray tube
cannot be pulsed due to the limited focal spot loading, and thus it must be operated in
continuous mode. On the Albion scanner is been positioned a table with a hole to
accommodate the breast during the clinical exam.
Fig. 1.20: Comet X-ray tube: 640 Watt and 80 kVp to 8 mA. It has a 0.4 mm x 0.4 mm focus and it is positioned to 47
mm from the top of the tube housing and used a window to turn on and off the X-ray beam during the CT scan
acquisition. Also has a W anode and used a water cooled anode.
Because it is essential to image the entire woman’s breast tissue in a screening
examination, is essential a good image coverage towards the posterior of the breast (chest
wall). To make this both the detector and the X-ray tube are positioned as close to the
bottom of the patient table and the X-ray tube also needs to have the X-ray focal spot
positioned near the physical end of the tube housing. In this way the central ray strike the
detector near its top, resulting in the half cone geometry. This geometry permits to
produce CT images so that all of the tissue up to chest wall can be captured in the
scanner’s field of view (FOV). A model is shown in fig. 1.21 in the scanning position. It
is extremely important in a breast cancer screen to image all of the breast tissue.
Therefore, the patient table was designed to have a 5 cm depression covering the central
37
region of the tabletop. However trial and error has been required in order to study which
tabletop designs are better capable of imaging the entire breast.
Fig. 1.21: A model is shown positioned on the breast CT scanner, with her right breast positioned in the pendant
geometry and in the scanning position.
This depression allows the woman’s thoracic region to collapse into the scan plane, such
that the X-ray tube and detector can image up to the chest wall (pectorals muscle). In this
prototype scanner the flat panel detector is positioned on a gantry arm, opposite the X-
ray source. The gantry arm is mounted on a motion control system (Kollmorgen,
Radford, VA) which consists of the high precision bearing, angle encoder, and motor
(fig. 1.22). The top of the X-ray tube housing is seen in figure 1.20, with the metallic
cylinders protruding from both sides of the tube housing. These metal cylinders are high
current solenoids used to rapidly activate an X-ray shutter system, which is used to turn
on and off, under computer control, the X-ray beam during the CT scan acquisition. The
actuation time is approximately 133 ms, and thus is acceptable for activating the X-rays
before and after the 17 second scan. In addition to the moveable X-ray shutter, the X-ray
tube assembly includes a lead collimator which focuses all primary radiation onto the
rectangular detector, positioned approximately 90 cm away. With this geometry, the X-
ray detector system becomes the primary barrier for radiation protection. For additional
radiation shielding, the panels which surround the bCT gantry were fabricated to include
an internal layer of lead, laminated to the external aluminum support. In addition to the
floor and the 4 mm thick stainless steel table, these panels provide the secondary
radiation shielding to reduce the scattered radiation levels in the room during the scan.
38
Fig. 1.22: The Kollmorgen Servo Motor, Housed Direct Drive Rotary (DDR) D081M. It is a motion control system
which consists of the high precision bearing, angle encoder, motor and 13 ft-lb continuous torque.
It was thought that 17 s was a reasonable time for a woman to hold her breath, in the
prone position during the scan. However, to reduce motion artifacts (which occur when
the patient moves during the acquisition) the scan time should be reduced. To satisfy this,
a second prototype breast CT scanner, Bodega, has been built which has an acquisition
time of 9 s.
The second prototype scanner, named Bodega, is fabricated in 2006 (fig. 1.23). This
system used a 1000 Watt Comet X-ray tube, allows 12.5 mA operation at 80 kVp for
example, a Varian PAXSCAN 4030CB flat panel detector (fig. 1.19) and a motion
control system Kollmorgen (fig. 1.24). The Bodega scanner has principally the same type
of components as the first prototype however this system is higher than the earlier
prototype, allowing a more comfortable access by the mammography technologist in
positioning the breast. An additional advantage is the stairs that allow the patient to
access the table top more easily. The second prototype also has an addition degree of
freedom into its design. The X-ray tube was mounted on a vertical ball drive, allowing
vertical motion during the rotational motion of the tube. Thus, instead of the strictly
circular motion of the Albion system, Bodega is capable of any number of scan
geometries, such as a “potato chip” trajectory, circle, line, etc. These other trajectories
have the potential to overcome some of the Fourier sampling limitations of cone beam
CT. The larger dimensions of this scanner also allow the incorporation of PET detectors
(fig. 1.23). The PET component, reported by Wu et al. [40], utilizes 32 detector blocks
set in 2 square panels. Each block consists of a 9 by 9 array of 3 mm×3 mm×20 mm
39
lutetium oxyorthosilicate (LSO) crystals, coupled by a tapered fiber light guide to a
position-sensitive photomultiplier tube (R5900-C8, Hamamatsu Photonics).
Fig. 1.23: At the top, Internal and external of Bodega system. Bodega’s components are similar to that Albion scanner
but this system is higher than the earlier prototype and is equipped with stairs that allow to the technologist and the
patient respectively to access the table more easily. At the bottom, the PET hardware installed into the Bodega scanner.
These PET heads consist of a 36 by 36 array of 3 mm × 3 mm × 20 mm LSO crystals, coupled to arrays of position
sensitive photomultiplier tubes. The two PET heads rotate 180° around the breast on a separate gantry system, which in
turn sits on top of the CT gantry. [39]
Fig. 1.24: Kollmorgen Servo Motor of Bodega scanner, Housed Direct Drive Rotary (DDR) D103M 100 ft-lb
continuous torque.
40
A resistive network reduces the output of the 16 blocks in each panel to four position
signals [41]. Triggering is accomplished using standard NIM electronics while pulse
shaping is achieved using a custom-built fast spectroscopy amplifier. Data is acquired
using a PCI Power DAQ board PD2-MFS-2M/14 (United Electronic Industries, Inc.,
Walpole, MA) characterized for PET applications by Judenhofer et al. [42]. The detector
heads are mounted on a gantry that allows radial translation, vertical translation and
rotation around the breast. When not in use, the detector heads are retracted behind lead
shields to avoid exposure to X-ray flux. Data is acquired in “step and shoot” mode, and a
typical scan takes approximately 10 minutes [43]. Spatial resolution under optimal
circumstances is approximately 2.5 mm FWHM, typical energy resolution is 25% at 511
keV, and sensitivity is approximately 1.7% at the center of the field of view [40].
Calibration and reconstruction
To reconstruct the images is necessary first of all a calibration procedure. The calibration
procedure consists in:
1) Geometric calibration
2) Detector calibration
3) HU calibration
4) Reconstruction
1) Geometric calibration
The geometric calibration is performed, every day before to scanning patients, to
determinate accurately dimensions and angles of the breast CT scanner, in order that the
reconstruction algorithm reproduces the acquisition geometry of the data. The geometric
calibration of the scanner is performed by imaging a phantom consisting of a vertical row
of Pb ball bearings (BB’s), in the scanner field of view (fig. 1.25). The position of each
BB is tracked over a 2π acquisition of images and the trajectory of each BB follows an
elliptical path, as shown in figure 1.25. The calibration data is fit to a series of equations,
and five key geometric calibration parameters are determined algebraically. The
parameters include the x and y coordinate of the ray that is normal to the detector in both
the vertical and horizontal dimensions, the rotation angle θ of the detector around this
point and the source to isocenter distance (SIC), and the source to detector distance
(SID).
41
Fig. 1.25: The geometric calibration of the scanner. The geometric calibration of the scanner is performed by imaging a
phantom consisting of a vertical row of Pb ball bearings (BB’s), in the scanner field of view. The position of each BB
is tracked over a 2π acquisition of images. The trajectory of each BB follows an elliptical path.
2) Detector calibration
The flat panel detector systems have a characteristic structured noise pattern thereby is
necessary a detector calibration process. However the PAXSCAN 4030CB is widely
linear thereby not required non-linear methods. The detector calibration process consists
into perform a series of individual exposures at 15 different mA levels, from 0.2 mA to
10 mA, under computer control, with the gantry completely stopped. At each kerma rate
are acquired 300 frames and for each pixel on the image is calculated the mean gray scale
(GS) value. After determining the response of the detector for each 2x2 binned detector
elements on the entire kerma rate system, a linear regression is used to determine the
slope (units of GS/mA) and intercept (units of GS) of the linear response. This
calibration process is also a supervision technique for finding bad detector elements. The
calibration process makes use of the gain image and an offset image (with no X-rays
incident on the detector) is acquired just previously to the acquisition. As shown in fig.
1.26, the gain and offset images are used in a process called flat field correction.
Fig. 1.26: The process calibration, named “flat field correction”, makes use of the gain image and an offset image (with
no x-rays incident on the detector) acquired just prior to the acquisition. [39]
42
3) HU calibration
Computed tomography provides by means of the CT number (CT#), or Hounsfield unit
(HU), the relative electron density of tissue. HU is defined as:
HU = 1000 x ( - w
w)
where µ is the attenuation coefficient and µw is the attenuation coefficient of water.
Attenuation coefficients depend on electron density, atomic number (Z) and on the
quality of the beam used in the CT scanner. For materials with an average atomic number
similar to that of water, the HU lie on or near a straight line that passes through HU= -
1000 for air and HU= 0 for water. Bone substitutive materials give HU values that lie
above this line, with different scanners. Fig. 1.27 shown the original image and image
after HU calibration.
Fig. 1.27: The figure shows the original image, on the left, and after HU correction, on the right.
4) Reconstruction
Once the projection images are corrected, they are preprocessed before to CT
reconstruction. A 30 x 40 pixel region of interest (ROI), corresponding to I0, is identified
on the breast and is computed the log ratio of average intensity (fig. 1.28A):
µt(x, y) = ln (
)
The next phase is the filtered back projection reconstruction process, using an 80 kVp X-
ray spectrum with either 0.3 mm or 0.2 mm of Cu filtration. These corrected values are
then used in a Feldkamp [44] style filtered back projection cone beam reconstruction
algorithm to produce a high resolution ~5123 CT volume data set (fig. 1.28B).
43
Fig. 1.28: (A) The preprocessed projection image; (B) The back projection reconstruction process. [39]
Image quality
The image quality is evaluated with the Modulation Transfer Functions (MTF). MTF
describes the quality of the imaging system with respect to its spatial resolution
properties, i.e. MTF (f) describes the ability of the system to distinguish variations in the
spatial distribution of the incident photon flux. As shown in fig, 1.29 the resolution
decreases towards the periphery of the field of view, at larger radius from the isocenter,
as consequence of the combined effects of the motion of the continuous X-ray source, as
opposed to pulsed X-rays, around the breast and the 33 ms frame time of the detector.
Fig. 1.29: The spatial resolution of the breast CT scanner, as characterized by the Modulation Transfer Function
(MTF), from the center (black line) to the edge (blue line) of the scanner field of view is reduced due to the interplay
between the x-ray tube rotation around the breast and the detector frame time of 33 ms. [39]
44
Radiation dose
The methods for computing the mean glandular dose to the breast in the geometry of
breast CT, with the women prone, breast pendant, and the x-ray tube rotating 2Л around
the breast are Monte Carlo simulations [45]. The Monte Carlo analyses done for this
purpose [46] and the breast CT technique factors (kVp, filtration, and mAs), maintaining
X-ray quantum noise to a reasonable level, are adjusted to deliver a mean glandular
radiation dose to the breast comparable to that two-view mammography. In computing of
delivered radiation dose must be taken into account of a layer of skin over the breast that
from the breast images is demonstrated to be about 1.5 mm thick [47]. This implies that
current DgN values for mammography (where the X-ray energies are low) are too low. In
addition, many medical physicists assume a breast with 50% adipose and 50% glandular
tissue , for breast dosimetry. Recent observations however have revealed the fact that the
average volume glandular fractions are much smaller, with the average volume glandular
fraction falling somewhere in the 15% to 20% range. This, too, leads to higher DgN
coefficients. The consequences of the above observations are that the dose to the breast
from mammography is higher than previously thought – perhaps 50% higher. This means
that the radiation levels in breast CT could be increased from the levels that are currently
used, to be equal to two view mammography. However a relationship between diameter
breast and compressed breast thickness is shown in fig. 1.30 and in the same figure is
shown a two-view mammography dose versus compressed breast thickness so is possible
determine the delivered dose level to uncompressed breast which doesn’t overcome that
of two-view mammography.
Fig. 1.30: On the left, relationship between diameter breast and compressed breast thickness. On the right,
relationship between two-view mean dose and compressed breast thickness.
45
Patient imaging
The UC Davis’s group is been the first to perform clinical trial. The phase I started in
2004 involving 10 healthy volunteers. The phase II, started in 2005, has involved women
with diagnosed lesions. At late 2009 were performed breast CT imaging on over 220
women, with over 50 cases performed both with and without the addition of iodinated
contrast agent. The women is placed prone with the breast to be imaged hanging into the
hole at the center of the table and is encouraged to extend her breast as far as possible
into the depression in the table. During CT image acquisition, women must hold their
breath for 17 seconds (for Albion scanner). The images of breast CT (fig. 1.31) show that
breast CT does an excellent job at demonstrating mass lesions. Although
microcalcifications are clearly visible in some of these breast CT images,
microcalcifications detection performance could be improved. An improvement in the
spatial resolution comes from design of a pulsed X-ray source system. In an analysis on
69 of earlier non-contrast breast CT cases, Lindfors [48] found that breast CT is superior
to mammography for the detection of mass lesions in the breast, but mammography
remained significantly better in the representation of microcalcifications. Nevertheless,
only a significant minority of breast cancers are seen based upon microcalcifications
alone, and thus the superior mass detection of breast CT may compensate in part for its
reduced performance in microcalcifications detection performance.
Fig. 1.31: This figure shows a series of breast CT images from different women, with non-contrast. These images are
all coronal sections through the breast. Is noted the difference in the characteristic parenchyma pattern for each women.
A large spiculated mass is seen in the upper left image, with associated microcalcifications. The breast CT image on
the lower right has a large field of microcalcifications. [39]
46
1.3.2 The prototype scanner at University of Rochester
The University of Rochester group has designed and constructed a high resolution cone-
beam breast CT scanner. Also has performed a computer simulation study and a series of
phantom and specimen experiments to validate their system.
Prototype scanner
Fig. 1.32 shows this CBBCT prototype scanner. It consists of a rotating assembly on
which mounted a Varian’s Rad 71SP x-ray tube and a Varian’s PaxScan 4030CB flat
panel detector. On this rotating assembly is placed a patient table which has a hole to
permit the woman’s breast to hang during the imaging. This system has two degrees of
freedom: the continuous rotation of the gantry around the breast, thanks to a slip ring
technology, to achieve a continuous circle scan and a controlled vertical motion during
the rotation to perform a spiral scan. The system can shift along vertical axis of 20 cm
maximum with a speed up to 1 rev/s. The gantry, with X-ray tube and the flat panel
detector, during data acquisition rotates of 360°, acquiring up to max 300 projection
images at 30 fps for 10s. Then the acquired images, 768 x 1024 x 16 bits, reconstructed
using Feldkamp’s algorithm [49] with a modified Shepp-Logan filter.
Fig. 1.32: Cone-beam breast CT scanner: a Varian’s Rad 71SP X-ray tube and a Varian’s PaxScan 4030CB flat panel
detector mounted on a rotating assembly. Above this rotating assembly is placed a patient table. [50]
Phantom and patient studies
The Rochester’s group performed an initial evaluation on phantom to evaluate
performances of this prototype in terms of uniformity, noise, CT number linearity, low
contrast resolution and high contrast spatial resolution [50]. To make this they used three
different breast phantoms, cylindrical water phantoms, of three different sizes: the small
of 12cm x 10cm x 8.2cm; the medium of 18cm x 15cm x 9cm and the large of 21cm x
47
17cm x 10cm. Each phantom of uncompressed shape consists of a phantom body, made
of a resistant urethane gel to simulate breast tissue, and internal targets, of diameter from
1 mm to 10 mm, for different purpose [50].
The results of this initial studies performed on phantoms have shown that the CBBCT
scanner really removes structure overlap and the image quality is better detecting all
masses and calcifications of different size (fig. 1.33) with a glandular dose level
equivalent to that of a single two-view mammography exam of the breast.
Fig. 1.33: On the left, the slice of the medium breast phantom and on the right, the slice of the large breast phantom.
This images clearly shown calcifications and tumors of different sizes. [50]
Clinical study at the University of Rochester Medical Center are performed on two
groups of subjects. The first group is compound by volunteer, women of age upper than
40 years, with the goal to show characteristic of cone-beam breast computed tomography
respect to mammographic exam. The second group is compound by women with
abnormalities evaluated by physical exam or imaging modality with the goal to compare
CBBCT images with that of digital mammography.
The results performed on patients indicate that the CBBCT imaging system is better into
detect small breast tumors and provide a better visibility of calcifications compared with
the conventional mammography system [50] as shown in fig. 1.34.
Fig. 1.34: Results performed on a patient show clearly two adjacent tumors. On the left a tumor of 0.27 mm thick and
on the right of 5.5 mm thick. [50]
48
1.3.3 The design of the prototype scanner at University of Erlangen, Germany
The group of Medical Physics in Erlangen started to study a dedicated breast CT imaging
technique in 2006. Its goal was to construct a tool capable to meet the following
demands: full 3D views; good soft tissue and tissue density differentiation; contrast-
enhanced imaging; high spatial resolution of 100 µm; mean glandular dose of below 5
mGy; patient comfort without breast compression; biopsy integration. The breast
imaging approach chosen is a spiral CT: the table moves at a constant speed while the
gantry rotates around the patient during scan. This system uses direct converting CdTe
crystals in photon-counting mode with a detector pixel size of (100 µm)2. The detector
mounted on a gantry in front of X-ray tube voltage of 60 kV and focal spot size of 100 or
300 µm with 3 mm aluminum filtration. The gantry moving up and down allowing spiral
motion. The focus-isocenter distance is 300 mm while detector-isocenter distance is 150
mm; these parameters are very important for X-ray power and the size of the CT system.
On the gantry system mounted a patient table, which can move allowing easy patient
access, with breast aperture. The design of this prototype is shown in fig. 1.35 [51].
During scan the patient is prone on table and the gantry rotates, acquiring 2000
projection per 360°, with 2s rotation time, in a time that doesn’t exceed 10s for each
breast. The X-ray tube voltage is 60 kV while the mAs value adjusted to deliver an MGD
of 3 mGy.
Fig. 1.35: Prototype scanner of Erlanger’s University. The figure shows the gantry, on which mounted X-ray tube and
detector, which can move up and down allowing spiral acquisition and more comfortable patient’s access. Above the
system placed a patient table with breast aperture. [51]
49
Photon counting detector
Single photon counting detectors have been developed for nuclear medicine application,
such as positron emission tomography (PET) and single photon emission computed
tomography (SPECT) but some research groups have tried to develop photon counting
systems for X-ray imaging applications [52, 53]. Single Photon Counting detectors allow
counting the number of interacting photons only in energy higher than threshold energy.
Each photon interacting in the detector material and depositing an energy (or charge)
higher than a threshold level produces a hit. The possibility to set a threshold implies that
noise can be eliminated. This increases the signal-to-noise ratio and the dynamic range.
This device requires signal discrimination and counting logic in each picture element, a
complex microelectronics task for large area and high granularity detectors.
Phantom studies
The Erlangen group performed studies to verify spatial resolution and visibility of
different lesions in the breast of breast CT, using various phantoms which simulate an
average breast. These phantoms have cylindrical shape with a diameter of 14 cm and a
length of 10 cm. For evaluate spatial resolution are used water cylindrical phantoms with
inside a tungsten (W) wire of 10 µm diameter positioned 2 mm from the center. While
for evaluate the visibility of different lesions is used a homogeneous phantom with 80%
adipose tissue and 20% glandular tissue, with inside soft-tissue lesions and areas of
microcalcifications. To simulate soft-tissue lesions are chosen spheres of 1, 2 and 5 mm
in diameter while microcalcifications were made of calcium hydroxyapatite of 100, 150
and 200 µm in diameter (fig. 1.36).
Fig. 1.36: On the left, water cylindrical phantoms with a tungsten wire of 10 µm diameter. On the right homogeneous
phantom with soft-tissue lesions and microcalcifications and relative magnification. [51]
50
The acquired image reconstructed with a standard Feldkamp-based filtered back-
projection (FBP) software package with a iterative method (MBIR = model-based
iterative reconstruction) based on the ordered subset convex (OSC) algorithm [54] to
evaluate soft-tissue lesion, with a voxel size of (150 µm)3
and a smooth filter kernel, and
microcalcifications, with a voxel size of (50 µm)3 and a sharp kernel.
By phantom studies [51], it resulted that: lesions of 2 and 5 mm diameter in adipose
tissue are clearly detectable both with 100 and 300 µm focal spot size while glandular
lesions of 1 mm diameter weren’t enough detectable; microcalcifications resolved down
to 150 µm diameter while calcifications with 100 µm diameter are not visible with 300
µm focus size but detected for 100 µm focus size. All this with delivered dose levels that
not exceeds those of mammography screening.
1.3.4 The prototype scanner at University and INFN Napoli
The prototype scanner for laboratory investigations on CBBCT realized at the University
and INFN Napoli by Medical Physics group (within research projects: BREAST-CT in
2007-2008, and BCT in 2009-2011), is the first built in Europe. The cone-beam breast
CT prototype has been designed, manufactured and assembled completely in-house. It is
a non-clinical bench-top prototype composed of a step motor rotating gantry, a
microfocus W-anode X-ray tube, a flat panel detector, custom acquisition software and
cone-beam CT reconstruction software (fig. 1.37). The gantry can host a compact pinhole
gamma camera for SPECT imaging, based on a photon counting CdTe pixel detector
(fig. 1.38). The scanner is mounted on an optical bench (1.5 x 1.8 m2) and housed in a
lead-shielded (3 mm Pb) box of 2x2x2.6 m3.
Fig. 1.37: This photo shows principal components of University and INFN prototype CBBCT. (1) Microfocus X-ray
tube; (2) High resolution Flat panel detector; (3) Rotating gantry.
51
Fig. 1.38: First European prototype (5 in USA) for Cone-Beam Breast CT/SPECT for laboratory investigations,
composed by: X-ray tube (1); flat panel detector (2); rotating gantry (3); pinhole compact gamma camera (4); PMMA
breast phantom (5). [55]
The X-ray source
The source used in this study is an X-ray tube (SB-80-250, Source-Ray Inc., Bohemia,
NY), shown in fig. 1.38. The tube has a W tungsten anode and it operates in the range
35-80 kVp with a tube current up to 0.25 mA and the rise\fall time from and to standby is
less than 0.25 s. The tube also has a carbon filter window, a 50 µm minimum focal spot
and an inherent filtration of 1.8 mm Al. For providing additional filtration, copper filters
can be used.
For studies of micro breast CT with the same scanner, which will not be discussed in this
thesis, a microfocus X-ray tube (Hamamatsu model L8121-03) has been used. The X-ray
tube has a W tungsten anode and it operates in the range 40-150 kVp with a tube current
of up to 500 µA. This tube can work at different focal spot sizes, of 5, 20, 50 µm. The
focal spot of 5 µm is ensured only at the power of 4 W, otherwise the smallest focal spot
size is 7 µm (fig. 1.39 and table 1.1). The tube presents a carbon fiber window, an
inherent filtration of 1.8 mm Al and an additional filtration of 0.2 mm Cu. Its computer
control is via a serial interface.
Fig. 1.39: Microfocus X-ray Source (Hamamatsu model L8121-03).
52
Part Number L8121-03
Abstract 150 kV sealed type
Tube Voltage 40-150kV
Tube Current 10-500µA
Min. X-ray Focal Spot Size 5µm
Max Output 75W
X-ray Beam Angle 43degrees
FOD (Focus to object distance) 17mm
Table 1.1: X-ray tube key specifications
The Flat Panel Detector
The detector used is a CMOS (Complementary Metal-Oxide Semiconductor) Flat Panel
Sensor model C7942CA-02 (Hamamatsu, Japan) (fig. 1.40). It is digital X-ray image
sensor newly developed as key device for non-destructive inspection, biochemical
imaging. It is composed of a matrix of 2400 x 2400 pixels with a CsI:Tl scintillator layer
(0.2 mm thick) for X-ray indirect detection and a 1mm thick Al sheet placed at a distance
of 9mm from the scintillator surface. The Flat Panel detector has a sensitive area of 12cm
x 12cm with 50µm pixel pitch, (table 1.2). It can be operated (internal trigger) at a rate of
2 fps (1 x 1 binning), 4 fps (2 x 2 binning) or 9 fps (4 x 4 binning. This detector is a
lightweight and compact flat panel sensor consisting of a sensor board and a control
board. The sensor board also has 8 charge-sensitive amplifier arrays each having 300-
channel amplifiers with a horizontal shift register. Analog video signals are amplified as
the charge on each video line by 2400 channel charge amplifiers with CDS (Correlated
Double Sampling) circuits added, and are output each of 8 amplifier arrays. The control
board converts the analog video signal into a 12-bit digital signal and outputs it to an
external frame grabber, (IMAQ PCI-1424, National Instruments, Austin, TX, USA, RS-
422 interface), through the 12-bit parallel port; the signal digitization is at 12 bit/pixel.
Fig. 1.40: CMOS Flat Panel Sensor model C7942CA-02 (Hamamatsu, Japan).
53
Part Number C7942CA-02
Pixel size 50 x 50
Photodiode area 120 x 120
Number of pixels
Number of active pixels
2400 x 2400
2240 x 2344
Readout Charge amplifier array
Video output (Data 1-12) RS-422 (differential) 12 bit
Scintillator CsI
Table 1.2: General ratings.
The X-ray tube and flat panel detector placed on a mechanical assembly, consisting of
two rotating arms positioned along horizontal X axis. The X-ray tube and flat panel
detector located on opposing arms. The scanner comprises 8 computer-controlled step
motors: one controls the gantry rotation, other six control two arms making them move
on the three axes XYZ, and another rotator is placed behind the flat panel detector, for
aligning its row-columns axes with the vertical (Z) axis of the scanner. The arms can
rotate up to a maximum radius of 50 cm, and translating the X-ray tube and/or the flat
panel, along the horizontal axis, the system magnification can be regulated. According to
the standard scan protocol the gantry rotation speed is 6°/s with a frame rate of 7 fps.
Also used four-blades W collimator (2 mm thick) at the X-ray tube, which shapes the
beam so as to produce, e.g., a truly cone-beam, a half cone-beam or a fan beam.
Fig. 1.41: On the left final version of prototype scanner. On the right shown adopted geometry: the X.ray tube and
detector, mounted on rotating gantry, rotates around the breast during acquisition.
In final version on the mechanical assembly will be placed a patient bed with a hole for
accommodate the breast during clinical exam. The adopted geometry by system is that in
which woman is in prone position with one breast at time pending from a hole in the
54
patient bed. The X-ray tube and detector, mounted on rotating gantry, will be rotated
around the breast during acquisition allowing tomographic reconstruction according to
vertical axis (fig. 1.41). This configuration geometric-anatomic of the gantry allows to
reduce, through an opportune screen, the radiation to the chest wall of the patient.
Moreover the breast, in pendant position, grows away naturally from the chest wall
allowing a full vision of back zone.
CBBCT scanner geometry
To define scanner geometry was chosen a 3D fixed coordinate system (X, Y, Z) in which
object space is placed at the origin O of this system. The vertical axis, Z, is chosen as
rotation axis, parallel to the X-ray tube, which placed vertically with the anode on the
side of the top of the scanner. While the horizontal axis is defined as the axis which
intersects the Z axis and passes through the X-ray focal spot. According to this employed
system define the coronal, axial and sagittal plane as the X-Y, X-Z and Y-Z planes,
respectively. Also can be define another 2D coordinate system (u, v), but on the detector
plane, where u and v are parallel to the X and to the Z axis, respectively (fig. 1.42).
Fig. 1.42: CT scanner geometry: 3D coordinates system (X, Y, Z) on the scanner isocenter and 2D coordinate system
on the detector plane. [56]
55
The phantom is placed at the scanner isocenter (point in which the central axis of the
beam intersects the axis of the gantry), while the X-ray source and the flat panel detector
are placed along the X axis, at a source-to-object distance (SOD) and source-to-detector
distance (SDD), respectively. In this way it is possible to determine the image
magnification factor, as:
M =
The chosen geometry, thanks to the collimator blades at the X-ray tube, to irradiate
phantoms is half cone-beam geometry, as shown in fig. 1.43. To evaluate dose
distribution and imaging are used PMMA breast phantom with six TLDs inside.
Fig. 1.43: Top and side view of the half cone-beam geometry. α1 and (α2+α3) are the fan and cone angle, respectively.
[56]
The PMMA breast phantom
The breast phantom is an object with a shape which simulates the hanging breast of
average size in a breast CT exam. It is made of PMMA (polymethyl-methacrylate, 1.19
g/cm3): PMMA is used as a material with linear attenuation coefficient close to that of
soft tissue for diagnostic range X-rays. As shown in fig. 1.44 the breast phantom has the
geometrical form of a hemi-ellipsoid of rotation on a cylindrical base, of 14 cm diameter
which simulates the average diameter of the uncompressed breast at the chest wall [16],
shared in two-parts, having half-axes of 7 cm (breast radius at chest wall) and 9.5 cm
(breast length), and a cylindrical base of 3.5 cm thick simulating the chest wall. In its
mid-plane there are six disk cavities, 1mm depth and 12 mm diameter, to locate TLDs:
56
three on the periphery of the phantom and three on the longitudinal phantom axis in
PERtop
, PERmid
, PERbot
and AXtop
, AXmid
, AXbot
positions, respectively.
TLD
TLDs (Thermo luminescent dosimeters) are used in dosimetry measurements to evaluate
the delivered dose during exams. In this study they were chosen to measure the dose
distribution into the breast phantoms. They are of the type: TLD-100 Harshaw, USA,
from Thermo Fisher Scientific; of size: 3mm x 3mm x 0,9mm. One TLD located in each
of the six cavities of the breast phantom: AXtop
, AXmid
, AXbot
, PERtop
, PERmid
, PERbot
positions (fig. 1.45). The TLDs utilized in this study are independently calibrated and
characterized in laboratories by Medical Physics group at the University of Naples.
Fig. 1.44: Drawing of the breast phantom: hemi-ellipsoid of rotation on a cylindrical base, with six cavities in its mid-
plane to locate TLDs. In this figure units are in cm.
Fig. 1.45: two halves of a breast phantom, hemi-ellipsoid on a cylindrical base, with six disk cavities to locate six
TLDs. [57]
TLD Homogeneity
A group of 105 TLDs annealed, before of the exposure to an X-ray beam, to 400°C for
one hour and after to 100°C for two hours. Are then left to cool to room temperature and
only after two days may be exposed to an X-ray beam. Exposure used parameters are: 80
57
kVp, 0.25 mA and 600s exposure time at a distance of 62.5 cm from the source.
Depending on the read value of the charge, (from minimum of 197.8 nC to a maximum
of 248.0 nC with an average value of 220±11 nC), TLDs are grouped into 10 groups.
TLD Calibration
The calibration factor defined for each dosimeter (CFi) is:
CFi =
In which, Qi = read-out value after exposure to dose, in nC; Di = dose expressed as Air
Kerma (that is, the transferred energy by the photons of a X-ray beam to electrons by
ionization per unit mass of air, in Gy) and <B> is the background signal, obtained
reading 20 TLDs that not exposed to the X-ray beam. The uncertainty on calibration
factor obtained by adding the uncertainty on the collected charge and on air kerma
measurements. Calibration measurements for one TLD at a time, are repeated at four
different energies (from 50 to 80 kVp in 10 kVp steps) and at five different mAs (10, 50,
100, 250, 500).
1.4 Breast computed tomography with synchrotron radiation
The synchrotron is a circular and cyclic particle accelerator in which the magnetic field,
necessary for bending the trajectory of the particles, and the variable electric field, which
accelerates the particles, are synchronized with the beam of the particles. In the
synchrotron the particles are held on a closed orbit by a series of bending magnets
arranged in a ring. At each revolution particles, which must be accelerated, pass through
a radiofrequency cavity, placed in one or more straight sections of the accelerator, and
their energy is incremented by an amount ΔE; at the same time since their energy
increases, the magnetic field of the bending magnets is increased in such a way that the
average trajectory remains stable.
The synchrotron is based on one or more accelerators of particles that act as injection
system and a final stage consisting of the "light machine". The electrons, generated from
a thermionic source, are accelerated by a linear accelerator up to the energy of some tens
or hundreds of MeV and then placed in a synchrotron which accelerates them further to
energy greater than 1 GeV. Subsequently, they are sent in a storage ring, where they
rotate at constant energy. Inside the ring magnetic fields generated by bending magnets,
58
undulators and wigglers1 produce synchrotron radiation. With appropriate systems, the
synchrotron light obtained at each passage of the electrons in the magnets is then
channeled in so-called "lines of light", under high vacuum, where you select the desired
wavelength and at the end of which are located user experimental setups. In fig. 1.46 is
shown a schematic representation of a synchrotron with its main elements and in fig. 1.47
there is a basic diagram of a synchrotron for the production of radiation.
Fig. 1.46: Schematic representation of a synchrotron with the following main elements: 1) detector; 2) injector; 3) the
focusing magnet (quadrupole); 4) bending magnet (dipole); 5) cavity to radio frequency.
Fig. 1.47: Basic diagram of a synchrotron for the production of radiation.
Synchrotron source provides highly collimated and bright beams with linear or circular
polarization. Particularly, undulators and wigglers provide beams collimated both in the
1 Magnetic multipole devices consisting of two opposite rows of magnets alternating in polarity, with the field
direction perpendicular to the electron beam, by which the electrons run through a slalom trajectory achieving an
increased production of synchrotron light compared to a conventional bending magnet
59
horizontal and in the vertical directions. Furthermore the source has a very small size
(about 0.1 mm) and a high spatial coherence. In addition the synchrotron source can
provide X-rays of well-defined wavelength and narrow or broad bands of wavelength
(wavelength tunability). Light from an undulator is emitted only in narrow bands of
wavelengths: individual bands can be selected and the center can be modified, by
changing the undulator working parameters; while a bending magnet or a wiggler
generates a broad band and the desired wavelength can be filtered by a monochromator.
The use of monochromatic beams allows, in principle, to reduce the diffused background
of the image, with an X-ray dose constant, increasing the image contrast and image
quality. These features can be applied, for example, in mammography. The source used
in mammography, usually molybdenum anode X-ray tube, has a spectrum of wavelength
which contains two peaks at about 0.63 Å and 0.71 Å, (i.e. 19.4 keV and 17.6 keV
energy of photons), which fall in the region spectrum optimum for mammography (0.59-
0.73 Å i.e. 17-21 keV). The two peaks are superimposed on a continuous spectrum (due
to bremsstrahlung), which extends below the optimal range for mammography,
producing a background which deteriorates the image contrast. The synchrotron radiation
can reduce this problem because it allows acquiring mammographic images with a single
wavelength, which can be adjusted depending on the thickness and density of the breast,
allowing also dose optimization. Many synchrotron facilities and related synchrotron
sources have been used for radiological technique. In particular I will deal the breast CT
with synchrotron radiation developed at ELETTRA, the synchrotron radiation facility in
Trieste, Italy, started as part of the Synchrotron Radiation for Medical Physics
(SYRMEP) project.
1.4.1 The SYRMEP beamline at the ELETTRA, Trieste
The synchrotron radiation for medical physics (SYRMEP) beamline, designed at the
Elettra synchrotron radiation facility in Trieste, is used for medical diagnostic radiology
study since 1996. A synchrotron radiation machine produces a laminar, monochromatic
and tunable beam, at a distance of about 23 m from the source. These characteristics of
the beam allow respectively to: reduce scattered radiation, increasing the image quality;
remove beam hardening artifacts and select the energy according to the organ thickness
and composition determining an important reduction in the delivered dose.
60
Particularly in this thesis work I will introduce results of breast CT with synchrotron
radiation carried out at the SYRMEP beamline at Elettra. The patient bed used at the
SYRMEP beamline is shown in fig. 1.48: the patient is placed prone with one breast
pending from a hole in the bed and compressed as in conventional mammography and
will be scanned vertically through the beam.
Fig. 1.48: Patient bed and scanning system used at the SYRMEP beamline at Elettra, for breast mammography and
tomography with synchrotron radiation. [58]
It consists of (fig. 1.49):
- A bending magnet of the Elettra storage ring which acts as a radiation source;
- A double-crystal Si (111) monochromator which allows adjusting the beam
energy in the range between 8 keV to 35 keV. Both crystals must be perfectly
parallel to each other to avoid the non-uniformities in the monochromatic beam;
- A tungsten slit system, to define the beam shape and to avoid unnecessary
delivered dose;
- A ionization chamber, used to measure the exposure and to calculate the delivered
dose at the sample;
- A linear array detector, aligned with the beam by means of two orthogonal
translation stages. In addition at two orthogonal translation stages there are a
vertical translation stage and a rotational stage to perform CT acquisitions.
61
During acquisitions the detector is stationary in front of the beam while the sample
rotates in front of it; the rotation axis of the sample must be orthogonal with respect to
the detector plane to avoid artifacts on the reconstructed image as regulated by means of
two goniometric system.
Fig. 1.49: Principal components of the SYRMEP beamline at Elettra with relative distances. [58]
62
Chapter 2
Measurements with a CBBCT prototype and with a SR beam
In this chapter I describe the measurements performed by the Medical Physics group
which were analyzed in this thesis.
2.1 The experimental setup of breast CT with synchrotron radiation at
Elettra, Trieste
The experimental setup used at Elettra SYRMEP beamline for imaging and dose
measurements in breast CT with synchrotron radiation is shown in fig. 2.1. The beam,
from the source, passes through a tungsten slit system providing a parallel beam. After
the slit system there is an ionization chamber (IC) to monitor the beam flux. At the
scanner isocenter is placed vertically a breast phantom, at a distance of 72.4 cm from the
IC and 88.2 cm from flat panel detector.
Fig. 2.1: The experimental setup for breast CT with Synchrotron radiation.
Phantoms
Two phantoms were used in this study: phantom 2 and phantom 3. They are made of
PMMA in the form of a rotational ellipsoid divided in two parts, for simulating an
uncompressed breast, with a cylindrical basis, for simulating the chest wall.
The phantom 2, shown in fig. 2.2-2.3, consists of a hemiellipsoid of half-axes of 7cm x
9.5cm with a cylindrical bases of 7cm x 3.5cm. It contains two series of hole inserts for
measurements of spatial resolution, positioned orthogonally to the major semiaxis, and
six small disk cavities (12mm diameter x 1mm depth) for positioning the TLDs (3mm x
3mm), for dose measurements: three places along the axis of rotation (Axtop
, Axmid
,
Axbot
) and three along the edge (PERtop
, PERmid
, PERbot
). The phantom 3, shown in fig.
63
2.4-2.5, has the same dimensions of the phantom 2 but has a large cavity in the center,
containing two cylindrical inserts (insert A and insert B) of PMMA with 7cm diameter
and two heights, (4cm for the insert A and 3cm for the insert B). It has ten internal holes
of different size (table 2.1), for measuring the spatial resolution.
Fig. 2.2: Photo of the phantom 2. On the left, phantom 2 closed. You see inserts for measurements of spatial resolution.
On the right, phantom opened. You see six disk cavities for the positioning of TLDs.
Fig. 2.3: Scheme of the phantom 2 with the size of the holes. On the left, you see six cavities for housing TLDs: 3
places along the axis of rotation (Axtop, Axmid, Axbot) and 3 along the edge (PERtop, PERmid, PERbot). On the right, you
see inserts for spatial resolution.
Fig. 2.4: Photo of the phantom 3 with the cylindrical inserts contained in it.
64
Fig. 2.5: On the left, axial scheme of the phantom 3 in which we see, in blue, the holes 8mm, 4mm, 2mm, 1mm and
0.5mm diameter. On the right scheme is a 3D plot of the phantom.
The holes of the phantom’s inserts were filled with different materials to simulate the
adipose tissue of the breast, calcifications and other tissues of different density. The table
3 shows a list of the filled holes, the materials used and the size of the holes. Fig. 2.6
shows a scheme of the insert of the phantom 3 with holes of various sizes and filled by
various materials, as listed in table 2.1. The approximate sizes of egg shells fragments,
contained in phantom inserts are: 2.2mm x 1.4mm; 0.5mm x 0.4mm; 1.1mm x 0.9mm.
N° hole Φ (mm) Insert A (Φ = 7cm)
Content ρ (g/cm3)
Insert B (Φ = 7cm)
Content ρ (g/cm3)
1 8 Egg shells 2.35
Fragments
Bologna fat
and egg shells -
fragments
2 0.5 Air - Air -
3 1 Powder of CaCO3 2.93 Air -
4 2 Nylon wires:
Φ1 = 1.0mm 1.14
Φ2 = 0.6mm 1.18
Air -
5 4 Water 1.0 Olive oil 0.886
6 1 Nylon wire:
Φ3 = 0.8mm 1.19
Air
7 0.8 Nylon wire:
Φ4 = 0.7mm 1.04
Air -
8 0.2 Air - Air -
9 0.4 Air - Air -
10 0.6 Air - Air -
Table 2.1: Size and content of the holes of the inserts A and B of the phantom 3.
65
Fig. 2.6: Scheme of the insert of the phantom 3 with sizes holes.
2.1.1 CT Imaging measurements
The scheme of the acquisition geometry for imaging is shown in fig. 2.7. The digital
detector is a CMOS flat panel (CsI:Tl scintillator) with a pixel sizes of 50 µm. The
isotropic resolution for CT reconstruction is (100 µm)3 or (200 µm)
3.
Fig. 2.7: Scheme of the acquisition geometry for imaging.
360 projection images of the two phantoms were acquired at different energies of the
incident beam (20, 24, 28, 30, 32, 34 keV), turning the phantom in steps of 1 to 360
degrees. The entire phantom has been scanned by the beam (transverse dimensions of
120mm x 4mm) by means of a vertical translation with steps of 2mm.
The measurements at Elettra, were taken in two shifts:
- The first shift in July 2008: the insert A and the insert B of the phantom 3 were
scanned at a single height, corresponding to the slice where there are the insert’s
holes with the energy of the incident beam of 28, 24 and 20 keV.
- The second shift in November 2009: the phantom 2 with two plastic supports
containing five microcalcifications each, was scanned with the energy of the
66
incident beam of 34, 32, 30, 28 and 24 keV with steps of 2mm. The size of
microcalcifications varies from 350µm to 450µm.
Table 2.2 shows a scheme of all acquired and reconstructed phantoms at various energies
while in table 2.3 are reported the inherent Aluminum beam filters which have been used
for each acquired phantom of table 2.2. Particularly for phantom 2 has been used external
filter of 4 mm too.
E (keV) 34 32 30 28 24 20
Phantom 3 Insert B Insert B Insert A
Phantom 2 Phantom 2 Phantom 2 Phantom 2 Phantom 2 Phantom 2
Table 2.2: Scheme of acquired and reconstructed phantoms at various energies. At 34, 32 and 30 keV has been
acquired the phantom2 with microcalcifications; at 28 and 24 keV have been acquired the insert B of the phantom3 and
the phantom2 with microcalcifications; finally at 20 keV has been acquired the insert A of the phantom3.
E (keV) 34 32 30 28 24 20
Phantom 3 4.0mm 5.0mm 2.75mm
Phantom 2 7.5mm 7.9mm 7.0mm 5.0mm 2.25mm
Table 2.3: Scheme of inherent Aluminum filters which have been used for each acquired phantom at various energies,
corresponding to the table 2.2.
2.1.2 Image processing
The work in this thesis included the reconstruction of the acquired images using the
ImageJ code and the commercial CT reconstruction software, Cobra.
ImageJ is a “public domain” software for the image processing available on the NIH
(National Institute of Health) website, www.nih.gov (fig. 2.8).
To suppress the electronic noise and to eliminate the malfunctioning lines of the detector
array on each acquired image, offset and gain correction (i.e. subtraction of a dark field
image and division by a uniform-irradiation image also called flat-field) was performed.
Thus, the final image was:
Ic =
where Ii is the input image, D the dark image, F the flat-field and Ic the corrected image.
When imaging a single slice of the phantom with 360 projections, first a stack of the
input image was created. In the case of multiple acquisitions where the phantom was
shifted vertically to be scanned completely, in order to obtain the correct final image a
macro was created with ImageJ.
67
Fig. 2.8: On the left, ImageJ software available on the NIH (National Institute of Health) website, www.nih.gov. On
the right, the logo of the Feldkamp’s filtered back projection reconstruction software (COBRA by EXXIM Computing
Corp. Pleasanton, CA, USA).
Fig. 2.9: View of the screen of Cobra main parameters and corresponding explanations.
68
The images have been then converted into raw files using a macro executed with ImageJ
and raw data have been then reconstructed using a filtered back projection software:
COBRA by EXXIM Computing Corp. Pleasanton, CA, USA, (fig. 2.8). In fig. 2.9 is
shown the screen shot of its main panel for setting the parameters of the CT
reconstruction. The CT values were scaled to represent material density (mg/cm3) in the
phantom.
In the following are shown the reconstructed images of CT slices in the phantom at
different energies. For a quantitative evaluation of the processing image quality the
following figure of merit will be used:
The image noise:
N = σ
where σ is the standard deviation of density values in the background (σB) or in
the detail (σi) of the CT slice image.
The contrast to noise ratio (CNR):
CNR =
where σB and ρB are, respectively, the standard deviation and the density mean
value in the background and ρi is the density mean value in the detail.
The contrast to noise ratio per unit dose (CNRD) calculated as the ratio of CNR
to the square root of the absorbed dose:
CNRD =
where CNR is the contrast to noise ratio and dose is the air kerma at the phantom
position.
The percent coefficient of variation (COV (%)):
COV (%) =
* 100
where σi and ρi are, respectively, the standard deviation and the mean density
value in the background.
The signal to noise ratio (SNR):
SNR =
where ρi is the mean density in the detail and σB is the standard deviation of the
density in the background.
69
To evaluate the mean density and its standard deviation in the background (ρB and σB
respectively) and in the details (ρi and σi respectively) of each CT slice, a rectangular
ROI (5 x 5 pixels) has been identified.
The following shows the common parameters to all measures:
- Distance from the axis of the phantom to the slit = 72.4 cm
- Distance from the axis of the phantom to the ionization chamber = 61.5 cm
- Distance from the axis of the phantom to the flat panel = 88.2 cm
- Scan angle = 360° in 360 steps
- Modality = step and shoot
For the inserts A and B of the phantom 3, 360 projections have been acquired at a single
height of the insert (corresponding to the slice where there are holes filled with different
materials) in about 1.17 s, while the phantom 2 has been acquired with vertical steps of 2
mm in about 8 minutes.
In these images various artifacts are observed which degrade the detail contrast visibility:
- Streaks artifacts occurring as black stripes in the image. They are often seen in
CT images around materials that block most X-rays, such as metal or bone. These
streaks can be caused by undersampling, photon starvation, motion, beam
hardening, or scatter.
- Ring artifacts: they are due to malfunctioning detector pixels.
Phantom 3
For the phantom 3 the images correspond to the insert A at 20 keV and the insert B at
two different energies, 28 and 24 keV. The projections of the phantom 3 were acquired at
4x4 pixel binning (560 x 50 pixels) for a scan angle of 360° with a step of 1°, for a total
of 360 projections, in a step-and-shoot modality. In the same temporal sequence and with
the same acquisition parameters have been acquired flat-field and dark-field images to
perform offset and gain corrections. The CT reconstructions of the phantom 3 have been
made with a Shepp-Logan filter, an isotropic voxel dimension of 200 x 200 x 200 µm3
and final 3D CT matrix size of 448 x 448 x 32 voxels.
Table 2.4 shows the current variation during acquisition for phantom 3 at different
energies. In the figures below are shown the reconstructed slices of the insert A and B of
the phantom 3 (fig 2.6) at different incident beam energies, 28, 24 and 20 keV.
Particularly in figures 2.10, 2.14 and 2.18 are shown slices of the reconstructed images.
70
28 keV 24 keV 20 keV
Θ1
I1 (nA)
0°
0.091
0°
0.115
0°
0.334
Θ2
I2 (nA)
176°
0.090
360°
0.113
180°
0.321
Θ3
I3 (nA)
300°
0.088
360°
0.319
Table 2.4: Current variation during acquisition for phantom 3 at different energies.
The figures 2.10 and 2.14 show how, scrolling through the slices of the insert it is clearly
possible to observe the various materials within the insert holes, up to the hole of 0.2 mm
diameter, as reported in table 2.1. Instead in the case of the beam at 20 keV, (fig. 2.18),
holes below 1mm diameter are not well visible due to streak artifacts from the calcium
detail. Figures 2.11, 2.15 and 2.19 show a magnified view in which different materials
are evident. Finally in figure 2.12, 2.13, 2.16, 2.17 and 2.20 are shown the density linear
profiles evaluated along a diameter of the CT slice, where on X axis is the distance in
pixel and Y axis is the density in mg/cm3.
Fig. 2.10: Reconstructed slices of the insert B of the phantom 3 at the incident beam energy of 28 keV. Details
observed are holes filled with air, egg shells fragments, olive oil and animal fat.
71
Fig. 2.11: On the left, axial view of the insert B of the phantom 3 at 28 keV and on the right, magnified view of the
details in which egg shells fragments, animal fat and air are evident.
Fig. 2.12: On the left, density linear profile along the selected line of the third slices, on the right, which passes through
the phantom holes filled with animal fat, egg shells fragments, olive oil and air (28 keV).
Fig. 2.13: On the left, density linear profile along the selected line of the third slices, on the right, which passes through
the phantom holes filled with air (28 keV).
72
Using a Gaussian fit, from profiles of figure 2.12 and 2.13 the FWHM (full width at half
maximum) in mm of the detail profile and the relative error values have been obtained.
These values are shown in table 2.5 and compared with the holes expected diameter (fig.
2.6).
Φexpect 1mm 1mm 0.8mm 0.6mm 0.5mm 0.4mm
FWHM
± σ (mm)
0.87±0.01 0.93±0.02 0.74±0.08 0.63±0.02 0.57±0.03 0.57±0.03
Table 2.5: The FWHM (mm) and the relative error values calculated using a Gaussian fit from profiles at 28 keV.
The spatial resolution, can be measured analyzing an intensity profile. In an intensity
profile the FWHM of the line profile across the image of the detail, is related to spatial
resolution and to effective detail size: by subtracting in quadrature from the FWHM
widths the nominal size of the smallest detail, an estimate of the BCT scanner spatial
resolution can be obtained.
In particular, the spatial resolution of these CT scans at 28 keV, considering the hole of
0.4mm diameter, in the profile in fig 2.12, and the corresponding FWHM (table 2.5), is
of 0.41mm.
Fig. 2.14: Reconstructed slices of the insert B of the phantom 3 with the incident beam of 24 keV. Are observed holes
filled with air, egg shells fragments and animal fat. The hole containing animal fat and egg shells fragments shows the
so-called streaks artifacts, due to the heterogeneity of the objects contained in it and to the difference of absorption.
73
Fig. 2.15: On the left, axial view of the insert B with details, of the phantom 3 at 24 keV. On the right, magnified view
of the details in which egg shells fragments, olive oil, animal fat and air are evident.
Fig. 2.16: On the left, density linear profile along the selected line of the slice, on the right, which passes through the
phantom holes filled with animal fat, egg shells fragments, olive oil and air (24 keV).
Fig. 2.17: On the left, linear profile along the selected line of the slice, on the right, which passes through the phantom
holes filled with air (24 keV).
74
Fig. 2.18: Reconstructed slices of the insert A of the phantom 3 with the incident beam of 20 keV. Holes are filled with
CaCO3, egg shell fragments, nylon wires and air. This slice presents streaks artifacts very pronounced.
Fig. 2.19: On the left, axial view of the insert A of the phantom 3 at 20 keV and on the right, magnified view of the
details in which nylon wires are evident.
Fig. 2.20: On the left, density linear profile along the selected line of the slice, on the right, which passes through the
phantom holes filled with egg shells fragments, CaCO3, nylon wires (20 keV).
75
In the following are reported the table with measured dose mean values (expressed as air
kerma, AK) in mGy, calculated as the mean of the measured dose values from TLD
dosimeters (table 2.6) and the table (table 2.7) with the calculated parameters for the
image quality for the insert A e B of the phantom 3 at 28, 24 and 20 keV. In order to
evaluate image quality parameters a ROI (5 x 5 pixels) on the background (PMMA) and
on the details, placed in the insert A and B holes of the phantom 3, has been selected and
the mean value and the standard deviation have been measured by the software ImageJ.
From these values have been then calculated: the percent COV, CNR, CNRD and SNR.
28 keV 24 keV 20 keV
AK (mGy) 58.4 82.8 154.7
Table 2.6: Dose mean values (expressed as air kerma, AK), in mGy, measured at various energies for phantom 3.
From the measured values, reported in table 2.7, it can be deduced that all the details
with a high CNR and SNR (> 5.5) are clearly visible. In figure 2.21 is shown the trend of
the values of the CNR as a function of material density. Furthermore, we observe that
the CNR values of the eggshell and CaCO3 details are very high and show a spatial
structure, as is shown in fig. 2.22, while the air level has a very high noise and its signal
level is saturated (fig.2.22).
Fig. 2.21: Trend of the values of the CNR as a function of material density: CNR increases with material density.
Points represent the nylon wire, animal fat, olive oil, CaCO3 and eggshell fragments values, respectively.
76
Phantom
3 insert
Material ρexpect
(mg/cm3)
ρmeas.
(mg/cm3)
σ
(mg/cm3)
COV
(%)
CNR CNRD
(mGy-1/2
)
SNR
Insert B
28 keV
PMMA 1190 1194 18 1.5
Olive oil 890 880 29 3.3 17 2.2 49
Animal fat 870 916 28 3.1 15 2 51
Eggshell
fragments
2350 6692 - - 122 16 149
Air cavity 1000 18 20 - 62 8.2 1
Insert B
24 keV
PMMA 1190 1194 28 2.3
Olive oil 890 761 39 5.1 7 0.7 13
Animal fat 870 890 48 5.4 5 0.6 22
Eggshell
fragments
2350 6939 - - 186 20.4 224
Air cavity 1000 3 27 - 34 3.7 0.1
Insert A
20 keV
PMMA 1190 1199 115 9.6
CaCO3 2930 3885 - - 58 4.7 84
Nylon wire 1140 577 26 4.5 5.4 0.4 5
Eggshell
fragments
2350 1727 157 - 15 1.2 17
Air cavity 1000 -104 38 - 23 1.9 1.8
Table 2.7: Density mean value, standard deviation, COV (%), CNR, CNRD and SNR evaluated on CT slices for insert
B and A of the phantom 3. It is also shown the expected density value for the different materials.
77
Phantom 2
Phantom 2 with microcalcifications, from 350 to 450 µm size, has been scanned
vertically in steps of 2mm and for each step 360 projections were acquired with incident
beam energy of 34, 32, 30, 28 and 24 keV, at 1x1 pixel binning (2240 x 68 pixels) for a
scan angle of 360° with a step of 1° in a step-and-shoot modality. An image stack has
been produced with acquired projections for each step. All the stacks have been then
processed and concatenated using a plugin of the Image J software. In order to perform
offset and gain corrections, flat-field and dark-field images have been acquired in the
same temporal sequence and with the same acquisition parameters. The CT
reconstructions of the phantom 2 have been made with a Shepp-Logan filter, an isotropic
voxel dimension of 100 x 100 x 100 µm3 and final 3D CT matrix size of 896 x 896 x 64
voxels.
Fig. 2.22: 3D plot of the insert B of the phantom 3. It shows the different materials structure, in particular can be
observe the eggshell fragments structure and the saturated air.
Figures 2.23 and 2.24 show the axial views of the processed images at 34, 32, 30, 28 and
24 keV. In these images are clearly seen, inside a phantom hole, the microcalcifications.
In fig. 2.25 are shown the sagittal views of the same phantom, at various energies, (34,
32, 30 and 28 keV), in which are evident the two holes of the phantom with five
microcalcifications inside and in fig. 2.26 the line profile along a diagonal containing
microcalcifications is shown. While fig. 2.27 shows the volume views of the processed
images at 34, 32 and 24 keV, in which an image part of the phantom was cut so as to
show the two holes with the five microcalcifications. The phantom shape is different
78
from the actual one because the stacks corresponding to the upper part of the phantom
have not been acquired. Furthermore in figures 2.28 and 2.29 are shown the 3D surface
plot at 32 keV and 28 keV respectively. The five pronounced peaks are due to the five
microcalcifications, the rest is the air contained in the hole of the phantom.
Fig. 2.23: Magnified axial view of the processed image with the incident beam at 34 keV: are clearly seen the
microcalcifications inside a phantom hole.
a) b)
c) d)
Fig. 2.24: Axial views of the processed images with the incident beam at: a) 32 keV; b) 30 keV; c) 28 keV and d) 24
keV respectively.
79
a)
b)
c)
c)
Fig. 2.25: Sagittal views of the processed images at: a) 34 keV; b) 32 keV; c) 30 keV and d) 28 keV. In all cases are
visible the five microcalcifications inside phantom holes.
Fig 2.26: On the left, density linear profile along a diagonal, shown on the right, containing three microcalcifications,
at 28 keV.
80
a)
b)
c)
Fig. 2.27: Volume viewers of the processed images at: a) 34 keV; b) 30 keV and c) 24 keV. The phantom shape is
different from the actual because the stacks corresponding to the upper part of the phantom have not been acquired.
81
Fig. 2.28: On the left, 3D graph of the intensities of pixels in a pseudo color images (non-RGB images) of the selected
ROI in the picture on the right, at 32 keV.
Fig. 2.29: On the left, 3D graph of the intensities of pixels in a pseudo color images (non-RGB images) of the selected
ROI in the picture on the right, at 28 keV.
82
In the following are reported the table with measured dose values (expressed as air
kerma, AK), in mGy (table 2.8) and the table (table 2.9) with the calculated parameters
for the image quality for the phantom 2 at 34, 32, 30, 28 and 24 keV. In order to evaluate
image quality parameters a ROI (5 x 5 pixels) has been selected on the background
(PMMA) and on the phantom 2 holes containing CaCO3, to simulate microcalcifications.
From the CNR and SNR measured values, reported in table 2.9, it can be deduced that all
the microcalcifications are clearly visible (CNR > 5).
34 keV 32 keV 30 keV 28 keV 24 keV
AK (mGy) 50 90.2 95 112 180
Table 2.8: Dose values measured at various energies for phantom 2.
Phantom2 Material ρ
(mg/cm3)
COV(%) CNR CNRD
(mGy-1/2
)
SNR
34 keV PMMA 1196 59 5
CaCO3 2820 - - 28 4 48
32 keV PMMA 1200 70 0.1
CaCO3 3629 - - 35 3.7 52
30 keV PMMA 1195 63 0.1
CaCO3 3889 - - 43 4.4 62
28 keV PMMA 1197 87 0.1
CaCO3 2852 - - 19 1.8 33
24 keV PMMA 1191 85 0.1
CaCO3 3460 - - 27 2 41
Table 2.9: Density mean value, standard deviation, COV (%), CNR, CNRD and SNR evaluated on CT slices for the
phantom 2. The CaCO3 expected value is 2.93 g/cm3.
83
2.1.3 Dose distribution into the phantom
To evaluate the absorbed dose both in air and in the volume of phantom 2, thermo
luminescence dosimeters (TLDs) at different energies of the SR beam (34, 32, 30, 28, 24
and 20 keV) have been used. A first set of measurements has been taken at 28, 24 and 20
keV and a second set of measurements has been taken at 34, 32, 30, 28 and 24 keV. The
TLDs were placed in the six holes situated both along the vertical axis of the phantom
(Axtop
, Axmid
, Axbot
) and in the periphery (PERtop
, PERmid
, PERbot
), as shown in fig. 2.3.
During irradiation the phantom has been rotated by 360° around the vertical axis at each
step. All measurements have been performed by a vertical scan of 2 mm and in the first
set of measurements at 24 keV also of 3 mm. In fig. 2.30 is shown the scheme for the
dose measurements in the phantom with the ionization chamber, for measuring the air
kerma to calculate the calibration factor which converts TLD signal into dose.
Fig. 2.30: Scheme for the measurement of absorbed dose in the phantom 2.
The values obtained from the readings of the TLD (in nC) must be multiplied by the
calibration factor (CF) to convert the collected charge in dose (mGy):
CF =
and
=
Where AKair
is the air kerma, Q is the collected charge and ΔAK and ΔQ are the
respective measurement uncertainties.
To measure the air kerma, the dosimeters were exposed to the beam in air at the same
beam energies as for measurements in phantom. The TLDs have been positioned inside
thin plastic supports at the same distance from the axis of rotation of the phantom from
the beam slit (fig. 2.31). Measurements at each energy have been repeated five times
under the same conditions. The final value of accumulated charge is given by the average
value of these five estimates and by corresponding standard deviation.
84
Fig. 2.31: Scheme of the measurement of TLDs in air to calculate the air kerma. CI = ionization chamber
The following shows the parameters common to all measurements:
- Distance from the axis of the phantom to the slit = 72.4 cm
- Distance from the axis of the phantom to the CI = 61.5 cm
- Texp = 61s
In the following tables are: the values of air kerma (AK) in mGy, the photon flux
(γ/mm²), the collected charge in the TLD (TLD reading) in nC, the absorbed dose by
TLD (D) in mGy, calculated as the product of the TLD reading for the calibration factor
(D = TLD reading * CF), at different energies for the first set of measurements (28, 24
and 20 keV). In the same table have been reported the normalized dose ratio values,
calculated as the ratio of the collected charge in the TLD with respect to the inner
position (Axbot
). The first row of the table contains the value of the energy (E), the beam
width (Δy) and the value of the calibration factor (CF). In addition the mean value and
standard deviation of the measured dose values have been calculated.
E=28KeV Δy=2mm CF = 0.19 (mGy/ nC)
Position TLD# AK
(mGy)
n° γ/mm²
(x108)
TLDreading(
nC)
D±ΔD
(mGy)
Normalized
dose (%)
PERtop 5 (57÷60) (6.73÷7.05) 30.85 5.86 221
PERmid 50 29.25 5.56 209
PERbot 65 26.06 4.95 187
Axtop 14 26.81 5.09 192
Axmid 29 18.9 3.59 135
Axbot 41 13.97 2.65 100
Average 58.5 6.89 4.6±1.2
Table 2.10: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the inner position
(Axbot), of six TLDs exposed at 28 keV.
85
E=24KeV Δy=2mm CF = 0.18 (mGy/ nC)
Position TLD# AK
(mGy)
n° γ/mm²
(x108)
TLDreading
(nC)
D±ΔD
(mGy)
Normalized
dose (%)
PERtop 91 (78÷87) (6.76÷7.47) 34.78 6.26 359
PERmid 98 32.1 5.78 331
PERbot 99 26.88 4.84 277
Axtop 79 27.23 4.90 281
Axmid 71 14.46 2.60 149
Axbot 76 9.7 1.75 100
Average 82.5 7.11
4.4±1.8
Table 2.11: Air kerma, photon fluence, collected charge dose and normalized dose ratio, respect to the inner position
(Axbot), of six TLDs exposed at 24 keV with aperture beam of 2mm.
E=24KeV Δy=3mm FC = 0.18 (mGy/ nC)
Position TLD# AK
(mGy)
n° γ/mm²
(x108)
TLDreading
(nC)
D±ΔD
(mGy)
Normalized
dose (%)
PERtop 93 (70÷73) (6.06÷6.29)
21.46 3.86 231
PERmid 101 18.97 3.41 205
PERbot 58 17.84 3.21 192
Axtop 84 19.07 3.43 206
Axmid 61 8.57 1.54 93
Axbot 69 9.27 1.67 100
Average 71.5 6.17
3±1
Table 2.12: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the inner position
(Axbot), of six TLDs exposed at 24 keV with aperture beam of 3mm.
E=20KeV Δy=2mm FC = 0.22 (mGy/ nC)
Position TLD# AK
(mGy)
n° γ/mm²
(x108)
TLDreading
(nC)
D±ΔD
(mGy)
Normalized
dose (%)
PERtop
45 (144÷160) (8.38÷9.32)
81.53 17.55 140
PERmid
67 79.03 17.01 136
PERbot
104 36.09 7.77 621
Axtop
68 19.78 4.26 340
Axmid
83 8.99 1.94 155
Axbot
73 5.81 1.25 100
Average 152 8.85
8.3±7.3
Table 2.13: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the inner position
(Axbot), of six TLDs exposed at 20 keV.
86
In the following are shown :
- The table of the TLD distances measured from the outer edge of the phantom
(table 2.14);
- The graph of the normalized dose ratio values, calculated as the ratio of the TLD
readings with respect to the inner position (Axbot
), as a function of TLD distance
from the edge of the phantom, in cm, for the measurements at 28, 24 and 20 keV
(fig. 2.32);
- The graph of the normalized dose ratio values as a function of TLD distance from
the edge of the phantom, in cm, for two measurements at 20 keV with a vertical
translation steps of 2mm and 3mm (fig. 2.33).
- The table (table 2.15) and histogram (fig. 2.34) of the normalized dose ratio
values as a function of energy (28, 24 and 20 keV), for beamwidth of 2mm, for
each position of the TLDs (PERtop
, PERmid
, PERbot
, Axtop
, Axmid
, Axbot
).
TLDs position TLD distance from the outer edge
of the phantom
PERbot 1.4
PERmid 1.6
PERtop 1.7
AXtop 4.3
AXmid 6.1
AXbot 7
Table 2.14: TLD distances measured from the edge of the phantom.
Fig. 2.32: Graph of the normalized dose ratio values as a function of TLDs distance from the edge of the phantom, in
cm, for the measures of the first shift (28, 24 e 20 keV).
87
Fig. 2.33: Graph of the normalized dose ratio values as a function of TLDs distance from the edge of the phantom, in
cm, for the two measures to 24 keV with a step of vertical translation of 2mm and 3mm.
TLDs E = 28keV 24keV 20keV
PERtop 221 359 140
PERmid 209 331 136
PERbot 187 277 621
AXtop 192 281 340
AXmid 135 149 155
AXbot 100 100 100
Table 2.15: Normalized dose ratio values in percent, respect to Axbot
, as a function of energy (with beamwidth of
2mm), for different TLDs positions (PERtop
, PERmid
, PERbot
, Axtop
, Axmid
, Axbot
).
Fig. 2.34: Histogram of the normalized dose ratio values as a function of energy (28, 24 and 20 keV), for beamwidth of
2mm, for each position of the TLDs (PERtop
, PERmid
, PERbot
, Axtop
, Axmid
, Axbot
).
88
The following are tables with the values of air kerma (AK) in mGy, the collected charge
in the TLD (TLD reading) in nC, the photon flux (γ/mm²), the absorbed dose by TLD (D)
in mGy calculated as the product of the TLD reading for the calibration factor (D = TLD
reading * CF) at different energies for the second set of measurements (34, 32, 30, 28, 24
keV). In the same table the normalized dose ratio values, calculated as the ratio of the
collected charge in the TLD with respect to the inner position (Axbot
), have been
reported.
E=34KeV Δy=2mm CF = 0.22 (mGy/ nC)
Position TLD# AK
(mGy)
n° γ/mm²
(x108)
TLDreading
(nC)
D±ΔD
(mGy)
Normalized
dose (%)
PERtop 75 19.8 3.43
102.1 22 164
PERmid 49 106.7 22.99 171
PERbot 17 94.25 20.31 151
Axtop 101 93.92 20.24 151
Axmid 61 74.9 16.14 120
Axbot
20 62.3 13.42 100
Average 19.2±3.7
Table 2.16: Dose, photon fluence, collected charge and normalized dose ratio, respect to the inner position (Axbot), of
six TLDs exposed at 34 keV.
E=32KeV Δy=2mm CF = 0.14 (mGy/ nC)
Position TLD# AK
(mGy)
n° γ/mm²
(x108)
TLDreading
(nC)
D±ΔD
(mGy)
Normalized
dose (%)
PERtop 26 31.5 4.87
218.9 30.14 163
PERmid 57 216.5 29.81 161
PERbot 59 205.6 28.31 153
Axtop 31 203.4 28.01 152
Axmid 21 151.8 20.90 113
Axbot 98 134.2 18.48 100
Average 26±5
Table 2.17: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the inner position
(Axbot), of six TLDs exposed at 32 keV.
89
E=30KeV Δy=2mm CF = 0.21 (mGy/ nC)
Position TLD# AK
(mGy)
n° γ/mm²
(x108)
TLDreading
(nC)
D±ΔD
(mGy)
Normalized
dose (%)
PERtop 63 31 4.23
173.4 36.43 185
PERmid 84 150.5 31.62 160
PERbot 105 136.3 28.64 145
Axtop 94 141.1 29.09 150
Axmid 95 109.9 23.09 117
Axbot 97 93.91 19.73 100
Average 28±6
Table 2.18: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the inner position
(Axbot), of six TLDs exposed at 30 keV.
E=28KeV Δy=2mm CF = 0.19 (mGy/ nC)
Position TLD# AK
(mGy)
n° γ/mm²
(x108)
TLDreading
(nC)
D±ΔD
(mGy)
Normalized
dose (%)
PERtop 76 38.8 4.02 135.1 26.22 229
PERmid 47 131.6 25.54 224
PERbot 4 124.7 24.20 212
Axtop 18 119.9 23.27 204
Axmid 100 85.2 16.54 145
Axbot
91 58.87 11.43 100
Average 21.2±5.9
Table 2.19: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the inner position
(Axbot), of six TLDs exposed at 28 keV.
E=24KeV Δy=2mm CF = 0.18 (mGy/ nC)
Position TLD# AK
(mGy)
n° γ/mm²
(x109)
TLDreading
(nC)
D±ΔD
(mGy)
Normalized
dose (%)
PERtop 41 157 1.35
174.1 31.78 334
PERmid 65 158.3 28.90 304
PERbot 70 142.6 26.03 274
Axtop 29 126.4 23.07 243
Axmid 19 76.93 14.04 148
Axbot 16 52.07 9.51 100
Average 22.2±8.7
Table 2.20: Air kerma, photon fluence, collected charge, dose and normalized dose ratio, respect to the inner position
(Axbot), of six TLDs exposed at 24 keV.
90
Below are shown:
- the graph of the normalized dose ratio values, as a function of TLDs distance
from the edge of the phantom, in cm, for the measurements at 34, 32, 30, 28 and
24 keV, (fig. 2.35);
- The table (table 2.21) and histogram (fig. 2.36) of the normalized dose ratio
values as a function of energy (34, 30, 32, 28 and 24 keV), for beamwidth of
2mm, for each position of the TLDs (PERtop
, PERmid
, PERbot
, Axtop
, Axmid
, Axbot
).
Fig. 2.35: Graph of the normalized dose ratio values as a function of TLDs distance from the edge of the phantom, in
cm, for the measures of the second shift (34, 32, 30, 28 and 24 keV).
E = 34keV 32keV 30keV 28keV 24keV
PERtop 164 163 185 229 334
PERmid 171 161 160 224 304
PERbot 151 153 145 212 274
AXtop 151 152 150 204 243
AXmid 120 113 117 145 148
AXbot 100 100 100 100 100
Table 2.21: Normalized dose ratio values in percent, respect to Axbot, as a function of energy (with beamwidth of
2mm), for different TLDs positions (PERtop, PERmid, PERbot, Axtop, Axmid, Axbot).
91
Fig. 2.36: Histogram of the normalized dose ratio values as a function of energy (34, 30, 32, 28 and 24 keV), for
beamwidth of 2mm, for each position of the TLDs (PERtop
, PERmid
, PERbot
, Axtop
, Axmid
, Axbot
).
An important observation about the CT system geometry with synchrotron radiation is
the following: since the beam is laminar, to irradiate the whole volume of the phantom is
necessary to perform a stepped scan of the sample. Since the beam profile is not flat,
there is an area that is irradiated twice, i.e. a double absorbed radiation dose (fig. 2.37).
Fig. 2.37: The plot number 1 shows the beam profile, the plot number 2 shows the same beam profile shifted of 2mm
and plot number 1+2 shows the sum of two profiles. As shown, since the beam profile is approximately Gaussian, there
is an area that is irradiated twice, i.e. a double absorbed radiation dose.
92
2.2 The experimental setup for CBBCT prototype at the University of
Napoli
At the University of Napoli and INFN two experimental setups were used: low energy
and high energy setup. In the first, (fig. 2.38) the X-ray tube and the flat panel detector
were in a fixed position while the breast phantom is rotated during the acquisition, with
the tube focal spot at the same height of the Axmid
position. For this setup, the distances
source-to-isocenter and source-to-detector were respectively 56.2cm and 62.6cm. The X-
ray beam, in addition to inherent filtration of 1.8mm Al, was filtered by 0.05mm Cu and
it has an output cone angle of about 40 degrees. In the high energy setup, the X-ray tube
and the flat panel detector are placed on the arms of the gantry, which rotates around the
phantom with a maximum rotation radius of 50cm. To have a high energy X-ray photon
beam, the X-ray beam is filtered by 0.2mm Cu. The experimental high energy setup used
imaging and dose measurements in cone-beam breast CT, is shown in fig. 2.39. The
breast phantom is placed vertically at the scanner isocenter from the top, to simulate a
pendant breast from a bed. In this study the isocenter was 38.5cm from the source, the X-
ray beam was horizontal and the cone angle was 40 degrees. During the acquisition the
gantry, with the X-ray tube and the flat panel detector, rotates for 360 deg to acquire 360
projections in about 300 s.
Fig. 2.38: The low energy setup. The X-ray tube and the flat panel detector were in a fixed position while the breast
phantom is rotated during the acquisition.
93
Fig. 2.39: Experimental setup for CBBCT at the University of Napoli. Shown: X-ray tube (1); flat panel detector (2);
rotating gantry (3); pinhole compact gamma camera (4); PMMA breast phantom (5). [55]
To control the acquisition procedure a software was realized using C programming
language in the Windows/CVI National Instruments environment, whose main graphical
panel shown in fig. 2.40. It is divided in three parts: in the left panel there are all the
settings to regulate the X-ray tube parameters (kV, mA) and to monitor them during the
acquisition; in the center there are all the settings to control the flat panel detector and
finally, in the right panel there are all motor controls.
Fig. 2.40: Main screen of the software to control the x-ray tube (on the left), flat panel detector (in the center) and the
motor. [59]
94
The PMMA phantom used for imaging and dose studies are respectively: phantom 1 and
phantom 2. They have the form of a hemiellipsoid of rotation with a cylindrical base. The
hemiellipsoid has half-axes of 7 and 9.5cm, which simulate an uncompressed breast of
14cm diameter at the chest wall while the cylindrical base has radius of 7cm and height
of 3cm and simulate the chest wall.
The phantom 1, unlike phantom 2, has a large cavity in the center, where it is possible to
put two cylindrical inserts (insert A and B). Each insert has a set of holes which are
parallel to the chest wall-nipple direction with a size from 8 to 0.2mm diameter. Each
hole was filled with various materials to simulate details with different contrast.
Particularly inside the 8mm hole the insert A (fig. 2.41) contains animal fat with eggshell
fragments (of size 2.2, 1.1 and 0.5mm), to simulate calcifications, and olive oil, to
simulate adipose tissue, inside the 4mm hole. The insert B (fig. 2.41), contains distilled
water inside the 8mm hole, extravirgin olive oil (0.89 g/cm3 density) inside the 4mm
hole, four nylon wires (0.6mm diameter, 1.11 g/cm3 density) inside the 2mm hole,
CaCO3 calcium carbonate powder (2.93 g/cm3 density) inside the 1mm hole and two
holes of 5mm diameter filled with air.
The phantom 2, which consists of two half-parts, contains in one half-part six cylindrical
cavities (12mm diameter x 1mm depth) for housing TLDs to measure the dose
distribution (fig. 2.42). In addition this phantom contains two sets of details,
perpendicular to the CT rotation axis: a set of six cylindrical holes, with the same depth
but con different diameter (two of 1mm, two of 1.5mm and two of 2mm), placed along
the perpendicular axis at the chest-nipple direction; the second set is a 4 x 4 array of
holes with a diameter of 2.5, 2, 1.5 and 1mm respectively for each four holes in a same
row (fig. 2.42).
Fig. 2.41: Scheme of the insert A and the insert B of the phantom 1, with holes of different sizes filled with various
substances.
95
Fig. 2.42: Scheme of the phantom 2 with six cylindrical cavities (12mm diameter x 1mm depth) for housing TLDs and
two sets of details.
2.2.1 Imaging measurements
Imaging measurements have been performed by the medical physics group at the
University of Napoli and the results reported below were taken from references [59, 60,
61].
To measure the air Kerma and the Half Value Layer (HVL) of the X-ray beam an
ionization chamber model 20X6-6, Radcal Corporation, Monrovia, CA; sensitive volume
6 cm , pen shaped, diameter 24.5mm) has been used. In addition a Cu filter of 100µm in
front of X-ray tube has been used. The HVL measurements have been obtained by
varying the tube voltage from 50 to 80 kVp in step of 10 kVp at the isocenter of the
setups, and with aluminum foils of 1mm thickness. From the HVL values they calculated
the beam effective energy (Eeff) values (table 2.22). The air kerma has been measured at
the isocenter (38.5cm from the tube focal spot), by varying the X-ray tube voltages in the
range 50-80 kVp, in step of 10 kVp and fixed the tube current at 0.25 mA. In addition the
tube load varied in the range 10-100 mAs depending on the exposure time, 40-400s.
From these values has been derived the tube output (mGy/mAs) and the calculated MGD
for a 50/50 breast (50% glandular and 50% adipose tissue) 14cm diameter, as shown in
96
table 2.23, using the normalized glandular dose coefficients (DgNCT) provided in ref.
[45].
Tube Voltage (kVp) HVL (mm Al) Eeff (keV)
80 4.22 39
70 3.67 36
60 3.25 34
50 2.63 31
Table 2.22: HVL values and corresponding effective energy measured at different tube.
Tube
Voltage
(kVp)
Tube
load
(mAs)
Tube
output
(mGy/mAs)
Air
kerma
(mGy)
DgNCT
(mGy/mGy)
Calculated
MGD
(mGy)
80 15 0.34 5.10 0.70 3.6
70 15 0.25 3.75 0.65 2.4
60 30 0.17 5.10 0.60 3.1
50 60 0.10 6.00 0.54 3.2
Table 2.23: Values of air kerma, tube load, DgNCT and calculated MGD at various tube voltages.
To perform imaging measurements with this CBBCT setup, the phantom has been placed
at the scanner isocenter at a distance of 385mm from the X-ray source and 110mm from
the detector, with scans from 50 to 80 kVp in steps of 10 kVp, with a corresponding
calculated MGD values reported in table 2.23. The phantom projections have been
acquired at 4 x 4 pixel binning (560 x 586 pixels, 200 x 200 µm pixel size), for a scan
angle of 360° with a step of 0.86°, for a total of 420 projections. The projections have
been then reconstructed with a Ram-Lak filter and a voxel size of 0.6 x 0.6 x 0.6mm3
(250 x 250 x 146 pixels) while to correct for the beam hardening artifact, the Jian and
Hangmian algorithm [62] has been used, while ring artifacts were not corrected.
In figures 2.43-2.46 are shown the axial and sagittal views of the phantom 2 at different
tube voltages (80, 70, 60 and 50 kVp) and at the linear profile along the diameter of a
13cm axial slice. We can observe that all sets holes filled with air are clearly visible
down to 1mm diameter in the scans at all tube voltages. Using a Gaussian fit, from the
profile in fig. 2.43-2.46, the FWHM (mm) values have been calculated (table 2.24).
97
Fig. 2.43: On the left axial and sagittal views of the phantom 2 acquired at 80 kVp; on the right linear profile along the
diameter of a 13cm axial slice of the same phantom.
Fig. 2.44: On the left axial and sagittal views of the phantom 2 acquired at 70 kVp; on the right linear profile along the
diameter of a 13cm axial slice of the same phantom.
98
Fig. 2.45: On the left axial and sagittal views of the phantom 2 acquired at 60 kVp; on the right linear profile along the
diameter of a 13cm axial slice of the same phantom.
Fig. 2.46: On the left axial and sagittal views of the phantom 2 acquired at 50 kVp; on the right linear profile along the
diameter of a 13cm axial slice of the same phantom.
By subtracting in quadrature from the FWHM widths the nominal size of the detail, an
estimate of the BCT scanner spatial resolution can be obtained. For example, the spatial
resolution of these CT scans at 50 kVp (corresponding to an effective energy of about
28.7 keV), considering the hole of 1mm and the corresponding FWHM (table 2.24), is
0.63mm.
99
Φ 1mm 1mm 1.5mm 1.5mm 2mm 2mm
FWHM
(mm)
50 kVp 1.18 1.27 1.54 1.62 1.94 2.02
60 kVp 1.21 1.30 1.55 1.58 1.97 2.00
70 kVp 1.17 1.18 1.44 1.44 1.91 1.87
80 kVp 1.11 1.28 1.53 1.54 2.03 2.00
Table 2.24: FWHM values for each detail of diameter Φ, at different tube voltages, obtained by the Gaussian fit to the
line profiles of the air-filled details in fig. 2.43-2.46. There is little variation in the detail resolution at different tube
voltages.
In figures 2.47-2.48 are shown the linear profiles along the details in axial views of the
insert A and B, at 80 kVp. In fig. 2.49 is shown, instead, the axial view of the insert A of
the phantom 1 at 80 kVp and a magnification of the details containing animal fat and
three eggshell fragments. We can observe that all details are well visible.
Fig. 2.47: In the upper right is shown the linear profile of the insert A, along the region of interest 1, containing details,
and the bottom right the linear profile along the region of interest 2, as shown in the image in the top left corner, at 80
kVp.
100
Fig. 2.48: In the upper right is shown the linear profile of the insert B, along the region of interest 1, containing details,
and the bottom right the linear profile along the region of interest 2, as shown in the image in the top left corner, at 80
kVp.
Fig. 2.49: On the right, axial view of the insert A of the phantom 1 at 80 kVp and on the left, magnification of the
details containing animal fat and three eggshell fragments.
101
In the work [60] the same insert A and B of the phantom 1 have been acquired with the
same modality. For a quantitative evaluation, the image quality parameters (contrast,
CNR and CNRD) have been calculated (table 2.25).
Phantom 1
insert
Material Detail
size(mm)
Contrast
(HU)
CNR CNRD
(x 104 Gy
-1/2)
B CaCO3 1 424 14 8.3
B Nylon+air 2 264 10 5.9
A Air 2 451 14 8.2
B Air 4 764 16 9.5
A CaCO3 4 1141 26 15
B Animal fat 8 157 2.4 1.4
A Olive oil 8 143 2.7 1.6
Table 2.25: Detail contrast, CNR and CNRD evaluated on CT coronal slices for phantom 1.
In another study [61], an investigation on the visibility and detectability of the
microcalcifications has been performed, using different detector pixel size (50, 200 µm),
at various air kerma (5, 7.5, 9 mGy) at X-ray tube voltage 80 kVp (Eeff = 44.4 keV).
Phantom 2 with small eggshell fragments, for simulating microcalcifications, positioned
in the cylindrical holes for the TLD dosimeters housing, has been used. The total
dimensions of microcalcifications were in the range of 350-450 µm or 500-700 µm. To
evaluate the microcalcifications visibility the SNR and the image contrast (CNR) in a
ROI (corresponding to 37 pixels selecting a flat panel pixel size at 200 µm and to 256
pixels at 50 µm), have been calculated, for different detector pixel size (4 x 4 binning and
1 x 1 binning) and different air kerma (table 2.26). In figure 2.50-2.52 the coronal views
and the profiles along the selected line are shown. We observe that, by increasing the
delivered air kerma (at the same flat panel acquisition pixel), the SNR and CNR increase,
but at 50 µm pixel size are too low.
Air Kerma
(mGy)
Flat panel acquisition pixel
(µm)
SNR CNR
5.0 200 3.2 0.26
7.5 200 4.6 0.29
9.0 50 1.9 0.12
Table 2.26: SNR and CNR values evaluated on the acquired images at different air kerma and with different flat panel
pixel size.
102
Fig. 2.50: A) Coronal view of central hole containing five microcalcifications, acquired at 200 µm flat panel pixel size,
at 80 kVp and at an air kerma of 5.0 mGy. B) The same slice shown in A) processed using a FFT band pass filter. C)
Linear profile along a diagonal containing three microcalcifications (B, A and D). The microcalcifications FWHM is
also indicated.
Fig. 2.51: A) Coronal view of central hole containing five microcalcifications, acquired at 200 µm flat panel pixel size,
at 80 kVp and at an air kerma of 7.5 mGy. B) The same slice shown in A) processed using a FFT band pass filter. C)
Linear profile along a diagonal containing three microcalcifications (B, A and D). The microcalcifications FWHM is
also indicated.
103
Fig. 2.52: A) Coronal view of central hole containing five microcalcifications, acquired at 50 µm flat panel pixel size,
at 80 kVp and at an air kerma of 9.0 mGy. B) The same slice shown in A) processed using a FFT band pass filter. C)
Linear profile along a diagonal containing three microcalcifications (B, A and D). The microcalcifications FWHM is
also indicated.
2.2.2 Dose distribution into the phantom
Dose measurements have been performed by the medical physics group at the University
of Napoli and the results reported below were taken from ref. [63].
In this study to measure the air Kerma and the Half Value Layer (HVL) of the X-ray
beam both for low and high energy setup an ionization chamber have been used. The air
Kerma has been measured at the isocenter with X-ray tube voltages in the range 50-80
kVp, in step of 10 kVp. The tube output (mGy/mAs) is shown in fig. 2.53. The HVL
measurements have been obtained by varying the tube voltage from 50 to 80 kVp in step
of 10 kVp at the scanner isocenter using aluminum foils and the ionization chamber. The
beam effective energy (Eeff) values, both for high and low energy setups, are reported in
tables 2.27 and 2.28 respectively. Dose measurements inside phantoms have been
performed in an effective energy range of 35.3-44.4 keV. In each of the six phantom
cavities were positioned three TLDs. The phantom, with eighteen TLDs inside, has been
exposed to the X-ray beam at tube voltages of 50 keV, 60 keV, 70 keV and 80 keV; and
with a fixed tube current of 0.17 mA. TLDs charge values (Q in nC), charge per air
104
Kerma (Q/AK) and dose (Q x FC) in the Axbot
position, both for high and low energy
setups, are shown in tables 2.29 and 2.30 respectively.
Fig. 2.53: Plot of tube output (air Kerma per mAs) as a function of the tube voltage at the isocenter of low, on the left,
and high, on the right, energy setup.
Tube Voltage (kVp) HVL (mm Al) Eeff (keV)
80 5.6 44.4
70 4.9 41.6
60 4.2 38.7
50 3.4 35.3
Table 2.27: HVL values and corresponding effective energy measured at different tube voltages for high energy setup.
Tube Voltage (kVp) HVL (mm Al) Eeff (keV)
80 3.0 33.5
70 2.7 31.9
60 2.4 30.5
50 2.0 28.7
Table 2.28: HVL values and corresponding effective energy measured at different tube voltages for low energy setup.
105
kVp
Eeff (keV)
mAs
AK (mGy)
Q (nC) Q/AK (nC/mGy) D (mGy)
80 kVp
44.4 keV
33.2 mAs
6.6 mGy
27.8±0.3 4.2 5.6
70 kVp
41.6 keV
55.1 mAs
7.1 mGy
28.8±0.2 4.0 5.2
60 kVp
38.7 keV
97.9 mAs
7.8 mGy
28.8±0.1 3.7 5.8
50 kVp
35.3 keV
231.7 mAs
9.3 mGy
32.6±1.2 3.5 5.7
Table 2.29: TLDs charge values (Q in nC), charge per air Kerma (Q/AK) and dose (QxFC) measured in the Axbot
position for high energy setup.
kVp
Eeff (keV)
mAs
AK (mGy)
Q (nC) Q/AK (nC/mGy) D (mGy)
80 kVp
33.5 keV
73.75 mAs
20.3 mGy
64.68 3.2 13.1
70 kVp
31.9 keV
73.75 mAs
15.3 mGy
46.28 3.0 8.3
60 kVp
30.5 keV
73.75 mAs
11.3 mGy
28.75 2.5 5.7
50 kVp
28.7 keV
73.75 mAs
7.3 mGy
13.39 1.8 2.3
Table 2.30: TLDs charge values (Q in nC), charge per air Kerma (Q/AK) and dose (QxFC) measured in the Axbot
position for low energy setup.
106
To determine the dose distribution inside the breast phantom the normalized dose ratio
values have been calculated, i.e. the ratio of the charge values in the six positions
(PERtop
, PERmid
, PERbot
, Axtop
, Axmid
, Axbot
) with respect to the intermost position
(Axbot
). These results are shown in fig. 2.54 for both high and low energy setup.
a) b) Fig. 2.54: Normalized dose ratio values respect to the intermost position, Axbot. a) For high energy setup. b) For low
energy setup.
By these results it is possible to derive that for low energy setup (28.7-33.5 keV) the dose
variation is more significant with respect to high energy setups (35.3-44.4 keV). In fact
for low energy the percentage variation range it is 11%-82% while for high energy it is -
7%-33%. In addition for low energy setup the maximum variation with respect to the
inner position is 66% in the radial direction and 63% in the axial direction; while for high
setups energy the maximum variation with respect to the inner position is 33% in radial
direction and 2% in axial direction. So the absorbed dose is more homogeneous with
high energy beams than with low energy beams. A cause of this radial non-homogeneity
for low energy derives from the lower penetration of X-rays.
2.3 Comparison between SR based and CBBCT based results
In this section we will compare results of measurements in phantoms, both for imaging
and for radiation absorbed dose, obtained with polychromatic and monochromatic X-ray
beams, whose peculiar characteristics have been highlighted in previous paragraphs. The
comparison is qualitative because it is two systems with totally different characteristics.
The following are the observations related to various parameters:
107
Image quality
Considering the profiles for monochromatic beam (fig. 2.12, 2.13, 2.20) and those
for polychromatic beam (fig. 2.47, 2.48), and the values of the CNR in the table
2.7 and 2.25, it is clear that the details in the first case (monochromatic beam),
have higher contrast. However also the details in the case of polychromatic beam
are clearly visible.
Material Phantom1
insert
Detail
size (mm)
CNRCBBCT Phantom3
insert
Detail
size (mm)
CNRSR
CaCO3 B 1 14 A 1 58
Nylon+air B 2 10 A 2 5.4
Air A 2 14 B 2 62
Air B 4 16 A 4 23
CaCO3 A 4 26 B(Eggshell) 8 122
Animal fat B 8 2.4 B 8 15
Olive oil A 8 2.7 B 4 17
Table 2.31: CNR data calculated with polychromatic beam (CNRCBBCT) and monochromatic beam (CNRSR). It is
evident that the details in the case of monochromatic beam have higher contrast so they are more visible.
Microcalcifications visibility
Images in fig. 2.25-2.29 and the processed data in table 2.9 (SNR and CNR) show
that all microcalcifications are perfectly visible at all monochromatic beam
energies; while in the case of polychromatic beam (fig. 2.50-2.52 and table 2.26),
the microcalcifications are just visible and are not well defined.
Imaging artifacts
There are many different types of CT artifacts which degrade the detail contrast
visibility: beam hardening for polychromatic X-ray source, streak artifacts and rings
for monochromatic and polychromatic X-ray source. To minimize the beam
hardening effect, a number of techniques can be used including filtration, calibration
correction and beam hardening correction software. In fig. 2.10, 2.14 and 2.18 rings
and dark streaks, uncorrected artifacts, are shown; while in figure 2.43-2.46 rings in
the CT slices are less obvious.
Spatial resolution
The spatial resolution of CT scans at 28 keV with monochromatic X-ray beam,
considering the hole of 1mm (profile fig. 2.12), and the corresponding FWHM (table
108
2.6), is of 0.41mm. Instead the spatial resolution of CT scans at 50 kVp
(corresponding to an effective energy of about 28.7 keV) with polychromatic X-ray
beam, considering the hole of 1mm (profile fig. 2.46), and the corresponding FWHM
(table 2.24), is of 0.63mm. The spatial resolution of the system with monochromatic
beam is therefore better than a system with polychromatic beam, in the given
experimental conditions.
Acquisition geometry
A synchrotron radiation imaging setup implies a laminar, monochromatic and tunable
beam, in our case at a distance of about 23 m from the source. These characteristics
of the beam allow respectively to: reduce scattered radiation, increase the image
quality; remove beam hardening artifacts and select the energy according to the organ
thickness and composition determining an important reduction in the delivered dose.
An important observation about the CT geometry with synchrotron radiation is the
following: since the beam is laminar, to irradiate the whole volume of the phantom it
is necessary to perform a linear scan with steps of the same size as the beam
transverse size, of 2 or 3 mm in our case. When the beam moves, since the beam
profile is not flat, there is an area that is irradiated twice, i.e. a double absorbed
radiation dose (fig. 2.37). Also, to perform CT acquisitions a rotating bed is needed
since the source is fixed in space. This is a disadvantage for the patient comfort, the
acquisition time and image quality because it could arise motion artifacts.
Instead the CBBCT uses an half cone-beam geometry. This geometry requires only
the rotation of the gantry around the breast to acquire the data for reconstructing the
entire breast volume. In addition, both the detector and the X-ray tube are positioned
close to the depression in patient table: in this way the entire area of interest is
available (from chest wall to nipple), with the possibility of choosing between
different fields of view (FOV), in relation to the size of the region to be examined.
Radiation dose
Comparing the values of the normalized dose ratio, calculated as the ratio of the
charge collected in the six positions of the TLD with respect to the position Axbot
,
both monochromatic (2.16-2.20) and polychromatic beams (table 2.29-2.30 and fig.
2.54), it is clear that: at similar effective energies, the normalized dose and the
109
absorbed radiation dose is always greater in the case of monochromatic beam with
respect to the polychromatic beam.
E = 34keV 32keV 30keV 28keV 24keV
PERtop 164 163 185 229 334
PERmid 171 161 160 224 304
PERbot 151 153 145 212 274
AXtop 151 152 150 204 243
AXmid 120 113 117 145 148
AXbot 100 100 100 100 100
Table 2.32: Normalized dose ratio values in percent, respect to Axbot, as a function of energy for monochromatic beam.
E = 33.5keV 31.9keV 30.5keV 28.7keV
PERtop 132 140 114 181
PERmid 141 125 148 182
PERbot 139 144 166 164
AXtop 119 120 140 163
AXmid - 111 115 131
AXbot 100 100 100 100
Table 2.33: Normalized dose ratio values in percent, respect to Axbot, as a function of energy, for polychromatic beam.
110
Conclusions
In this thesis is presented an investigation, in terms of data analyses, on the problem of
X-ray imaging dedicated to the breast and absorbed radiation dose, with two different
technologies: CT with X-ray beams in cone-beam geometry and CT with synchrotron
radiation in parallel beam geometry. Having access to the experimental data from the
beamline at ELETTRA, with synchrotron radiation, measurements on phantoms at
different energies of the incident beam (from 34 to 20 keV) have been processed with
analysis and reconstruction software ImageJ and Cobra respectively. Various figures of
merit for the image quality, such as noise, CNR (contrast to noise ratio), SNR (signal to
noise ratio), COV (coefficient of variation) have been evaluated. The distribution of the
absorbed radiation dose by the TLD, placed in six different holes at midplane in the
phantom, has been evaluated. The processed data have been compared with published
data by the medical physics at this Dept. of Physics, with the prototype “cone-beam
breast computed tomography” (CBBCT) with a polychromatic beam. The result of this
comparison, reported in the last paragraphs of this thesis, reveals the peculiarities of both
technologies. In particular, the comparison between the two technologies with respect to
image quality, microcalcifications visibility, imaging artifacts, spatial resolution,
acquisition geometry and radiation dose, evidences that the synchrotron radiation
provides a better image quality, with a high contrast of the details, and a very good
visibility of the microcalcifications, at a high absorbed dose (up to 28 mGy). Since the
synchrotron radiation beam is laminar, to irradiate the whole volume of the phantom it is
necessary to perform a linear scan of the phantom height at 2-3 mm steps. When the
beam moves, since the beam profile is not flat, there is a side area in each step that is
irradiated twice, i.e. a double absorbed radiation dose; moreover to perform CT
acquisitions a rotating bed is needed as the source is fixed in space. This is a
disadvantage for the patient comfort, the acquisition time and image quality and it could
give rise to motion artifacts. On the other hand the cone beam geometry provides an
adequate image quality, with good detail contrast, with a lower dose (up to 13.1 mGy)
than that measured with synchrotron radiation. In addition, the patient lies prone on the
table, which is fixed, with the breast pendant, while the gantry with the X-ray tube and
flat panel detector rotate under the bed. In addition, both the detector and the X-ray tube
are positioned as close to the bottom of the patient table, in this way the entire area of
111
interest is covered (from chest wall to nipple), with the possibility of choosing between
different fields of view (FOV), in relation to the size of the region to be examined. This
study confirms the literature data regarding the intrinsic high quality of CT images
acquired with a synchrotron radiation beam, as well as confirming the practical
advantages of the breast CT exam with an under-table gantry setup and X-ray tubes, for
use in the clinical environment.
112
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Acknowledgements
At the end of this work I want to thank, first of all, my supervisors, prof. Paolo Russo and
prof. Giovanni Mettivier for their availability and their fundamental contribution to the
development of this thesis. Thanks to them I developed the desire to learn to do better
and not to be discouraged by the difficulties but always try, because “how hard it was
does not matter, the result is what counts!”
To Francesca, for advancing and supporting me while elaborating my thesis.
To M. Luisa, she spent time studing with me and now I know I have found a true friend
and I can always count on her.
To Magda and Maddalena, always available to listen to me anytime I needed some help.
Also I would like to thank prof.ssa Ester Piedipalumbo and prof. Salvatore Solimeno.
They encouraged me to overcome my shyness.
And now I want to thank Agnese, my mum for all sacrifices she has made to bring me up
and for the patience and understanding she showed me.
To my brothers and sisters, Luigi, Teresa, Gilda and Giovanni, who patiently support me.
I am not an easy person to deal with! And also newcomer, the little Alessia, who has
brought so much joy into our lives.
To Ovidio, my dear dad who cares about me and always protect me from above. Thank
you for the lessons of life you gave to me, even though our time together was so short!
To Gilda and Tammaro, my grandparents. Especially, thanks to my grandmother Gilda
for living with me for six years, she treated me like a daughter.
Finally, I’d like to thank my boyfriend, Mario. Since we met, he has filled my heart and
my life with joy. Thanks for trust, encouragement, understanding, attention and
protection that you show me, day by day.
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