Università degli Studi di Napoli Federico II...Corso di Laurea Triennale in Fisica TESI DI LAUREA...

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

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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...

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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

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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

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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

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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

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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

http://www.nih.gov/

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


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


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







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


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


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


), of six TLDs exposed at 24 keV with aperture beam of 2mm.................85



), of six TLDs exposed at 24 keV with aperture beam of 3mm………….85


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





), of six TLDs exposed at 30 keV………...................................................89






, 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


, as a function of energy for

monochromatic beam……………………………………………………………...………….109


, 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

http://history.nih.gov/exhibits/genetics/sect2.htm

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.

http://www.nih.gov/

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.

http://www.nih.gov/

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


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


(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



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


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


, 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.


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.


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



89


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




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




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



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

).


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.



98





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.




also indicated.

103




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.


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.


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.

Università degli Studi di Napoli Federico II...Corso di Laurea Triennale in Fisica TESI DI LAUREA...

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