Download V. Images and Results with Si

Characterization of a Dual Energy
Angiographic System
with CCD+FOS and Si-strips Detectors
G. Baldazzi Member, IEEE, D. Bollini, A.E. Cabal Rodriguez, M. Gambaccini, M. Gombia, P. Grybos
J. Lopez Gaitan, G. Pancaldi, F. Prino, L. Ramello, L. Roma, P. L. Rossi,
A. Sarnelli, K. Swientek, A. Taibi, A. Tuffanelli, M. Zuffa

We have investigated a non-conventional
angiographic imaging methodology called Dual Energy
Angiography (DEA) based on a quasi-monochromatic X-rays
source.
An experimental DEA apparatus was developed. The Bragg
monochromator, mounted on a standard W-anode X-ray tube,
generates two thin parallel beams with energies peaked before (at
EL = 31.1 keV) and after (at EH = 35.3 keV) the iodine K-edge (EK
= 33.2 keV). Images result as the difference between the
logarithms of the transmitted intensities of the two beams.
A dedicated dynamic phantom, simulating different tissues
absorption and containing calibrated vessels, was scanned under
the monochromator slits by a PC-controlled scanning system.
Two types of detectors were used and tested: a linear CCD
coupled with FOS and a Si-strips detector with energy window
discriminator.
In this work resulting imaging capabilities of the DEA
experimental system were compared with those of a commercial
SIEMENS ANGIOSTAR digital subtraction angiographic
apparatus.
Abstract--
I.
INTRODUCTION
Digital Subtraction Angiographic technique (DSA) is one of
the most invasive and risky radiodiagnostic procedures [1]. In
fact, the long X-ray fluoroscopy time – needed to verify the
insertion and site of catheters and wire guides, the injection of
the iodine contrast medium (with a concentration of 320370
G. Baldazzi, D. Bollini and M. Gombia are with the University of Bologna,
Dep.t of Physics, viale Berti Pichat 6/2, 40127 Bologna, Italy (telephone: +39051-2095152, e-mail: [email protected]).
G. Baldazzi, D. Bollini, G. Pancaldi and M. Zuffa are with the INFN
Section of Bologna, viale Berti Pichat 6/2, 40127 Bologna, Italy.
A.E. Cabal Rodriguez is with CEADEN, Havana, Cuba.
M. Gambaccini and A. Sarnelli are with the University of Ferrara, Dep.t of
Physics, via Paradiso 12, 44100 Ferrara, Italy.
P. Grybos and K. Swientek are with the AGH University of Science and
Technology, Dep.t of Nuclear Electronics, Krakow, Poland.
J. Lopez Gaitan is with the Universidad de los Andes, Bogota, Colombia.
F. Prino and L. Ramello are with The University of Piemonte Orientale
Dep.t of Science and Advanced Technologies and INFN, via Bellini 25/G,
15100 Alessandria, Italy.
L. Roma is with the University Hospital S. Orsola, via Massarenti 9,
40138 Bologna, Italy.
P. L. Rossi is with the University of Bologna Health Physics Service, via
Massarenti 9, 40138 Bologna, Italy.
mgI/ml) and the sequences of radiological frames could
represent a risk for patient [2].
In particular, visualization of small vessels, as in pediatric
patients, needs higher iodine concentration (370 mgI/ml) and
longer fluoroscopy time (20 minutes or more), that could
determine very high dose values (also Entrance Skin Dose
values of 1 Gy or more) if added to the necessity of repeating
X-rays exposure for multi-projection diagnostic images.
Because traditional DSA – also known as “temporal DSA” –
is based on logarithmic subtraction of images acquired before
and after contrast medium injection, only the part of the X-rays
spectrum with energy surrounding the iodine K-edge value
(E=33.17 keV) produces the angiographic image. So, large part
of the polychromatic spectrum contributes only to patient dose
or image penumbra [3].
Different imaging systems based on added filtration of the
primary beam were proposed [4] but difficulties involving the
reduction of the Signal to Noise Ratio (SNR) represent a limit
to the clinical application of these techniques.
Recently, the application of the DEA technique with high
intensity highly monochromatic synchrotron radiation
[5],[6],[7],[8],[9],[10],[11], generated angiographic images
with reduction of contrast medium concentration and patient
dose, but cannot represent a practical solution for clinical
applications.
To overcome these fundamental limitations, we have
developed and tested a new imaging system [12] based on the
Bragg diffraction of the primary polychromatic X-ray beam on
a highly oriented pyrolitic graphite mosaic crystal.
The experimental DEA apparatus utilize a custom Bragg
monochromator fitted on a W-anode X-ray tube and produces
two thin (1504 mm2) and quasi-monochromatic parallel
beams. The detector system was constituted by two
HAMAMATSU S7199 linear image sensors CCD coupled
with Fiber Optic plate with CsI(Tl) Scintillator (FOS) mounted
on a micrometric motion system. Phantoms were scanned
between the monochromator and the detector and an adapted
version of the logarithmic subtraction, that we have called
hybrid subtraction, provide the enhanced image of the vessels
morphology.
A second type of detector was used. It was a Si strips linear
detector [13], [14]. This detector has spectroscopic capabilities
and permits to work in an energetic window to reduce
scattering noise.
On the other hand, the difference between the intensities of
the two beams, the energetic shift (that is the difference in
energy measured along a transversal section of each slit) and
intrinsic properties of the detectors could represent limitations
of the DEA system.
For this reasons, the aim of this work is to characterize and
evaluate the performance of our DEA apparatus, comparing the
image results with a professional DSA angiograph used in
clinical applications.
II. MATERIALS AND METHODS
A. X-ray source: the Monochromator
The polychromatic beam generated by a conventional X-ray
tube impinge on a 002-pyrolytic graphite mosaic crystal
(28601 mm3 with certified mosaic spread 0.26) with an
incidence angle 3.2. This is the Bragg angle [15] needed
to select diffracted beams with peaks energies lower and higher
of the iodine K-edge by means of two slit collimators [16].
The choice of the mosaic graphite crystal was made to gain
in diffracted photonic intensities [17], to have spectra with
sufficiently great fluence and a FWHM of about 3 keV.
Fig. 1. Spectrometric analysis of the beams along coronal plane
(perpendicular to the slits): spectrua was acquired using a nitrogen cooled
HPGe detector with a collimator of 0.21 mm2 translated under the slits.
In order to define the energetic properties of the beams as
well as the photonic fluxes, a spectrometric analysis was
performed. Figure 1 shows the shape of the quasimonochromatic energetic beams. Their properties are
summarized in Table I. According to Zachariasen theory for
the Bragg diffraction on graphite mosaic crystal [18], the
photonic flux of the Low Energy Beam (LEB  EL = 31.1
keV) is roughly 3 times more intense of the High Energy Beam
(HEB  EH = 35.3 keV).
Table I: Beams Properties @ 70 kV, 25 mA anode voltage and current,
25 mm slit-detector distance.
LEB
HEB
Mean Energy
(Peak Centroid in keV)
31.1
35.3
FWHM
(keV)
2.3
2.3
Mean Photon Flux
(photons×mm-2×mA-1s-1)
1.9E+04
5.2E+03
B. The CCD detector
The detector, produced by HAMAMATSU, is made with two
CCD chips mounted side to side closely (with only 130200
m of dead space) and coupled with CsI(Tl) FOS. The CCD
pixel size is 4848 m2 and the overall detector size is
15361282 square pixels with an active area of
147.4566.144 mm2 (plus the dead space between the chips).
CsI(Tl) scintillator is grown as needle-shaped crystals, each
having a diameter of about 2 m to increase internal
conversion efficiency between X-rays and light photons The
CCDs were mounted inside a vacuum proof case of Cu with the
dedicated electronics.
C. The Si-strips detector
The detector consists of two one-dimensional silicon array (to
cover contemporarily both beams) with photon counting
capabilities. A double energy threshold is implemented to
exclude scattered radiation. Any array has 384 silicon
microstrips of 100 m pitch and 300 m Si-wafer thickness.
Each array is equipped with 6 RX64 ASICs. The ASIC
includes a charge preamplifier, a shaper, a discriminator and a
20-bit counter for each of its 64 channels [14].
D. Dynamic phantom
The test object used to evaluate imaging properties of the
systems is shown in Figure 2. It consists of 5 different epoxy
blocks with catheters of different diameters: (2.000.05) mm,
(1.800.05) mm, (1.600.05) mm, (1.000.05) mm,
(0.800.05) mm and (0.400.05) mm.
This phantom has a particular design that can confuse the
Automatic Exposure Control (AEC) always active in the
Siemens angiograph. So, the epoxy blocks are surrounded by a
lead layer that masks the X-rays excess. On the same support, a
bar pattern (for modulation transfer function analysis) and
inserts with high and low contrast were embedded in order to
obtain a multi-purpose object to test the physical and imaging
properties of every angiographic system.
Catheters are filled automatically by a pump: this feature
permits to analyze the vessels perfusion as well as to acquire
mask image – in DSA – and subtracting images without
phantom movements to eliminate image artifacts.
III. DEA IMAGE PROCESSING
A. DEA imaging
Two images (LEB and HEB images) were recorded scanning
the dynamic phantom along the orthogonal direction of the
beams. Images are composed of N frames, each one with 25
rows of 3072 pixels. The image matrices contain (N25)3072
numerical data with 12 bits of dynamical range.




SI (i, j )   ln HEB final (i, j )   ln LEB final (i, j ) ,
where i  1,...,25  N j  1,...,3072.
(1)
This allows to reduce background noise in subtracted image
and to enhance iodine contrast.
IV. IMAGES AND RESULTS WITH CCD DETECTOR
Water
Bone
Phantom 1
Phantom 2
Phantom 3
Phantom 4
Phantom 5
Lin. Abs. Coeff. (1/cm)
2.5
Energy
range of
interest
for
D.E.A.
2.0
1.5
1.0
In order to evaluate and compare the image properties, we
have collected images of the dynamic phantom with our D.E.A.
system and with a commercial D.S.A. apparatus, filling
catheters with different iodine concentration (starting from 185
mgI/ml – i.e. ½ of typical iodine concentration injected in
pediatric patient during an angiographic procedure - to 20
mgI/ml).
0.5
0.0
20
30
40
50
60
70
80
Energy (keV)
Fig. 3.a: Images of the dynamic phantom filled with 185 mgI/ml. On the
left, DEA image compared with the DSA image on the right. The Phantom 1
is on the top, the Phantom 5 on the bottom.
Fig. 2: Every block of the dynamic phantom was characterized in terms of
linear absorption coefficient  (1/cm) with a X-ray monochromatic source.
The first block on the left is the phantom 1 and so on.
Images are weighted by means of a flat field process, to
correct the different efficiency and dark current of each CCD
pixels and are normalized for X-ray tube fluctuations by
reading the counts of a custom miniaturized solid-state
exposure meter fitted inside the monochromator.
B. Hybrid subtraction
Since the energy spread between the two quasimonochromatic beams is of 4 keV, it is not possible to apply
the classic logarithmic subtraction algorithm between LEB and
HEB images because of the energy dependence of the
absorption coefficients. If applied to DEA, the classic DSA
algorithm produce an incomplete background contrast
cancellation, reducing visibility of vessels filled by contrast
medium.
For this reason, we have implemented a hybrid algorithm that
weight HEB and LEB images by using two parameters (and
) correlated to the differences of  between tissues:
Fig. 3.b: The same as 3.a but with 95 mgI/ml.
For each image and for each block of the dynamic phantom,
an average contrast profile of the vessels was calculated in
order to define the sensitivity of the systems. In particular, for
every block and for every iodine concentration was selected a
ROI as great as the block analyzed (i.e., to minimize statistical
noise fluctuations), was measured 20 transversal profiles inside
the ROI and was evaluated the average contrast profile as the
mean of these 20 profiles.
Profile (Grey Level)
180
160
140
120
100
80
60
40
0
200
400
600
800
1000
1200
Distance (pixel)
Fig. 3.c: The same as 3.a but with 46 mgI/ml.
Fig. 4.b: Example of average contrast profiles evaluated on images. These are
DEA profile for the phantom 4. Different curves represent profiles obtained
with a different iodine concentration (from 185 mgI/ml to 20 mgI/ml). The
“overshoots” on the catheter edges are due to the convolution with a nonoptimized filter. Same graphs were calculated for every block obtaining
similar shapes.
100%
90%
80%
MTF
70%
60%
50%
40%
DEA
30%
20%
DSA
10%
Fig. 3.d: The same as 3.a but with 20 mgI/ml.
0%
0
1
200
3
4
5
180
Fig. 5: Modulation transfer Function (MTF) of the DSA angiograph in
comparison with the one of the experimental DEA apparatus.
160
140
120
V. IMAGES AND RESULTS WITH SI-STRIPS DETECTOR
100
80
60
40
0
50
100
150
200
250
300
350
Pixel #
Fig. 4.a: Example of average contrast profiles evaluated on images. These are
DSA profile for the phantom 4. Different curves represent profiles obtained
with a different iodine concentration (from 185 mgI/ml to 20 mgI/ml). Same
graphs were calculated for every block obtaining similar shapes.
Because of the large time required to scan an image with the
thin detectors (300 m) we have acquired only a reduced
number of catheters. Figure 6 shows subtracted images of the
only first tree vessels filled with growing concentrations of
iodine.
Relative intensity
300
0.0
-0.2
300 um pixel
Profile (Grey Level)
2
Frequency (lp/mm)
-0.4
200
-0.6
-0.8
Pixels
-1.0
100
0
100
200
300
0
0
20
40
60
80
100 um pixel
100
120
Fig. 6.a: The Phantom 1 filled with 185 mgI/ml. On the left the intensity
profile.
[3]
300
100 um pixel
[4]
200
[5]
100
[6]
0
0
20
40
60
80
300 um pixel
100
120
140
[7]
Fig. 6.b: The Phantom 1 filled with 95 mgI/ml.
0
20
40
60
80
100
120
140
[8]
300
300
100 um pixel
[9]
200
200
100
100
[10]
0
0
0
20
40
60
80
300 um pixel
100
120
[11]
140
Fig. 6.c: The Phantom 1 filled with 20 mgI/ml.
[12]
VI. DISCUSSION AND CONCLUSIONS
Fig. 4 shows the average profiles for the block labeled
phantom 4 when iodine concentration decreases. The DEA
system can visualize catheters much smaller than those seen by
the conventional angiograph. Smaller catheter is never seen by
DSA. But this may be related to its less spatial resolution (Fig.
5).
The more important thing is the better capability of DSA to
distinguish catheters filled with very small concentration of
iodine.
Besides, noise in CCD images is much higher then in DSA
ones. This is due to various elements: a) the lower photon
fluence and consequently the worst statistics, b) the mechanical
movements required for DEA, b) the X-rays energy shift in the
coronal direction, d) the post-processing algorithm
implemented on DSA are not yet well optimized.
In Si-strips images, the low fluence feels even more but
images are, in complex, less noisy for respect to the CCD
detector.
We believe that a more accurate filtering process and the use
of powerful RX tube may overcome the major problems of this
new technique.
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