Md Shahadat Hossain Akram1, Takayuki Obata1, Mikio Suga2, Fumihiko Nishikido1, Eiji Yoshida1, and Taiga Yamaya1
1National Institute of Radiological Sciences, Chiba, Japan, 2Chiba University, Chiba, Japan
Synopsis
Simultaneous PET/MRI system has attracted much because of its both
functional and anatomic imaging capability. In our laboratory, we have
developed a human head-size PET/RF-coil integrated modality to be used with
existing clinical 3T MRI system (Siemens Magnetom Verio). Eight PET detector
modules are integrated with a cylindrical 8-element Birdcage RF coil for
simultaneous PET and MRI imaging. In the design each detector has been
integrated in between two coil elements.
RF interference to PET detector circuits affects
PET performance. Also noise generated from PET circuits affects MR image
quality. For proper simultaneous operation, PET circuits in each
detector-module were installed inside a copper-shielded box. But shielding
materials very close to RF coil elements and close to imaging region adversely
affects MR imaging quality. In this study we performed Bo and SAR calculation of our hybrid system. Though there have noticeable changes in the Bo values, the SAR remains very low.Introduction
Dedicated PET inserts
for the existing MRI systems have got a great deal of interest in the recent
years. PET inserts for the brain imaging has got the most attention over the
last few years [1-3]. A head size PET insert have several benefits both in
terms of cost and image quality: (a) it will be close to the imaging region -
means better sensitivity and image resolution; (b) with fewer PET detectors and
electronics - means lower cost than the larger diameter body PET ring. In this
study we took the challenge of inserting the PET detectors in between the 8
elements of a Tx/Rx Birdcage coil in a plan to make the PET detectors closer to
the imaging region. In a previous report in [2], we published the performance
of our concept-prototype for 2 oppositely positioned PET modules. Here we have
developed a complete cylindrical PET ring with 8 PET detectors integrated with
the RF coil.
Methods and Materials
Corresponding electronic circuits of eight detectors were
installed inside copper shielded rectangular cubic boxes with the scintillation
crystal outside the shield boxes and each box is positioned in between 2 coil
elements (rungs). This way the PET detectors were closer to the imaging region
with the inner diameter of 26 cm whereas the coil diameter is 27 cm (Fig. 1a-b, Fig. 2). To avoid parallex error of PET modules and to
get reasonable image quality even near the PET detectors, we have
implemented depth-of-interaction (DOI) detectors with multi-pixel photon
counter (MPPCs) and application specific integrated circuits (ASICs).
For the scintillation crystal, we implemented 4-layer
depth-of-interaction (DOI) LGSO (Lu1.8Gd0.2SiO5:Ce, Hitachi Chemical Co., Ltd., Japan) detectors.
The scintillators were arranged as 12X4X4 matrix (Fig. 1(a))
block in which the dimension of each crystal cube was 2.9 mmX2.9 mmX2.9 mm and the gap
between the crystal blocks were filled up with 0.065 mm thick multilayer
polymer mirror of 98% reflectivity (Sumitomo 3M, Ltd.). The axial FOV of the
PET detectors was 11.73 mm. Two such sets of scintillator detectors were
arranged side-by-side as shown in Fig. 1(c-d), and 6 MPPCs (3 MPPCs for each
block of scintillators; S11064-050P, Hamamatsu Photonics, K.K.) were connected
with DOI scintillator sets to readout the PET signals. Each MPPC of dimension 3
mm X 3mm had 44 readout channels. A
weighted sum circuit (WSC) concept was implemented to reduce the total 96
readout channels of 6 MPPCs into 4 channels and amplifier circuits were
implemented with the readout circuit board.
The readout signals were collected in the data acquisition system outside the
MRI magnet room by 10-meter long coaxial cables.
Experiments were conducted in a 3 T clinical MRI system
(Siemens Magnetom Verio) by using a homogeneous nickel chloride solution
phantom of diameter 20 cm and axial length 11 cm.
Bo map was generated by the phase difference of two images
with two different echo times - 30 ms and 35
ms [5].
Following the electric properties tomography (EPT) method,
we calculated SAR from the B1 field distribution [6]. Also for the B1
field measurements, we implemented double angle method [7] with the following
sequence parameters: TR = 3 s, TE = 12 ms, Slice thickness
(ST) = 5 mm, 128 X 128 matrix, FOV = 300 mm. For the two images the
excitation flip angles were 60o and 120o
and, refocusing flip angles were 180o for both.
Results and Discussion
Fig. 3 illustrates
the B
o map for the central slices. Units are given in microT. For comparison
purpose we have given B
o maps for both the cases of RF coil with and
without PET modules. It is clearly seen that due to PET module Bo
values increase.
At the center slice the
inhomogeneity increases from 0.93 ppm without PET to 2.2 ppm due to PET detectors and, at 5
cm off-center it increases from 2.5 ppm to 4.1 ppm. The SAR maps (Fig. 5) for both transverse and coronal plane
were calculated from the B
1 distribution maps (Fig. 4). The value of local SAR
and average SAR was seen much smaller than maximum limit.
Change in SAR value for
PET detectors were very negligible (from .0021 W/Kg to .0027 W/Kg at the center
slice with spin-echo imaging for 3 minutes). We have found a significant change in the B
o field. This can be because of the ferromagnetic elements on the PET circuits and eddy currents in the shielding materials. We are working to further improve the system in near future.
Acknowledgements
No acknowledgement found.References
[1] A. Kolb, et al, Eur
Radiol 22 (2012) 1776–1788; [2] F. Nishikido, et al, Nucl Instru Method Phys
Res A 756 (2014) 6–13; [3] P. Olcott, et al, Phys Med Biol 60 (2015) 3459–3478; [4] M. C. Leifer, Magn Reson Med, 38 (1997)
726-732;
[5] AAPM Report No. 34,
Medical Physics, 19(1992)217-229.[6] U. Katscher, et al, IEEE Trans Med Imag, 28 (2009) 1365-1374. [7] E.K.Insko et al, JMR
103(1993)82-85.