J. Biomedical Science and Engineering, 2010, 3, 1085-1092 JBiSE
doi:10.4236/jbise.2010.311141 Published Online November 2010 (http://www.SciRP.org/journal/jbise/).
Published Online November 2010 in SciRes. http://www.scirp.org/journal/jbise
A 4-channel 3 Tesla phased array receive coil for awake rhesus
monkey fMRI and diffusion MRI experiments
Mark Haig Khachaturian1,2
1Athinoula A. Martinos Center for Biomedical Imaging, Massachusetts General Hospital, Harvard Medical School, Charlestown,
USA;
2Department of Nuclear Science and Engineering, Massachusetts Institute of Technology, Cambridge, USA.
Received 6 October 2010; revised 18 October 2010; accepted 20 October 2010.
ABSTRACT
Awake monkey fMRI and diffusion MRI combined
with conventional neuroscience techniques has the
potential to study the structural and functional neu-
ral network. The majority of monkey fMRI and dif-
fusion MRI experiments are performed with single
coils which suffer from severe EPI distortions which
limit resolution. By constructing phased array coils
for monkey MRI studies, gains in SNR and anatomi-
cal accuracy (i.e., reduction of EPI distortions) can be
achieved using parallel imaging. The major chal-
lenges associated with constructing phased array
coils for monkeys are the variation in head size and
space constraints. Here, we apply phased array tech-
nology to a 4-channel phased array coil capable of
improving the resolution and image quality of full
brain awake monkey fMRI and diffusion MRI ex-
periments. The phased array coil is that can adapt to
different rhesus monkey head sizes (ages 4-8) and fits
in the limited space provided by monkey stereotactic
equipment and provides SNR gains in primary visual
cortex and anatomical accuracy in conjunction with
parallel imaging and improves resolution in fMRI
experiments by a factor of 2 (1.25 mm to 1.0 mm iso-
tropic) and diffusion MRI experiments by a factor of
4 (1.5 mm to 0.9 mm isotropic).
Keywords: 4-Channel; fMRI; RF Coil Design; Monkey;
SNR; G-Factor; Diffusion MRI; Phased Array
1. INTRODUCTION
Awake monkey fMRI combined with conventional neu-
roscience techniques (e.g. electrophysiology, lesion stud-
ies, reversible deactivation) has to potential to under-
stand and map the functional network [1-5]. Also, awake
monkey fMRI can be used to validate diffusion tracto-
graphy methods [6-9] by benchmarking known functional
pathways. The majority of monkey MRI experiments are
performed with single coils [1,3,4]. Though previous
monkey studies have provided valuable information,
single coil full brain fMRI studies suffer from severe
EPI distortions at resolutions higher than 1.25 mm iso-
tropic for fMRI [1,3,4].
DTI studies at 1.5 mm isotropic have provided valuable
information on the the underlying structure of the monkey
brain [10], however considering the thickness of white
matter in the optic radiation is 1 mm, current DTI studies
do not possess the necessary resolution to characterize
the white matter structure throughout the brain. Since the
purpose of DTI is to measure the “structural connec-
tivity” between regions of the brain, resolutions of < 1
mm would be desirable. Also, more advanced diffusion
reconstruction techniques which are aimed at resolving
multiple fiber orientations can benefit from any SNR and
EPI distortion reduction [11-13]. Phased array coils in
conjunction with parallel imaging could provide gains in
SNR and anatomical accuracy (i.e., reduction of EPI
distortions) in order to achieve the resolutions of 1 mm
[14-17]. This would greatly aid fMRI and DTI studies to
map functional and structural connectivity [18,19].
The major challenges associated with constructing
phased array coils for monkeys are the variation in head
size and space constraints. Phased array technology is
based on placing the multiple receive channels as close
to the head as possible [20]. Therefore, the natural varia-
tion in monkey head size makes rigid coils impractical.
Another major concern is the space constraints associ-
ated with awake monkey fMRI experiments (head posts,
recording wells etc.). The monkeys is confined in a chair
with its headpost secured to the outside of the chair [3].
This leaves very little room for RF equipment and ren-
ders bird-cage coils ineffective and makes it difficult to
cover to the entire monkey brain with coils. Therefore,

M. H. Khachaturian / J. Biomedical Science and Engineering 3 (2010) 1085-1092
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large SNR gains with phased array coils are localized to
surface cortex and will not be present through the brain
as in human imaging. Finally, phased array coils for
awake money fMRI studies cannot interfere with the
vision of the monkey.
Here, we apply phased array technology to a 4-channel
phased array coil design which satisfies the criteria de-
fined above. The coil improves the resolution and ana-
tomical accuracy of full brain awake monkey fMRI and
diffusion MRI experiments. The phased array coil is
capable of adapting to different rhesus monkey head
sizes (ages 4-8) [15,20]. The methodology described in
this paper can be used in the development of phased ar-
ray coils for other primates and small animals.
2. MATERIALS AND METHODS
2.1. Single Coil-10 cm Transmit/Receive Coil
The large majority of full brain awake monkey fMRI
experiments have been performed with a single transmit/
receive coil [2-4,21]. A single transmit receive coil has
the advantage of being easy to setup and is not affected
by variations in monkeys head size. In order to quantita-
tively compare the benefits of the 4-channel phased array
receive coil, we designed a 10 cm transmit/receive coil.
A 10 cm transmit/receive coil was made from flexible
circuit board material (Dupont Pyralux, Durham, NC)
and glued (Gougeon Corp.-G5 Adhesive Hardener and
Resin Epoxy, Bay City, MI) to a thermoplastic base. The
coil was formed in the shape of a ‘saddle’ to improve the
uniformity of the field compared to a circular coil while
not affecting the vision of the monkey. The conductor
width was 4 mm. The coil had seven capacitors on it and
one variable capacitor (Voltronics, Denville, NJ) The coil
plugged into a single preamplifier (Advanced Receiver,
Burlington, CT). A schematic of the circuitry used in the
transmit/receive coil is shown in Figure 1(a). The physi-
cal shape of the single transmit/receive coil and transmit
only coil was the same.
2.2. Phased Array
2.2.1. 10 cm Transmit Only Coil
It would be ideal to use the body transmit coil of the
NMR scanner in conjunction with all phased array coils
because of its uniform magnetic field. The extent of the
body transmit coil field induces currents in the long re-
ceive cables and causes image artifacts. Long receive
cables are necessary because the preamplifiers must be
located outside the magnet so that they are not affected
by gradient switching. However, with a custom build 10
cm transmit only coil, the field of the transmitter is not
large enough to reach the long cables. Thus, functional
and diffusion MRI experiments are only limited in reso-
lution by the coil properties and not other hardware con-
siderations.
The transmit only coil was constructed in the same
manner as the 10 cm transmit/receive coil. However, a
detuning circuit was added to prevent the transmit coil
from interfering with the phased array receive coil dur-
ing image acquisition (see Figure 1(b)).
(a)
(b)
Figure 1. Circuit schematic of (a) single coil and (b) phased array coil.

M. H. Khachaturian / J. Biomedical Science and Engineering 3 (2010) 1085-1092
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2.2.2. 4-Channel Receive Coil
A model of a monkeys head was made from T1 ana-
tomical images using 3D stereolithagraphy (Medical
Modeling, Golden, CO). A monkey helmet was made
from fiberglass cloth (Bondo, Atlanta, GA) and shaped
to the monkey model. Three layers of fiberglass cloth
were glued together using epoxy to give the monkey
helmet a combination of strength and flexibility. The
four receive coils were made from flexible circuit board
material (Dupont Pyralux, Durham, NC) and glued to the
monkey helmet (Gougeon Corp.-G5 Adhesive Hardener
and Resin Epoxy, Bay City, MI). A standard capacitive
bridge match and PIN diode (MA4P4002B-402; Macom,
Lowell, MA) trap was used to detune the receive coils
[22]. Two coils were overlapped on the left side of the
helmet and two coils were overlapped on the right side.
The coupling, S12, between the overlapping coils was
measured (< −20 dB) before gluing the coil to the mon-
key helmet. The coupling between second nearest
neighbor coils was minimized using preamplifier de-
coupling (see next section). Figure 1(b) shows a circuit
schematic of the phased array and Figure 2 shows the
orientation of the coils relative to the brain. When posi-
tioning close to the monkey, the movement of the jaw
muscles do not affect the stability of the coil.
2.3. Preamplifiers
A preamplifier board was constructed and made portable
to accommodate the different stereotactic apparatus and
monkeys chairs used in awake and anesthetized MRI
experiments. The preamplifier board consists of eight
preamplifiers (Siemens Medical Solutions, Erlangen,
Germany), and four bias lines (±30 V) to tune and detune
the receive and transmit coils. Preamplifier decoupling
was employed using a cable trap with semi rigid coax
(Suhner UT-070 type) to suppress any residual coil cou-
pling from next nearest neighbor interactions [20,23].
2.4. Monkey Data Acquisition
MRI data were acquired from three juvenile male rhesus
(a) (b)
Figure 2. Schematic of the orientation of the placement of four
coils relative to the brain from the (a) superior and (b) sagital
view.
monkeys (macaca mulattas, M1, 5.5 kg, ID #3704, M2,
4.9 kg, ID #0404, M3, 5.7 kg, ID #4505). The data were
acquired on a Siemens Trio 3T MRI scanner located at
the Athinoula A. Martinos Center for Biomedical Imaging,
Massachusetts General Hospital (Charlestown, Massa-
chusetts). All procedures conformed to Massachusetts
General Hospital, Massachusetts Institute of Technology,
and the National Institutes of Health guidelines for the
care and use of laboratory animals (Subcommittee on
Research Animal Care protocol #2003N000338).
The monkey was placed into a magnet compatible ste-
reotactic apparatus (Kopf Instruments, Tujunga, Cali-
fornia) for an anaesthetized experiment to compare the
performance of the single coil to the phased array coil.
Anesthesia was maintained using ketamine and xylazine
(induction 10 and 0.5 mg/kg, i.m., maintenance with
ketamine only). Local anesthetic (lidocaine cream) was
applied to the ends of the ear bars and ophthalmic oint-
ment was applied to the eyelids to minimize discomfort
induced by the stereotactic apparatus. A heating pad was
placed beneath the monkey to keep it warm during the
scan session.
The monkey was placed in a monkey chair (Crist In-
struments, Washington, DC) during awake experiments.
All monkeys were implanted with a headpost (Crist In-
struments, Washington, DC) which was secured to the
monkey chair using two M5 peek plastic screws. Also a
rail system was used to slide the monkey chair in the
magnet to minimize motion. The following sequences
were used to quantify the behavior of the phased array
coil and single coil.
2.5. G-Factor Maps
Proton density gradient-echo images were acquired in
order to calculate g-factor maps of the phased array coil
for horizontal, coronal, and sagital slices [16,20]. Raw
k-space data was obtained with TR/TE/flip = 200 ms/
4.14 ms/20o, 1.5 mm single slice, 128 × 112, and 100 ×
87 mm FOV.
2.6. Coil Sensitivity Maps
Proton density gradient-echo images were also acquired
in order to perform a non-uniform signal normalization
on the T1 anatomical images to improve gray and white
matter contrast and aid in image registration (50 slices,
TR/TE/flip = 1190 ms/3.72 ms/8o, 1.0 mm isotropic, 96
× 96 matrix size).
2.7. SNR Maps
Proton density gradient-echo images (TR/TE/flip 200
ms/3.92 ms/20o, slice thickness 1 mm, matrix 128 × 128,
FOV 128 mm, BW 300) were obtained in the sagital,
axial, and coronal planes for SNR comparison. A noise

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reference scan was acquired by recording an image with
no RF excitation. The image noise matrix, Ψ, was cal-
culated from the receive and sample noise matricies as
described in [16,23] and the SNR was calculated using
the equation below [24] where, S, is the signal vector of
length n, where n is the number of coils (4 in this case).
1T
SNR   SS (1)
2.8. fMRI
Functional single shot echo planar images (EPI) were
acquired at two different resolutions (1.25 mm, 1.0 mm
isotropic) to compare the anatomical accuracy of the
single coil to the phased array coil.
2.8.1. fMRI-1.0 mm BOLD
Fifty horizontal slices were acquired using TE/TR = 26/
3290 ms (phased array coil), 26/5000 ms (single coil),
96 × 84 matrix, 90o flip angle at 1.0 mm isotropic resolu-
tion. The phased array coil had a lower TR because par-
allel imaging [16,17,25] was employed, specifically
generalized auto-calibrating partially parallel acquisi-
tions (GRAPPA, R = 2) [25].
2.8.2. fMRI-1.0 mm MION
Fifty horizontal slices were acquired on M1, M2, and
M3 using TE/TR = 24/2700 ms, 96 × 84 matrix, 90o flip
angle at 1.0 mm isotropic resolution using the phased
array coil. The Siemens AC88 gradient insert was used
(BW = 1305 Hz/voxel) to decrease the TR. GRAPPA with
an acceleration factor of two was employed [25].
2.9. Diffusion MRI
Diffusion MRI was acquired at two different resolutions
(1.25 mm, 0.9 mm isotropic). The 1.25 mm data was
acquired with the single coil and the 0.9 mm data was
used to explore the practical limits of the phased array
coil.
2.9.1. Diffusion MRI-1.25 mm (Single Coil)
The diffusion preparation used a twice-refocused spin
echo [26]. Thirty horizontal slices were taken of M1 at
1.25 mm (0 mm skip). The in-plane resolution was 1.25 ×
1.25 mm, with a matrix size of 74 × 74 (R = 2 was used
for the phased array coil). The sequence parameters were
TE/TR = 85/8300 ms, b = 700 s mm-2. The diffusion
gradient sampling scheme consisted of n = 60 directions
which were obtained using the electrostatic shell method
[27]. Ten images with no diffusion-weighting were also
obtained for a total of 70 acquisitions. The total acquisi-
tion time was 7 min 20 sec. The average of the 10 b = 0
raw images were compared for EPI distortions.
2.9.2. Diffusion MRI–0.9 mm DTI (Phased Array Only)
The diffusion preparation used a twice-refocused spin
echo [26]. Fifty seven horizontal slices were taken of M1
at 0.9 mm (0 mm skip). The in-plane resolution was 0.9 ×
0.9 mm, with a matrix size of 96 × 96. The sequence
parameters were TE/TR = 93/10100 ms, b = 700 s mm-2.
The diffusion gradient sampling scheme consisted of n =
60 directions which were obtained using the electrostatic
shell method [27]. Ten images with no diffusion-wei-
ghting were also obtained for a total of 70 acquisitions.
The total acquisition time was 11 min 48 sec. The diffu-
sion images from five acquisitions were averaged and
reconstructed using the diffusion tensor model [28].
2.10. T1 Anatomical Images
T1 anatomical images were acquired with an MPRAGE
sequence [29] with TR/TI/TE = 1910/1100/3.06 ms, α =
8o, 0.65 mm isotropic resolution, total acquisition time: 6
min 35 sec. The functional (EPI), proton density and
GRE images were registered to the T1 images for com-
parison between the single and phase array coils.
2.11. Image Registration and Visualization
Images were registered using the flirt command (rigid
registration, 6 degrees of freedom) in the FSL toolbox
(http://www.fmrib.ox .ac.uk). All visualization post-pro-
cessing was performed using custom software written in
Matlab (Version 6.5.1.199709 (R13) Service Pack 1).
DTI reconstructions were visualized with custom soft-
ware written in C++ and VTK (Version 4.2) (http://p ublic.
kitware.com/VTK). The tensors were visualized as
color-coded rectangloids.
3. RESULTS
3.1. Coil Properties
The coupling between the coils was measured using an
S12 measurement on a network analyzer. The two sets of
overlapping coils had < −20 dB decoupling between
them. Next nearest neighbor coupling was < −10 dB. In
addition, the preamp decoupling added −20 dB of de-
coupling. Therefore, all coils had < −30 dB of decoup-
ling between them ensuring each coil behaved as a sin-
gle element in the tuned state. PIN diode detuning
achieved > 35 dB of isolation between the tuned and
detuned states.
3.2. SNR
The SNR in the four individual coils of the phased array
is shown in Figure 3 and was calculated using (1).
Figure 4 compares the SNR of the single coil to the
phased array coil. The phased array coil has higher SNR
in surface cortex but has a less uniform profile. Table 1
presents the relative SNR of the phased array coil to the
single coil in regions relevant to fMRI studies.

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Figure 3. (a)-(d) SNR of individual coils.
Figure 4. (Top) SNR of single coil and (Bottom) SNR of
phased array coil. The SNR of the phased array coil is higher
than the single coil in surface cortex although the single coil
has a more uniform SNR profile.
Table 1. Relative SNR in the phased array coil relative to the
single coil in primary visual cortex (V1), the frontal eye fields
(FEF), and the lateral geniculate nucleus (LGN) in the left and
right hemispheres (LH, RH).
SNRPA/SNRSC
Region
LH RH
V1 2.10 2.03
FEF 0.76 0.74
LGN 0.80 0.78
3.3. g-factor
An important property of a phased array coil is its g-factor
map. The inverse g-factor map quantifies how much
SNR is lost in the brain during a parallel acquisition (R >
1) compared to an image with no acceleration, R = 1.
The parallel imaging capabilities of a 4 channel coil
limit the acceleration factor to two, (i.e., R = 2) [16].
Figure 5 presents the inverse g-factor maps for the
phased array coil for an acceleration factor of two in the
different phase encode directions. Less than 15% of the
SNR is lost during an acceleration factor of two in 85%
of the brain. The maximum g-factor of 1.35 (Figure 5(a))
occurs in the horizontal slice with left-right phase en-
coding.
3.4. fMRI
Functional EPI (BOLD) was acquired at 1.0 mm isotropic
resolution. The TR of the phased array coil was much
shorter (3290 ms) compared to that of the single coil
(5000) because parallel imaging (GRAPPA, R = 2) was
employed. Figure 6 presents EPI at 1.0 mm isotropic for
the single and phased array coil. The benefit of parallel
imaging with regards to EPI distortions is apparent
around the edge of the brain. A T1 image is presented for
reference.
Figure 7 presents functional EPI (MION) of M1, M2,
and M3. The images were acquired with the AC88 gra-
dient insert (Siemens, Erlangen, Germany). The SNR
profiles in the brain are relatively consistent between the
monkeys despite the natural variation in head size. M1
has the largest head and thus the signal in the middle of
the brain is the lowest compared to M2 and M3.
3.5. Diffusion MRI
Figure 8 presents the average of 10 T2 images (b = 0)
from a DTI scan from a single coil (1.25 mm) and phased
array coil (R = 2, 0.9 mm). The severe EPI distortions in
(a) (b)
Figure 5. 1/g-factor maps for the (a) left-right horizontal and
the (b) anterior-posterior horizontal phase encode directions for
two-fold GRAPPA acceleration, R = 2. The blue boxes indicate
where the maximum g-factor in the slice is located. In 85% of
the brain, the SNR decrease is less than 15% for two-fold
GRAPPA acceleration. The average SNR decrease in the brain
is ~50% for R = 3 (data not shown). The monkey’s brain is
outlined in black.

M. H. Khachaturian / J. Biomedical Science and Engineering 3 (2010) 1085-1092
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(a) (b) (c)
Figure 6. SNR of (a) single and (b) phased array functional
EPI at 1 mm isotropic resolution. Both images are on the same
scale. The phased array coil images where taken with a
GRAPPA acceleration factor of 2. The phased array image
does not posses the EPI distortions of the single coils are more
accuratley represents the T1 image in panel (c).
(a) (b) (c)
Figure 7. Functional EPI (with MION) acquired with the
phased array coil for (a) M1, (b) M2, and (c) M3 at 1.0 mm
isotropic resolution, TE/TR = 24/2700 ms. M1 has a larger
brain than M2 and M3 resulting in lower SNR in the middle of
the brain. However the SNR in all monkeys is sufficient (> 25)
to measure the MION response throughout cortex.
(a) (b) (c)
Figure 8. (a) Single coil (1.25 mm) and (b) phased array (0.9
mm) b = 0 T2-weighted images (R = 2) for a DTI scan. The
severe EPI distortions in the single coil image make registra-
tion very difficult with a (c) T1 image compared the phased
array coil image.
the single coil image make registration very difficult
with a T1 image. Figure 9 depicts the DTI reconstruction
of a 0.9 mm isotropic acquisition. The tensor map is
overlayed on a T1 with a rigid registration algorithm.
The phased array coil provides a much more accurate
anatomical image such that the rigid registration pre-
serves the tensor maps. ADC and FA maps calculated
from the tensors are presented in panels (b) and (c). The
phased array coil is capable of resolving the ADC and
FA throughout the white matter in the brain.
4. DISCUSSION
The SNR benefit of the phased array coil compared to
the single coil in surface visual cortex, make it ideal for
awake monkey fMRI retinotopic studies (Figure 3, Fig-
ure 4, Table 1). In addition the benefit of using the
phased array coil with parallel imaging improves the
resolution of awake monkey fMRI experiments at 3T by
a factor of 2 (Figure 6) (1.25 mm to 1.0 mm isotropic)
and diffusion MRI experiments by a factor of 4 (Figure
8 and Figure 9) (1.5 mm to 0.9 mm isotropic). The pres-
ervation of the anatomical integrity of the fMRI and dif-
fusion images results in accurate registration with T1
images with a rigid registration algorithm. Thus, func-
tional and diffusion tensor maps can be visualized on a
T1 image making tractography more accurate. During
parallel imaging, an SNR loss of less than 15% is pre-
sent over 85% of the brain with a maximum loss of 26%,
GRAPPA = 2 (Figure 5). The SNR decrease in surface
cortex due to parallel imaging is more than compensated
by the coils higher SNR in that region (i.e., 26% de-
crease during parallel imaging compared to 100% boost
compared to single coil). The coil does not interfere with
the eyes of the monkey or with the headpost because all
of the coils are on the side of the head (Figure 2). This
allows eye tracking to be performed during fMRI experi-
ments. In addition, the flexibility of the fiberglass helmet
allows the coil to easily adapt to different monkeys (Fi-
gure 7).
Though four coils were used in this phased array, future
phased arrays with more coils could be constructed using
the same techniques described in this paper. The general
methodology of making a monkey mold from T1 ana-
tomical images and then constructing a fiberglass helmet,
can be applied to primates and other animals. The appli-
cation of the coil should be taken into account when de-
veloping a phased array coil. For example, in fMRI
studies only interested in retinotopy, it may be more
suitable to use more surface elements and sacrifice the
coils performance in deep cortical structures. Also, for
T1 anatomical studies of rhesus monkeys, a headpost is
not necessary. This would allow one to put a coil on top
of the brain making the phased array coil superior to the

M. H. Khachaturian / J. Biomedical Science and Engineering 3 (2010) 1085-1092
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(a) (c)
Figure 9. (a) DTI, (b) ADC, and (c) FA maps acquired with the phased array coil at 0.9 mm
isotropic resolution. The phased array coil resolves the ADC and FA throughout the brain.
single coil throughout the brain.
5. CONCLUSIONS
A 3 Tesla 4-channel phased array coil was developed to
improve the resolution and anatomical accuracy (i.e.,
reduction of EPI distortions) of awake monkey fMRI
and diffusion MRI experiments. The novel aspect of the
phased array coil is that can adapt to different rhesus
monkey head sizes (ages 4-8) and fits in the limited
space provided by monkey stereotactic equipment. The
phased array coil had a SNR benefit of ~2 on surface
cortex. The phased array coil allowed parallel imaging to
be employed reducing EPI distortions. It improved the
resolution of fMRI studies by a factor of 2 (1.25 mm to 1
mm isotropic). When applied to diffusion MRI studies,
the phased array coil improved the resolution of diffu-
sion MRI images by a factor of 4 (1.5 mm to 0.9 mm
isotropic) compared to single coil studies. The flexibility
of the design allows the coil to adapt to monkeys with
varying head sizes.
6. ACKNOWLEDGEMENTS
This work would not have been possible without the guidance of Gra-
ham Wiggins and Larry Wald. The author is also grateful to Wim Van-
duffel, Hauke Kolster, John T. Arsenault, Leeland Ekstrom, Bechir
Jarraya, Andreas Potthast, Simon Sigalovsky, and Helen Deng for their
assistance with this project. This work was supported by IUAP 5/04,
EF/05/014, FWO G151.04, HFSPO RGY0014/2002-C, GSKE, NINDS
NS46532, NCRR RR14075, NCI CA09502, NCI 5T32CA09502,
GlaxoSmithKline, the Athinoula A. Martinos Foundation, the Mental
Illness and Neuroscience Discovery (MIND) Institute, and the National
Alliance for Medical Image Computing (NAMIC) (NIBIB EB05149)
which is funded through the NIH Roadmap for Medical Research.
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