J. Biomedical Science and Engineering, 2010, 3, 167-180
doi:10.4236/jbise.2010.32022 Published Online February 2010 (http://www.SciRP.org/journal/jbise/
JBiSE
).
Published Online February 2010 in SciRes. http://www.scirp.org/journal/jbise
Design and development of a new dedicated RF sensor for the
MRI of rat brain
Aktham Asfour
Joseph Fourier University of Grenoble, Grenoble, France.
Received 14 October 2009; revised 13 November 2009; accepted 16 November 2009.
ABSTRACT
The design and development of a new dual-frequency
RF probe-head are presented. This probe was ini-
tially dedicated for the MRI of both proton (1H) and
hyperpolarized Xenon-129 (HP 129Xe) in the rat brain
at 2.35 Tesla. It consists of a double-tuned (100 MHz-
27.7 MHz) volume coil, which could be used for both
transmitting and receiving, and of a receive-only sin-
gle-tuned (27.7 MHz) coil. The double-tuned coil con-
sists of two concentric birdcage resonators. The inner
one is a low-pass design and it is tuned to 27.7 MHz,
while the outer one, tuned to 100 MHz, is high-pass.
The receive-only coil is a surface coil which is de-
coupled from the double-tuned volume coil by an
active decoupling circuitry based on the use of PIN
diodes. A home-built Transmit/Receive (T/R) driver
ensures biasing of the PIN diodes in both volume and
surface coils. The original concepts of the design are
addressed, and practical details of realization are
presented. One of the underlying ideas behind this
work is to proceed well beyond the application to the
MRI of HP 129Xe. Actually, this design could be easily
adapted for a large palette of other MRI applications.
Indeed, we tried to make the design versatile, simple
and easy to replicate by other research groups, with a
low-cost, minimum development time and accepted
performances. The prototype was validated at 100
MHz and at 26.4 MHz (sodium-23 resonance fre-
quency at 2.35 T). MRI experiments were performed
using phantoms. In vivo 1H images and 23Na spectra
of the rat brain are also presented.
Keywords: MRI; RF Coils; Double-Tune; Active
Decoupling; Proton; Sodium-23; Hyperpolarized
Xenon-129
1. INTRODUCTION
Nuclear Magnetic Resonance (NMR) and Magnetic
Resonance Imaging (MRI) are important diagnostic and
analytical tools for biomedical studies. During the last
two decades, and since the MRI of hyperpolarized xe-
non-129 (HP 129Xe) was initially demonstrated in 1994
[1] in mouse lungs, the MRI of hyperpolarized gases
(129Xe, 3He…) have become widespread for a large pal-
ette of applications in biology and medicine. Particularly,
the MRI of injected or inhaled HP 129Xe is a promising
noninvasive method for the investigation of brain func-
tion and clinical use. In fact, several studies carried out
in our laboratory and by some research groups in the
world, have already demonstrated that the MRI of HP
129Xe may allow quantitative measurement of the cere-
bral blood flow (CBF) with a high spatial resolution
[2,3,4,5]. This physiological parameter, which is defined
as the blood supply to the brain in a given time and per
mass unit of brain tissue1, is tightly regulated to meet the
brain’s metabolic demands. The measurement of changes
in both global cerebral blood flow (gCBF)2 and regional
cerebral blood flow (rCBF)3 is of great value for func-
tional brain studies as well as for the diagnosis of a large
number of brain diseases (Alzheimer, epilepsy, Parkin-
son, ischemia…). Nevertheless, for such measurements,
especially those of the rCBF, to become quantitatively
reliable by the MRI of HP 129Xe, one must first over-
come several experimental and instrumental challenges.
In addition to the understanding of the behavior of the
129Xe in the brain, which is out of the scope of this paper,
many instrumental obstacles have to be actually resolved.
For example, the experimental set-up for polarization
build-up and storage is relatively complicated and re-
quire some know-how. One important aspect in building
this polarization set-up has already been addressed in our
previous works [6,7]. A second instrumental challenge
concerns the development of specific radiofrequency
(RF) probes for these experiments. In this paper, we deal
with this problem and we present an original and versa-
tile design of a dedicated RF double-tuned volume coil
actively-decoupled from a simple-tuned receive-only RF
surface coil.
1In healthy brain of adult humans, the CBF is regulated to an average of
50 milliliters of blood per 100 grams of brain tissue per minute.
2Value of the CBF measured over all the brain.
3Value of the CBF measured within a s
p
ecific brain re
g
ion.
To help the reader in better understanding the general
context of our research and the goal of this paper, one may

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
168
JBiSE
need to review or to learn some basic principles of a typi-
cal NMR experiment as well as of the NMR of HP 129Xe.
As it is well known, when a sample, consisting of
NMR-sensitive nuclei (1H, 3He, 129Xe; 23Na…etc.), is
subjected to a uniform static magnetic field B0, a net
macroscopic magnetization M of the sample appears in
the direction of B0. This magnetization is proportional,
roughly speaking, to the polarizing field B0, to the den-
sity of the nuclei within the sample and to the character-
istic gyro-magnetic ratio

of the nucleus being studied.
If the sample is subjected now to a short pulse (called
excitation pulse) of a second radiofrequency magnetic
field B1, applied at the characteristic Larmor frequency f0
(about 100 MHz for 1H, 27.7 MHz for 129Xe and 26.4
MHz for 23Na at B0=2.35 Tesla), the net magnetization,
M, may be “tipped” or rotated from its initial direction
(or from its equilibrium state) by an angle his angle is
called “flip angle”. It is proportional to B1 according to
the equation 1 ..B

  where the duration of the RF
pulse is. At the end of the excitation pulse, a resonance
signal (or the NMR signal) at the same frequency, f0, is
received. This signal, which is proportional to the mag-
netization M and to

, is processed to be used for ob-
taining a “fingerprint” of the environment of the nucleus
being studied.
At equilibrium and for a same magnetic field B0, the
NMR signal of a given 129Xe population is about 10000
times lower than the one that could be obtained from the
same volume of protons. This is because of the intrinsic
lower gyro-magnetic ratio and lower density of the xe-
non. Consequently, at equilibrium, the in vivo-NMR
signal of the injected or inhaled 129Xe is not exploitable.
To compensate this limitation, hyperpolarizing tech-
niques have been successfully used to dramatically in-
crease the magnetization of 129Xe before using it for the
in vivo experiments. In this case, the magnetization and
the NMR signal levels are typically enhanced by about
five orders of magnitude.
In NMR/MRI experiments, the well-tuned (to the fre-
quency f0) RF probe has usually a dual role: it ensures
the excitation of the sample (i.e., it creates the RF field
B1) and it receives the weak NMR signal from this same
sample. Usually, both excitation (transmission) and re-
ceiving processes are produced in the same coil. Never-
theless, there are cases where these processes are pro-
duced by two or more different coils. Generally, accord-
ing to considerations related to the spatial homogeneity
of the excitation field B1 and to the receiving sensitivity,
RF coils may be categorized, roughly speaking, as either
volume coils or surface coils. Volume coils create ho-
mogeneous spatial distribution of B1 and guarantee the
possibility to image deep organs. Surface coils allow
high sensitivity, and hence high signal-to-noise ratio
(SNR), for surface regions (the rat brain could be con-
sidered as surface region). These surface coils present
however a poor B1 homogeneity when compared with
volume coils. This homogeneity decreases generally
when one moves away for the centre of the coil.
Both volume and surface coils could be used for both
transmission and receiving. However, when B1 homo-
geneity during the transmission period is required, the
use of a transmit volume coil is necessary. On the other
hand, when this homogeneity is not an issue, a surface
coil may be preferred to optimize the SNR of the de-
tected NMR signals. The association of a volume coil,
for transmission, decoupled from a surface coil, for re-
ceiving, is usually used when both B1 uniformity and
high SNR are required.
For our experiments of measuring the rCBF by the
NMR of HP 129Xe, a dedicated RF probe has to be built.
Actually, the measurement method of the gCBF requires
knowledge of the global (over all the rat brain) flip angle
[2,3,4,5]. This angle depends, remember, on B1. The
determination of a global flip angle can be easily per-
formed. However, when local or regional cerebral blood
flow (rCBF) values (within a specific brain region or
inside a pixel of the MRI image) have to be measured,
the local flip angles (which depend on the local values of
B1) have to be accurately estimated. This estimation is
not a trivial task. This is why it’s generally assumed that
local flip angles are close enough to the global value.
Nevertheless, when a surface coil is used for transmis-
sion, the estimated global flip angle could be very dif-
ferent from the local values due the inherent
non-uniformity of the B1 field distribution. This may
produce inaccurate results in the rCBF values. For ex-
ample, Venkatesh et al. have noted a 20% difference be-
tween the decay times (parameter directly related to the
rCBF) obtained with a surface coil and those obtained
with a volume coil used for both transmission and receiv-
ing [3]. Moreover, since Kilian et al. [2] used a surface
coil in their human brain measurements, it is possible that
their results contain a systematic error due to inaccurate
estimation of the local flip angles. The same problem was
also encountered in our experiments [5]. In all these ex-
periments, the use of surface coils was initially motivated
by the desire to obtain a high SNR in the MRI images.
To overcome this problem, the idea we develop here
consists in using a volume coil for transmitting so as the
global flip angle will be close enough to the local ones
thanks to the high B1 homogeneity guaranteed by this type
of coils. The detection of the NMR signals will then be
performed by a surface coil positioned inside the volume
coil. In this situation the advantage the inherent high sen-
sitivity associated with the surface coil is retained.
The two coils must be decoupled so as when the first
is active, the second must be deactivated. In other words,
during transmission, the volume coil is active or “on”
while the surface coil is “off”. During receiving, the
transmit volume coil has to be deactivated and the sur-

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
169
JBiSE
face one has to be active. Both transmit and receive coils
must be well-tuned to the resonance frequency of 129Xe
(27.7 MHz at 2.35 T).
Moreover, for the rCBF measurements by MRI, the fi-
nal “ideal” probe must satisfy more stringent requirements.
Actually, in addition to the transmission and receiving at
the frequency of 129Xe, one has to be able to perform
1H-MRI experiments (at 100 MHz) without changing the
probe and without moving the sample under test (rat in
our application). In fact, acquiring 1H images is necessary
to localize anatomical structures in the rat brain. These
anatomical images are to be superposed to 129Xe images to
identify structures where inhaled or injected 129Xe could
be localized in the rat brain. On the other hand, the pro-
cedure of homogenization of the static magnetic field
(shimming), which is essential before every MRI acquisi-
tion, must be performed using 1H NMR signals. Actually,
since this shimming procedure requires repetitive excita-
tions and detections, it can not be accomplished on the
129Xe signals because the artificial magnetization (non
equilibrium state) of the HP 129Xe is not renewable when
destroyed by an RF excitation pulse.
In both anatomical image acquisition and shimming
procedures, a high SNR of the detected 1H signals is not
an issue and a same transmit-receive volume coil tuned
to 100 MHz could be adequate. This coil must however
be decoupled from both transmit volume and receive-
only 129Xe coils.
To summarize: the final “ideal” NMR probe that we
developed consists of a transmit volume coil actively
decoupled from a receive-only surface coil working at
27.7 MHz, and a transmit/receive volume coil for the 1H
MRI at 100 MHz. The decoupling between the 1H coil
and the 129Xe coils will be performed geometrically.
This novel configuration, which actually associates
three MRI coils in a same probe, is new and has not been
published. In fact, due to the variety of size and experi-
mental protocols encountered with small-animal imaging
experiments, it is usually advisable and may be neces-
sary to build a probe optimized for each experimental
set-up and for each application. However, it is important
to notice here that another objective of this work is to
proceed well beyond this specific application to the MRI
of HP 129Xe. Actually, the current design could be
adapted or could inspire other designs for a palette of
MRI applications in the rat brain. Generally, for such
applications, the “ideal” imaging/spectroscopy coil
would be one which provides high sensitivity, high B1
homogeneity, and operation on two or more frequencies
without retuning requirement or cables changes. The
problem of sensitivity can be addressed using surface
coils [8]. The homogeneity requirements can be met by
the use of adequate coil geometry (i.e., birdcage design
[9]). The problem of acquiring data from two different
nuclei (usually proton and a second nucleus like 129Xe,
23Na sodium, 31P…) requires the use of an efficient dou-
ble-tuned coil (volume or surface coils) [10,11,12,13,14,
15,16,17]. While these problems have been addressed
individually, there have been no published designs that
have integrated all these features into one design. The
goal of this work is to approach performances as out-
lined above. Another underlying idea is to make this
specific probe easy to replicate, by other research groups
with a low-cost, minimum development time and ac-
cepted performances. To allow for easy extension of this
design to other frequencies (nuclei) and dimensions, we
describe in some details the practical aspects of the
workbench design as well as the coil characterization
and MRI/NMR testing using phantoms. The probe is
validated at the 1H frequency and the sodium frequency
(26.4 MHz at 2.35 T) since this last is very close to the
xenon frequency. In vivo MRI images of the rat brain
were obtained at 1H frequency as well as NMR signals at
the sodium frequency.
All developed circuits and any other details on devel-
oped probe could be obtained by simply writing to au-
thors.
2. MATERIALS AND METHODS
2.1. The Double-Tuned Volume Probe
We built a double-tuned volume probe using the bird-
cage design [9]. This design ensures inherent high B1
homogeneity within the coil region. It is therefore
well-suited for the MRI of HP 129Xe. The design is
shown in its low-pass and high-pass versions in Figure 1.
It consists of two circular loops, referred to as end-rings,
connected by a number of equally spaced straight seg-
ments referred to as legs or rungs. Depending on the
version, the tuning capacitors are placed in the legs
(low-pass) or in the end-rings (high-pass).
A single-tuned birdcage will resonate in several
modes. A birdcage with N legs has N/2 resonance modes
(assuming N is even) [9,17], with the most useful to
MRI being Mode 1 (fundamental mode). In this mode,
the currents in the legs follow a cosine distribution as
one moves radially around the coil. This sinusoidal dis-
tribution of currents is essential to create a very ho-
mogenous transverse magnetic field B1.
For our application, it is necessary to double-tune
the birdcage. The sinusoidal optimum current distribu-
tion should be maintained at each frequency. Double-
tuning a birdcage resonator is not a trivial task and it has
been subject of many publications [10,11,12,13,17]. The
method to choose depends actually on the desired per-
formances at each frequency. Double-tuned birdcages
have been used at relatively low field ( T). Few
designs at higher fields have been proposed. For exam-
ple, Fitzsimmons et al. [10] presented a double-tuned
birdcage coil at 2 T. It was dedicated for the MRI of
proton (85.5 MHz) and the NMR spectroscopy of phos-
5.1

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
170
Each resonator consists of eight legs. This number
represents a compromise between B1 homogeneity and
complexity of realization. The inner coil, a low-pass bir-
dcage, had a diameter of 9 cm, a length of 13 cm, and
was tuned to the xenon frequency. The outer coil, a
high-pass design, had a length of 16 cm, a diameter of
10.5 cm, and was tuned to the proton frequency. We
verified that no proton resonance modes occur at the
xenon frequency. Conversely, no xenon modes occur at
the proton frequency. The inner birdcage was rotated
within the outer birdcage until minimum coupling oc-
curred. This positioned the legs of the inner coil midway
between the legs of the outer one. Also, the outer bird
phorus (34.6 MHz) of the rat abdomen. They derived
their coil in quadrature operation for both proton and
phosphorus channels. The coil was used for both trans-
mission and receiving at each frequency. One of the
originalities of the double-tuned birdcage that we pro-
pose here is that it has to work at higher field (2.35 T)
for the MRI of the rat brain. More important is the fact
that, to the knowledge of the author, a double-tuned
birdcage associated with an actively decoupled receive-
only surface coil has not been published.
The double-tuned birdcage we built consists of two
birdcages in a concentric configuration as it is seen in
Figure 2.
Figure 1. Spatial representation of the low-pass (left) and high-pass (right) birdcage
resonators. CH (high-pass) and CL (low-pass) are the tuning capacitors.
Figure 2. A simplified scheme of the developed imbricate birdcage resonators. The
inner coil is low-pass and the outer one is high-pass. In linear polarization operation,
the coils should be derived in such a way that their B1 fields are orthogonal allowing
efficient geometric decoupling between the resonators.
JBiSE

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
171
JBiSE
cage was made longer than the inner one, minimizing
coupling between the pairs of end rings. Realistically,
one cannot achieve perfect isolation between the two
coils since the legs of the coils are in close proximity.
The coupling is only reduced. For our application, since
the resonance frequencies of the coils are far away from
each others, this geometric decoupling could be accepted.
Quantitative measurements of the isolation between the
two coils are given in Section 3.
The calculation of the necessary tuning capacitances
for both coils was performed using the free Birdcage-
Builder software from the Penn State University [18].
The practical used values for these tuning capacitors
were close to the calculated ones.
Variables capacitors were used to perform a final fine
tuning when the coils are placed inside the magnet and
loaded by the sample (Figure 3). For the low-pass bird-
cage, the fine tuning is achieved by adjusting tow dia-
metrically opposed capacitors (CLT in Figure 3). This
method allows only a limited tuning range. Otherwise, it
destroys the original ideal sinusoidal current distribution
between the legs [17]. After testing the resonator inside
the magnet and with different loads (phantoms, rats), we
limited the tunable range of the resonator to sufficient
range of about 500 kHz, so as the current deviation from
its ideal is less than 10% [17].
For the high-pass resonator, the fine tuning is accom-
plished by connecting two diametrically opposed points
of one end-ring by a tuning variable capacitor CHT. In
practice, the connection of these points is made by adding
a second circular loop, parallel to the end-ring as shown in
Figure 3. To preserve the symmetry of the birdcage, the
adjustable tuning capacitors CHT was balanced by a fixed
capacitor of 8.2 pF connected across the second end ring
using a circular loop (not shown in Figure 3).
Both coils were derived in linear polarization opera-
tion in such a way that theirs B1 fields are orthogonal as
it was shown on Figure 2. We used balanced capacitive
coupling to match their impedances to the real value of
50 at the working frequency [17]. The coupling scheme
of the low-pass resonator is shown together with the
active decoupling circuitry in Figure 4. In this scheme,
the capacitive coupling, accomplished through the ca-
pacitors CT and CM1, CM2 and CM3, is preceded by a
balun (balanced-unbalanced). This balun is required for
the probe circuit, not only due to the necessity of equili-
brating the probe coil and consequently minimizing di-
electric coil losses in the sample, but also for blocking
unwanted common mode currents. The balun we built is
widely used in radio technologies [17,19]. It consists of
a simple “trap” circuit where the loop formed by the
coaxial cable is tuned by a capacitor, CB, to form a high
impedance circuit at the xenon frequency.
The capacitive coupling scheme for the low-pass bir-
dcage is straightforward to be generalized to proton res-
onator. It will not be given here, but the interesting rea-
der may simply write to the author for more details.
As it was outlined in a previous paragraph, the trans-
mit-only birdcage coil for 129Xe (inner coil) must be de-
coupled from the receive-only surface coil (described in
the next section). During transmission, this is mandatory
since a coupling between the coils would result in dele-
terious consequences for the homogeneity of the B1 field
and in a possible destruction of the receive coil. During
receiving, part of the electromagnetic energy received by
the receive coil will be dissipated in the transmit volume
Low-pass resonator High-pass resonator
Figure 3. The fine tuning of the resonators. Low-pass resonator: CL=340 pF (220 pF+120 pF chip ca-
pacitors from American Technical Ceramics Corp. ATC), CLT=1.5-55 pF (adjustable capacitor
NMAT55HVE from Voltronics). High-pass resonator (only one end-ring is shown): CH=40 pF (18
pF+12 pF chip capacitors from ATC Corp.), CHT=1-15 pF (adjustable capacitor NMAT15HVE from
Voltronics). To preserve the symmetry of the high-pass birdcage, a tuning capacitor (which balances CHT)
should be connected across the second end-ring (not shown in the figure). We used a fixed chip capaci-
tor of 8.2 pF to balance CHT.

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
172
JBiSE
Figure 4. The 129Xe birdcage electrical circuit associated with the active de-
coupling circuitry. CL (220+120 pF from ATC Corp.) are tuning capacitors, LC
(8 µH+ 4x1 µH from CoilCraft) are the RF choke inductors, L (90 nH) are
parallel inductors (from CoilCraft) and D are PIN diodes (UM4006 from Mi-
crosemi Corp). Matching is achieved by the use of balanced capacitive cou-
pling through ATC chip capacitors: CT=220 pF, CM1=120 pF, CM2=53 pF (10
pF+10pF+33pF). For fine matching, CM3 is a variable capacitor with capaci-
tance between 1.5 pF and 55 pF (NMAT55HVE from Voltronics). Notice that
, which is with good agreement with balanced capaci-
tive coupling condition [17]. The adjustable 1.5 pF-55 pF tuning capacitors
CLT for fine tuning are not shown (see Figure 3). The tuning capacitor CB for
the balun CB=180 pF+120 pF.
23TM M
CC CC
L
coil resulting in SNR and sensitivity losses. Therefore,
the interaction of the volume resonator and the receive
coil must be cancelled out.
Decoupling can be achieved in principle geometrically
(as in the case of the two concentric birdcages). How-
ever, when the coils are tuned to the same frequency, it is
difficult in practice to perfectly decouple them by simply
adjusting their mutual orientation. Hence, geometric de-
coupling should be accompanied or replaced by another
effective mean that cancels out or, at least, minimize the
induced currents in each coil.
This can actually be done by shifting the resonance
frequency of one coil with respect to the other during
appropriate periods. This detuning method is well de-
scribed and used to detune a birdcage coil in [20] for
example. In this relatively simple method, usually used
in commercial coils, the current in the detuned coil is not
cancelled out but only reduced.
In more efficient design, the current in the resonator is
cancelled out at the working frequency, improving the
isolation to a high level. In such a design, which is based
on the “pole insertion” technique, an inductance is con-
nected, using switching diodes, in parallel to one or
more (depending on the coil structure and type) tuning
capacitors, constituting a parallel resonant circuit (pole
insertion). If the pole is well-tuned to the working fre-
quency, the resulting “trap” circuit becomes an efficient
blocking RF current device. This method was described
for the case of a single-tuned high-pass birdcage [21]. In
our prototype, we demonstrate that this method can be
generalized to the case of a low-pass birdcage which has
not been published.
We realized an active decoupling through the use of
high voltage PIN diodes (UM4006, Microsemi Corp)
distributed across one set of the legs capacitors. The sch-
ematic of the decoupling circuitry is shown in Figure 4.
For each leg, two bias rails apply the appropriate DC
bias, through the RF choke inductors LC to the PIN diode
which, when it is forward-biased, switches an inductor L
in parallel with the tuning capacitor CL. If the value of
the inductor L is correctly chosen (2
0
.. 1
L
LC 

), the
resulting parallel circuit presents a high impedance at the
working frequency so as the RF current is cancelled out
in each leg and the transmit volume coil is “off” (re-
ceiving period). When the diode is reverse-biased, it
forms an open circuit and the inductor L is disconnected.
In this case, the coil is “on” (transmission period).
The inductors LC in the direct DC path serve as
chokes and ensure “isolation” between the RF part and

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
173
JBiSE
Figure 5. The double-tuned volume coil. We can see (right
picture) the four variable tuning and matching capacitors for
the two channels.
the DC bias circuit by minimizing RF currents on the
DC path.
Figure 5 shows the double-tuned volume coil.
2.2. The Receive-Only Single-Tuned Coil
A surface coil was chosen for the receive-only function.
It consists of a single circular loop of 4 cm of diameter
(Figure 6). The tuning capacitor Ct was distributed into
two series capacitors (2Ct) positioned in two different
physical sites of the loop to reduce the losses associated
with the electric field generated by the coil [22,23].
The decoupling of this coil from the transmit volume
coil was achieved using the “pole insertion” method as
illustrated in Figure 6. The PIN diode switches or not in
parallel the inductor Ls across one of the distributed tune
capacitors (2Ct). During the transmission period the coil
is “off” (diode is forward-biased), while it is “on” during
the receiving period (diode is reverse biased).
Figure 6. The receive-only coil and its active decoupling cir-
cuitry. The fixed tuning is 2Ct=2x220 pF+2x150 pF+100 pF
(chip capacitors from ATC). The adjustable capacitor of 1.5-55
pF (NMAT55HVE from Voltronics) is not shown. The match-
ing capacitors are: Cm1=68 pF+10 pF, Cm2=56 pF Cm3=1.5-55
pF (adjustable capacitors NMAT55HVE from Voltronics).
Ls=40 nH (3 parallel 120 nH inductor from CoilCraft) is the
parallel inductor. D is a PIN diode (UM4006 from Microsemi
Corp). Lc=8 µH+5x 1 µH are RF choke inductors (from Coil-
Craft).
Figure 7. The receive-only surface coil and its circuitry. The
loop was fixed on one side of a printed board and all compo-
nents were soldered on the other side.
Notice finally that the receive-only coil was matched
using balanced capacitive coupling (capacitors Cm1, Cm2,
and Cm3 in Figure 6). RF choke inductors LC are used to
“separate” the RF circuit from DC bias one. Figure 7
shows the developed single-tuned receive only coil.
2.3. The T/R Driver
PIN diodes for both transmit and receive-only coils were
biased through the use of a home-made T/R driver. This
driver provides the gated DC bias necessary for elec-
tronically turning coils “on” and “off”. We designed the
driver with a gated input (a TTL compatible signal sup-
plied by the NMR spectrometer) and two separate TX
and RX bias sections. The TX stage applies a –12 Vdc
bias (PIN diodes are reverse biased and transmit volume
coil is “on”) and a 5 Vdc bias (PIN diodes are forward
biased and transmit coil is “off”). The RX stage applies
–12 Vdc bias (PIN diode opened and receive-only coil is
“on”) and a 5 Vdc bias (PIN diode closed and the coil is
“off”). The driver contains current limitation resistors to
limit the maximum output current to 1 A for both TX
and RX sections.
The flexibility of the driver allows the possibility to
turn “off” and “on” both coils by using a manual switch
(without the need to a TTL signal) for fine tuning and
matching purposes inside the magnet. Another manual
switch allows also the use of volume 129Xe coil for both
transmission and receiving. In this situation, the receive-
only coil is switched “off” during all the experiment
(PIN diode forward biased).The design of the driver was
based on the use of bipolar transistors.
3. RESULTS AND DISCUSSIONS
Although this specific probe-head was initially devel-
oped for the MRI of HP 129Xe, we performed all the tests

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
174
JBiSE
and all the validation experiments at proton and so-
dium-23 (23Na) frequencies. Actually, it is important to
notice that the magnetization of HP 129Xe is artificial and
it is therefore not renewable when destroyed by the RF
excitation pulses. Clearly, this is not a convenient situa-
tion for coil characterization and adjustments. Sodium
frequency at 2.35 Tesla (26.4 MHz) is very close to the
129Xe one (27.7 MHz). We could therefore consider that
all the tests and measurements conducted at the sodium
frequency would be valid at the 129Xe frequency. Valida-
tion at the HP 129Xe frequency for in vivo applications
will make the subject of a future work.
During all the tests, the receive-only surface coil was
placed inside the volume probe in such a way that geo-
metric decoupling between it and the proton volume coil,
is always achieved. It is clear here that no geometric
decoupling could be realized between the receive-only
coil and the sodium volume coil. Only active decoupling
between these coils is achieved.
As it was mentioned in a previous section, a single-
tuned birdcage will resonate in several modes, with the
most useful to MRI being Mode 1. In this mode, a very
homogenous transverse magnetic field is created by the
sinusoidal distribution of the currents in the legs. It is
important for our application to know whether or not this
distribution of currents was disturbed by the double-
tuned configuration, and consequently to know whether
or not the B1 homogeneity, especially for the 23Na coil, is
preserved. Moreover, we are interested in knowing if the
receive-only coil would or not disturb the sinusoidal
distribution of the currents of the proton coil since only
geometric decoupling was realized between these coils.
Also, we wanted to know if the only-active decoupling
between this receive-only coil and sodium volume coil
would be sufficient to preserve the homogeneity of the
B1 without the need of geometric decoupling. All these
information could be obtained by measuring the currents
in the legs of both birdcages at the proton and sodium
frequencies. We measured the currents distribution in each
coil by placing a small sense loop adjacent to the legs.
The receive-only coil was switched off during all meas-
urements. At the sodium frequency, we measured the
currents in the legs of sodium birdcage (inner resonator),
as well as the currents induced in the proton legs (outer
resonator). Similarly, these currents were measured at the
proton frequency. Measurements are shown in Figure 8.
At the sodium frequency, the currents in the sodium
birdcage (inner coil) follow a cosine distribution as one
moves radially around the coil. Small currents are in-
duced in the proton resonator (outer coil). These currents
are in phase with and add to the currents in the sodium
coil. From this information, one can expect that the
transverse magnetic field B1 created by the sodium coil
would still homogenous and would not be disturbed by
the proton coil. At the proton frequency, the measure-
ments show strong out of phase currents in the sodium
coil. Despite the presence of these currents, the currents
in the proton legs still follow the desired sinusoidal dis-
tribution as one moves radially around the coil. The coil
still creates a homogenous B1 field. The presence of the
out of phase currents in the sodium coil will result, off
course, in losses in the efficiency of the proton coil when
compared to a single-tuned one.
With the simple measurement method of currents which
was based on the use a small sense loop, no changes on
currents distribution were observed with or without the
presence of the surface coil. One could then argue that
the presence of the receive-only coil, in its “off” state,
does not disturb the homogeneity maps of B1 fields of
both sodium and proton birdcages.
(a) Currents distribution at sodium frequency
(b) Currents distribution at proton frequency
Figure 8. Currents distribution in the resonators at (a) sodium
and (b) proton frequencies.

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
175
JBiSE
It’s important to evaluate the degree of isolation be-
tween the coils. We have conducted isolation measure-
ments between the different coils using a network ana-
lyzer. At sodium frequency, the isolation between the
proton and the sodium birdcages was greater than 25 dB.
At proton frequency, the isolation was lowered to about
16 dB. These measurements are in agreement with cur-
rent measurements. Actually, at the sodium frequency,
very small currents are induced in the proton coil and
one can expect good isolation. Conversely, at the proton
frequency, where large currents are induced in the so-
dium coil, the isolation is naturally lowered.
Isolation measurements between the receive-only coil
and the two volume coils were also conducted. At the
proton frequency, the isolation between the proton bird-
cage and the surface coil was about 26 dB. This isolation
is very good and confirms that the cosine current distri-
bution is not, or little, disturbed by the surface coil.
Geometric decoupling between these two coils should
sufficient for our application. Isolation between the so-
dium birdcage and the receive-only coil was also meas-
ured, with the only-active decoupling, to be about 31 dB.
This excellent isolation between the two coils is essential
in maintaining a transmit field homogeneity and in re-
taining the high sensitivity of the surface coil.
We have realized MRI images using the developed
probe. In the first time, images were acquired using phan-
toms. A variety of phantoms was used to test the devel-
oped probe in various conditions and configurations.
The first series of images was realized using the proton
birdcage and the sodium birdcage biased as trans-
mit/receive coils. The surface coil was in its “off” state
during both transmitting and receiving periods. The phan-
tom (see Figure 9) was consisting of three cylindrical
tubes with a same height (50 mm). The biggest tube (27
mm of inner diameter) was filled with saline solution
(pure water saturated with NaCl). The two small tubes
(12.5 mm of inner diameter) were filled with pure water.
H
2
O + NaCl
H
2
O H
2
O
Figure 9. The cross-section
(drawn to scale) of the
phantom used to validate the
double-tune birdcages.
(a)
(b)
Figure 10. The cross-sectional (axial) MRI images (displayed
to scale) of a phantom acquired at (a) proton frequency (100
MHz) and (b) sodium frequency (26.4 MHz). A spin echo
pulse sequence was used to acquire images. FOV (Field of
View)=70x70 mm2, repetition time TR=1000 ms, echo time
TE=31 ms, flip angle=90°. Slice thickness for proton images=2
mm. Sodium images were acquired with 2 signal averages and
without slice selection.
Figure 10 shows the cross-sectional MRI spin-echo
images of the phantom obtained at 100 MHz and 26.4
MHz with a field of view (FOV) of 70x70 mm2 and an
acquisition matrix of 128x64. Before transformation of
the acquired data to the image space, zero filling to
128x128 was applied.
In these images, we can see the high homogeneity of
the B1 produced at 26.4 MHz. This confirms the currents
measurements addressed above. The sodium birdcage was
not disturbed by the proton coil. The homogeneity of the
B1 created at 100 MHz is shown to be also sufficient.
The second series of images was realized with the
proton birdcage used for both transmitting and receiving.
The sodium birdcage was biased as a transmit-only coil,
while the surface sodium coil was biased as a receive-
only coil. This is actually the final configuration of use
of the developed probe for our application. A three-

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
176
compartment cylindrical phantom was used for this ex-
periment (Figure 11). It was formed of a small tube (inner
diameter=10.5 mm) filled with pure water and placed
approximately diagonally inside a larger tube (inner di-
ameter=28 mm) filled with a saline solution. Between
the two tubes, we have intentionally conserved an empty
space with a variable width as illustrated by Figure 11.
JBiSE
Cross-sectional MRI spin-echo images of this three-
compartment phantom were acquired at 100 MHz and at
26.4 MHz (see Figure12) with a FOV of 50x50 mm2
and an acquisition matrix of 128x64. Before transforma-
tion of the acquired data to the image space, zero filling
to 128x128 was applied.
We can recognize on the sodium images the typical
sensitivity of a surface coil. This sensitivity decreases
when one moves away from the center of the coil. We
can, off course, expect a better signal-to-noise ratio with
a surface coil when compared with a volume coil. This is
actually well-known and no more signal-to-noise ratio
characterization is necessary.
The validation of the developed probe was further
demonstrated in vivo in the rat brain. Firstly, spin-echo
images were obtained at the proton frequency. Figure 13
(respectively Figure 14) shows a set of cross-sectional
(respectively coronal) images.
Notice that, in the configuration of the developed probe,
the filling factor of the outer birdcage (proton coil) is not
optimum for the outer coil. Consequently, its SNR per-
formances are relatively reduced when compared with
those that could be obtained by a single simple-tuned
birdcage. Actually, it is well-known that the SNR is di-
rectly proportional to square root of the coil filling factor.
This point does not handicap the probe since high SNR
performances of the 1H coil are not an issue in our ap-
plication. Actually, the 1H will be only used for acquir-
ing localization (anatomic) images where signal averag-
ing could be applied to enhance the SNR. For shimming
purposes, there is no need to have a high SNR.
The developed probe was also tested in vivo at the so-
dium frequency on the same rat. The sodium birdcage
was biased as a transmit-only coil, while the sodium
surface coil was biased as a receive-only coil. Our goal
here is not to have sodium images of the rat brain, but
only to demonstrate a basic functioning of the sodium
coil in vivo. Only sodium NMR signal and spectra in the
rat brain were acquired as it is shown in Figure 15. We
obtained the typical sodium spectra of sodium in the rat
brain with a typical peak width of few ppm [24].
In vivo sodium MRI images were difficult to obtain
with classical MRI pulse sequences (spin-echo, gradi-
ent-echo) except if an incredible number of signal av-
erages (more than 20 hours of acquisition time) is used.
Actually, because of the short T2 transverse relaxation
time components of sodium in the rat brain (0.7-3 ms or
16-30 ms depending on the value of the magnetic field),
the acquisition of sodium images requires the imple-
H
2
O + NaCl
H
2
O
Empty space
Figure 11. A schematic diagram (drawn
to scale) of the cross-section of the
three-compartment phantom.
(a) (b)
Figure 12. Cross-sectional (axial) MRI images (displayed to scale) of the three-compartment phantom acquired at (a) proton
frequency (100 MHz) and (b) sodium frequency (26.4 MHz). A spin echo pulse sequence was used to acquire images.
FOV=50x50 mm2, flip angle=90°, echo time TE=30 ms, repetition time TR=2000 ms for proton and 300 ms for sodium. Slice
thickness for proton images=2 mm. Sodium images were acquired with no slice selection and with no signal averaging.

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
177
Figure 13. Cross-sectional MRI images of the rat brain acquired at 100 MHz. A spin echo pulse sequence was used to acquire
images. FOV (Field of View)=60x60 mm2, flip angle=90°, echo time TE=30 ms, repetition time TR=4000 ms, slice thick-
ness=1 mm. Acquisition matrix was 128x64. Before transformation of the acquired data to the image space, zero filling to
128x128 was applied.
Figure 14. Coronal MRI images of the rat brain acquired at 100 MHz. A spin echo pulse sequence was used to acquire images. FOV
(Field of View)=60x60 mm2, flip angle=90°, echo time TE=17 ms, repetition time TR=4000 ms, slice thickness=1 mm. Acquisition
matrix was 128x64. Before transformation of the acquired data to the image space, zero filling to 128x128 was applied.
JBiSE

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
178
(a) (b)
Figure 15. In-vivo sodium (a) NMR signal and (b) its spectrum in the rat brain. A hard non-selective excitation pulse and 512 signal
averages were used.
mentation of dedicated three-dimensional (3D) and short
echo-time pulse sequences (3D back-projection se-
quence and 3D GRE for example) [14,25]. The imple-
mentation of 3D sequences is out of the scope of our
research topics.
JBiSE
Nevertheless, we think that our probe presents no in-
trinsic limitation for acquiring in-vivo sodium images.
The only limitation is the available MRI sequence. To
demonstrate this without the need to implement new
sequences, we tried to acquire sodium image of a phan-
tom containing a sodium concentration as low as those
which are generally present in vivo. The sodium trans-
verse relaxation time of the phantom must be longer than
its in-vivo value to allow the use of a classical spin-echo
sequence. So, we used a syringe filled with 20 ml of a
physiological serum. The sodium concentration (about
0.1%) in this serum is identical, or at least close, to the
in-vivo sodium concentrations.
Figure 16 shows transverse, coronal and sagittal im-
ages of the phantom at both proton and sodium frequen-
cies. Sodium images were acquired with 64 signal aver-
ages.
All these results indicate that the performances of the
design presented here are good. We think that these per-
formances would be at least as good as, if not better, at
the HP 129Xe frequency. Although, the resonance fre-
quency of sodium and xenon are closed, no significant
noise from sodium nuclei will disturb the HP 129Xe sig-
nal. Firstly, the excitation plus at 27.7 MHz will not sig-
nificantly, if never, excite the sodium nuclei. Secondly,
the in vivo HP 129Xe signal would be more than nine or
ten orders of magnitude greater than an eventual in vivo
sodium signal. Moreover, since the quality factor of the
receiving coil is relatively high (about 100), its obtained
bandwidth is narrow and it will significantly limit the
out-of-band noise. In any case, no influence of the so-
dium nuclei was observed during our HP 129Xe NMR
experiment.
We are currently building another prototype that in-
cludes shielding of the birdcages to minimize external
interferences. Some modifications in the mechanical
structure of the probe should be achieved to allow for
easy delivering of the hyperpolarized xenon to the rat.
The next step of our work is the use of this new probe
for rCBF in the rat brain. This would be the subject of a
future publication. At the same time, a third prototype of
the probe is currently being studied for use at a higher
magnetic field (4.7 Teslas) for the same application.
Notice finally that the choice of sodium for a first
validation of the developed probe is very important, not
only because its frequency is close enough to the xenon
frequency, but also because of large palette of potential
applications of the sodium MRI for both humans and
small animals. Actually, MRI is proving to be a very
useful tool for sodium quantification in animal models of
stroke, ischemia, and cancer in small animals and in
humans. For example, 23Na MRI short echo times was
used to quantify absolute tissue sodium concentration in
patients with brain tumors showing an increase of so-
dium concentration relative to that in normal brain
structures [14,26]. In rabbit models of focal cerebral
ischemia, it was shown that there are changes in tissue
23Na signal levels following acute ischemia, which may
help to identify necrotic tissue and estimate the duration
of the ischemia [27,28]. Recently, the sodium and proton
diffusion MRI are shown to be good biomarkers for
early therapeutic response in subcutaneous tumors [26].
Recent advances in ultra-high field MRI hardware have

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
180
JBiSE
(a) (b) (c)
(d) (e) (f)
Figure 16. Cross-sectional, coronal and sagittal MRI images of the syringe (inner diameter=27 mm) filled with a physiological serum
acquired at (a), (b), (c) proton frequency and at (d), (e), (f) sodium frequency. A spin echo pulse sequence was used to acquire images.
FOV (Field of View)=70x70 mm2, flip angle=90°, echo time TE=17 ms, repetition time TR=2000 ms for proton and 500 ms for so-
dium. Slice thickness for proton images=1 mm. Acquisition matrix was 128x64 for proton and 64x32 for sodium. Before transforma-
tion of the acquired data to the image space, zero fillings to 128x128 for proton and to 64x64 for sodium were applied. Sodium im-
ages were acquired with 64 signal averages and without slice selection.
allowed the acquisition of 23Na images of the mouse
heart [14]. These examples show that in human and
animal models, there is a need to obtain interleaved co-
registered anatomical (1H) and physiological (23Na) in-
formation which may be useful for assessing disease and
effective therapeutic intervention. For this purpose, dou-
ble tuned volume and/or surface coils are required. One
of the main advantages of the probe that we presented is
the possibility to use it in different configurations ac-
cording to the application.
Complete schematics of the developed circuits, values
and references of all parts, or any other details of design
and realization could be obtained by simply writing to
the author.
4. CONCLUSIONS
In summary, we have reported the design and testing of a
new MRI probe consisting of a double-tuned volume
coil actively decoupled from a receive-only coil. This
developed probe provides proton and sodium images at
2.35 T. In-vivo and phantom 1H images were acquired
with the volume transmit/receive coil. 23Na images of
phantoms in both transmit/receive mode and in transmit-
only/receive-only mode were acquired. In-vivo NMR
signals in the rat brain are also obtained. The results
showed acceptable performances of the design. This
probe is relatively simple to build and it is low-cost. It
should be a suitable tool for the HP 129Xe NMR experi-
ments in the rat brain. In its final version, under con-
struction, it should enhance the quantitative measure-
ments of the rCBF and their reliability.
5. ACKNOWLEDGEMENTS
The author wishes to thank Christoph Segebarth, Research Director
and Director of the Team 5 of GIN (Grenoble-Institute of Neurosci-
ences) and Jean-Louis Leviel for supporting with enthusiasm the re-
search topic on the HP gases in the laboratory. I thank also Olivier
Montigon for his help in the realization of the T/R driver circuitry,
Vincent Auboiroux for his participation in the mechanical construction

A. Asfour / J. Biomedical Science and Engineering 3 (2010) 167-180
180
JBiSE
of the probe. Finally, I am very thankful to Emmanuel Barbier for his
support in using his developed MRI sequences and visualization soft-
ware and for helpful discussions. All these persons are working at the
Team 5 of GIN.
REFERENCES
[1] Albert, M.S. et al. (1994) Biological magnetic resonance
imaging using laser polarized 129Xe. Nature, 370,
199-201.
[2] Kilian, W. et al. (2004) Dynamic NMR spectroscopy of
hyperpolarized 12 9Xe in human brain analyzed by an up-
take model. Magnetic Resonance in Medicine (MRM), 51,
843-847.
[3] Venkatesh, A.K. et al. (2001) Using dynamic hyperpo-
larized xenon MR to measure brain perfusion. Proceed-
ing of the International Society of Magnetic Resonance
in Medicine, 9, 951.
[4] Swanson, S.D. et al. (1997) Brain MRI with la-
ser-polarized 129Xe. Magnetic Resonance in Medicine
(MRM), 38, 695-698.
[5] Duhamel, G. et al. (2002) Global and regional cerebral
blood flow measurements using NMR of injected hyper-
polarized xenon-129. Acad Radiol., 9(2), 498-500.
[6] Asfour, A. et al. (2006) Development of a fully digital
and low frequency NMR system for polarization meas-
urement of hyperpolarized gases. Proceedings of the
IEEE Instrumentation and Measurement Technology
Conference IMTC, Sorrento, Italy.
[7] Asfour, A. (2008) A new DAQ-based and versatile
low-cost NMR spectrometer working at very-low mag-
netic field (4.5 mT): A palette of potential applications.
Proceedings of the IEEE International Instrumentation
and Measurement Technology Conference I2MTC, Van-
couver, Canada.
[8] Ackerman, et al. (1980) Mapping of metabolites in
whole animals by 31P NMR using surface coils. Nature,
283, 167-170.
[9] Hayes, C.E. et al. (1985) An efficient, highly homoge-
nous radiofrequency coil for whole-body NMR imaging
at 1.5 T. Journal of Magnetic Resonance, 63, 622-628.
[10] Fitzsimmons, J.R. et al. (1993) Double Resonant quad-
rature birdcage. Magnetic Resonance in Medicine (MRM),
30, 107-114.
[11] Rath, A.R. (1990) Design and performance of a dou-
ble-tuned birdcage coil. Journal of Magnetic Resonance,
86, 488-495.
[12] Shen, G.X. et al. (1997) Dual-frequency, dual-quadrature,
birdcage RF coil design with identical B1 pattern for so-
dium and proton imaging of the human brain at 1.5 T.
Magnetic Resonance in Medicine, 38, 717-725.
[13] Lanz, T. et al. (2001) Double tuned 23Na 1H nuclear
magnetic resonance birdcage for application on mice in
vivo. Review of Scientific Instruments, 72(5), 2508-2510.
[14] Alecci, M. et al. (2006) Practical design of a 4 Tesla
double-tuned RF surface coil for interleaved 1H and 23Na
MRI of the rat brain. Journal of Magnetic Resonance,
181, 203-211.
[15] Leach, M.O. et al. (1986) The design and use of a
dual-frequency surface coil providing proton images for
improved localization in 31P spectroscopy of small le-
sions. Med Phys, 13, 510-513.
[16] Volotovskyy, V. et al. (2002) Doubly tunable double ring
surface coil. Concepts in Magnetic Resonance, B(17B),
11-16.
[17] Mispelter, J., Lupu, M. and Briguet, A. (2006) NMR
Probeheads for Biophysical and Biomedical Experiments:
Theoretical Principles and Practical Guidelines, Impe-
rial College Press.
[18] Chin, C.L. et al. (2002) BirdcageBuilder: Design of
specified-geometry birdcage coils with desired current
pattern and resonance frequency. Concepts in Magnetic
Resonance, 15(2), 156-163.
[19] The ARRL Handbook for Radio Amateurs, Published by
the American Radio Relay League, 1996.
[20] Streif, J.U.G. et al. (2003) A coil combination for mag-
netic resonance perfusion imaging of mice in vivo at 7 T.
Review of Scientific Instruments, 74(5), 2843-2848.
[21] Barberi, E.A. et al. (2000) A transmit-only/receive-only
(TORO) RF system for high-field MRI/MRS applications.
Magnetic Resonance in Medicine (MRM), 43, 284-289.
[22] Décorps, M. et al. (1985) An inductively coupled, series
tuned NMR probe. Journal of Magnetic Resonance, 65,
100-109.
[23] Hoult, D.I. and Lauterburg, P.C. (1979) The sensitivity of
the zeugmatographic experiment involving human sam-
ples. Journal of Magnetic Resonance, 34, 425-433.
[24] Naritomi, H. et al. (1987) In vivo measurements of the
intra- and extra-cellular +Na and water in the brain and
muscle by nuclear magnetic resonance spectroscopy with
shift reagent. BIOPHYS Journal, Biophysical Society, 52,
611-616.
[25] Schepkin, V.D. et al. (2006) Sodium and proton diffusion
MRI as biomarkers for early therapeutic response in
subcutaneous tumors. Magnetic Resonance Imaging, 24,
273-278.
[26] Ouwerkerk, R. et al. (2003) Tissue sodium concentration
in human brain tumors as measured with 23Na MR imag-
ing. Radiology, 227, 529-537.
[27] Bartha, R. et al. (2004) Sodium T2*-weighted MR im-
aging of acute focal cerebral ischemia in rabbits. Mag-
netic Resonance Imaging, 22, 983-991.
[28] Wang, Y. et al. (2000) Brain tissue sodium is a ticking
clock telling time after arterial occlusion in rat focal
cerebral ischemia. Stroke, 31, 1386-1391.