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Flourine-19 NMR probe design for noninvasive tumoral pharmacokinetics

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Flourine-19 NMR probe design for noninvasive tumoral pharmacokinetics

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Content 19F NMR PROBE DESIGN FOR NONINVASIVE
TUM ORAL PHARM ACOKINETICS
BY
HYUN KWON KIM
A Thesis Presented to the
FACULTY OF THE SCHOOL OF ENGINEERING
UNIVERSITY OF SOUTHERN CALIFORNIA
In Partial Fulfillment of the
Requirements for the Degree
MASTER OF SCIENCE IN BIOMEDICAL ENGINEERING
April 1995
Copyright 1995 Hyun Kwon Kim
This thesis, 'written by
. . . . . . . . . . . . . . . .H y u n ..J K w A n „ M m . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
under the guidance of Faculty Committee
and approved by all its members, has been
presented to and accepted by the School of
Engineering in partial fulfillment of the re
quirements for the degree of
Master of' Science
D ate A p r n - 20- 199 5_
Faculty Committee
.. ......
^ Chairman . ,,
L
ACKNOW LEDGMENTS
I am sincerely grateful to the many people who gave me
advice, assistance and encouragement during the course of this
work. In particular I am indebted to my advisor, Professor M.
Singh, and Professor W. Wolf for their positive direction and many
useful and encouraging discussions. I would also like to thank Ms.
C. Johnson and my wife, Julie, for typing and proofing this
manuscript. I am also grateful to the Electrical Engineering
department for the use of their laboratory and test equipment. I
wish to add my thanks to my former mentor, Professor H. Keyzer
for his continual encouragement.
ABSTRACT
A 19F coil for the Siemens Magnetic Resonance Spectrometer
(MRS) was not available for the pharmacokinetic studies involving
antitumor compounds in our clinic.
An MRS coil was required capable of being tuned to a 1H
frequency for MR imaging, and to a 19F frequency for MR
spectroscopy. Such a coil was designed and tested in an MR
spectroscopic in vivo study involving 5-fluorouracil on human
cancer test patients. The coil design had a high signal to noise
ratio and proved to be suitable for the pharmacokinetic studies
required.
TABLE OF CONTENTS
1 INTRODUCTION 1
2 SURFACE COIL THEORY
2.1 INTRODUCTION 3
2.2 RECEIVER COIL 6
2.3 TRANSMITTER COIL 9
2.4 TRANSMITTER/RECEIVER COIL 11
2.5 HOMOGENEITY OF COIL 1 3
2.6 SENSITIVITY OF COIL 18
2.7 COIL LOCALIZATION 2 0
2.8 FIELD INSENSITIVE PULSES 2 4
3 SURFACE COIL DESIGN PRACTICALITIES
3.1 INTRODUCTION 2 5
3.2 SHIELDING 2 7
3.3 COUPLING 2 8
3.4 L-C RATIO 3 1
3.5 SENSITIVITY VS. HOMOGENEITY 3 2
4 SURFACE COIL EXPERIMENT
4.1 COIL REQUIREMENTS 3 4
4.2 CIRCUIT DESCRIPTION 3 5
4.3 COIL TESTING (BENCH WORK) 3 8
4.4 COIL TESTING (INVIVO) 3 9
4.5 COIL MAPPING 4 0
4.6 VOI DETERMINATION WITH 4 0
RESPECT TO THE COIL
5 IN VIVO STUDIES 4 1
6 FIGURES 4 2
7 REFERENCES
1 INTR O D U C TIO N
One major difficulty encountered in the quantification of in
vivo MR Spectra is the inhomogeneous Bj field of many of the RF
coils. In such MRS studies the measured signal depends (among
other factors) on the observation frequency, the nucleus, the RF
pulse used, the Bj field produced by the coil at a given position,
the receiver sensitivity of the coil and the position and size of the
VOI (volume of interest). The nutation angle produced by a
particular pulse is dependent on the B i field strength giving rise to
a position-dependent nutation angle. By the principle of
reciprocity, the receiver sensitivity is also dependent on the
relative B 1 field produced by the RF coil. Both the effects of flip
angle and receiver sensitivity cause the amplitude of the received
NMR signals to depend on spatial position. This B i dependence can
be reduced by using pulses that are unaffected by inhomogeneous
B i fields, although the receive sensitivity will still depend on the
position of the spins.
There are two methods at present that take into account B i
inhomogeneity and may be used for quantification. The first is to
perform theoretical calculations using computer calculations of the
spatial dependence of the RF fields [1,2]. This method has been
used to evaluate the ISIS localization technique by observing the
effects of various parameters on contamination and signal loss
using a surface coil and a headcoil. The computer integrates the
signal from a VOI by dividing the chosen tissue volume into voxels
1
and determines the signal in each voxel. The spatial sensitivity of
various specific surface coil has also been calculated
theoretically (3,4).
A second method of accounting for Bj inhomogeneity is to
perform calibration experiments (5,6) of a VOI in a large phantom,
selected to correspond in position and loading conditions to that of
a clinical examination, in order to calculate absolute concentrations
of metabolites. This method has been assessed by Buchli and
Boesiger(7) where they compared the accuracy of different
quantification methods to try to find a method that would be easy
to handle and could be used routinely in a clinical environment.
They[8] compared ISIS localization with localization using a surface
coil for three different concentration standards: heteronuclear
calibration using the internal water signal, homonuclear calibration
using an external phantom in a separate experiment and
homonuclear calibration using an external phantom symmetrically
placed in the experimental arrangement, i.e., the phantom was
placed in an equally sensitive region of whichever coil was being
used as the tissue under investigation. They found the phantom-
based calibration strategies in combination with the ISIS
localization to be the most accurate method, giving measurement
errors of 5-7%. These errors arise due to losses inherent in the
ISIS technique resulting from incomplete inversion from the
pulses used and T1 losses between inversions giving rise to a low
selection efficiency(9,10).
2
The theoretical computational method has one advantage
over this phantom-based method in that it avoids the need for
separate calibration phantom experiments for each parameter
change in the ISIS experiment and yields the sensitivity of the
experiment over any desired VOI or subregion of the VOL This
technique may therefore be used to advantage when the VOI
encompasses two tissue types[ll]. The phantom-based
experimental method, however, has the advantage of being the
most similar in experimental parameters to the clinical
environment hence enabling more accurate determination of the
absolute concentrations of in vivo metabolites of 5 fluorouracil
(5FU).
2 SURFACE COIL THEORY
2.1 INTRODUCTION
In the radio domain transmit antennas are designed to
"radiate" their energy as far as possible, and reception antennas
are designed to pick up signal from distant sources. In both cases
the field of interest is far from the antenna body and can therefore
be described by simplified Maxwell equations disregarding current
and charge effects (so-called far-field approximation). The field
can be characterized in terms of travelling waves, where the
3
spatial and temporal, as well as the electric and magnetic behavior,
are inseparable.
NMR probes, in contrast, are designed to confine their field
and energy density to a well-defined volume; any distant field
distribution contributes to power and sensitivity losses, but not to
the desired spin interaction. Sensitivity to the region outside the
probe results in a low "filling factor" (defined by the fraction of the
sensitive volume occupied by the sample), in susceptibility to
external noise, and in reduction of overall sensitivity. As a
consequence, NMR probes are characterized much more by the
electric and magnetic fields caused by the current and charge
distribution on the conductor structure than by free-field effects
(near-field approximation). The full Maxwell equations including
charge and current distribution have to be applied to describe the
fields. NMR probes produce standing waves, with regions of high
electric field more-or-less separated from regions of high magnetic
field.
Thus, one will generally build radio antennas as open
structures with large dimensions of quarter to half wavelengths,
but NMR probes as closed structures with smallest dimensions
compatible with the target object. To illustrate the different
behavior, consider a straight wire carrying current J, representing
an elementary radio antenna. The magnetic field created by this
current encircles the wire and decays reciprocally with distance.
Bending this wire to a loop results in a prototype NMR probe
(Fig.l): inside the loop the field is reinforced by contribution of all
parts of the wire, while outside the loop it is "diluted". Within the
loop plane the magnetic field is fairly constant; outside it decays
much more rapidly than the field of the straight wire.
Both in radio-technology and in NMR the transmission
antenna is principally suitable for as a reception antenna as well.
In both cases the major difference is that the transmission antenna
has to handle powers in the range from 1W to several kW, while
the reception antenna should receive as little as pW to mW powers.
In the radio domain this power drop is simply due to geometric
"dilution". The solid angle at which the reception antenna appears
from the transmit antenna, and thus the power partition, drops
very rapidly with distance. In NMR there is ideally no energy
radiation into space; but energy is stored and concentrated in the
confined volume. The power continuously delivered to the NMR
probe during a pulse does not, however, lead to infinitely strong
fields in the sensitive volume, as resistive coil and sample losses
increase with field energy density. Equilibrium is reached when
the coil and sample losses match the power delivered per cycle. In
both directions, to and from the spin system, energy transport is
taxed very heavily by energy dissipation in the antenna structure
and in the sample. Thus, the energy available to perturb the
desired spins is unfortunately only a small fraction of the total
energy delivered to the system. Correspondingly, only a minor
part of the total spin energy released upon relaxation is available
to the receiver unit.
5
In many biomedical applications the transmission and
reception functions are accomplished by two different probes: a
large resonator (body coil) is used to excite the whole sample space
evenly. A smaller reception coil which focuses its sensitive volume
to the region of interest is applied to achieve higher sensitivity at
the cost of a more inhomogeneous sensitivity profile. This
approach introduces distinct problems resulting from the mutual
coupling of transmission and reception antenna, as has been
discussed in [12]. But apart from the electronic decoupling
circuitry to be introduced into the probe structures, this approach
does not alter the design goals of resonators and does not lead to
design differences between transmission and reception probes.
2.2 RECEIVER COIL
An oscillating magnetic dipole, positioned at point q in space
relative to a loop of wire, will induce a time-varying voltage in the
loop whose intensity is directly related to the strength of the
magnetic field B] that would be produced at point q by unit
current (DC) in the loop [13,14], This relationship between the
induced signal voltage and the (hypothetical) magnetic field
produced by unit current, known as the principle of reciprocity
[13], describes the oscillating signal voltage S(t)q induced in a NMR
receiver coil by a pressing magnetic moment M at q:
6
where B[ is the magnetic field produced by the unit current
flowing through the receiver coil.
For the NMR experiment with the static magnetic field Bo in
the z direction, the components of Bi and M which contribute to a
NMR signal are those in the transverse, or xy plane. Evaluating
Eq.(l) for transverse magnetization, the signal induced in the NMR
coil from a moment at point q can be written as
C2>
S(0q - (Bi)xy{q)Mzy(q)aJ0 sm (co0r 4- 1 radius result from reversal in the direction of B jx relative to
that of the homogeneous transmitter. While the uniform flip angle
results in increased signal intensity in the region lying within
approximately on radius of the x axis for the homogeneous-
excitation experiment relative to the single-coil experiment, the
dependence on B jx rather than Bixy results in reduced sensitivity
laterally outside this region[16].
The uniform flip angle over the sample volume also results
in an increased penetration depth relative to the single-coil
experiment. In the homogeneous-excitation experiment 90% of the
total signal is obtained from the sample volume located within 2.8
radii of the coil plane. By comparison, for the single-coil
experiment 90% of the signal is obtained within a depth of only
one radius[ 16,21]. The effects of the uniform flip angles and phase
mismatches result in a narrower and deeper interrogated volume
of the surface coil receiver. However, the total signal intensities
for the two experiments are similar for a single pulse at
equilibrium [16].
One important application of surface coil reception with a
homogeneosly excited sample is found in imaging[ 19.22]. This
application takes advantage of the high sensitivity over a localized
1 6
region afforded by the surface coil receiver, while optimizing
experimental conditions for imaging by providing uniform flip
angle excitation. In order to resolve spatially the NMR signal
contributions in the imaging experiment, each volume element is,
in effect, detected independently of the others. Thus assuming
that the volume element is small relative to the receiver B jxy
inhomogeneities, there is no signal cancellation due to
transmitter/receiver phase mismatch[16]. The image signal-
intensity distribution of the surface-coil receiver in the presence of
homogeneous excitation is, therefore, identical to the B jxy
distribution of the surface coil receiver, as plotted (Fig.2). Since no
signal cancellation results from transmitter/receiver phase
mismatch, the signal intensity is a function of Bixy instead of just
B jx, and the xy and xz signal distributions are no longer identical.
Comparison of the homogeneous excitation, surface coil
reception imaging experiment (equivalent to Fig.2) with the
homogeneous excitation, surface coil reception spectroscopy
experiment (Fig.4) shows that except for the x=0 and y=0 planes,
where the two experiments are equivalent (in magnitude), (a) the
interrogated volume of the imaging experiment extends further in
every direction than the spectroscopy experiment, and (b) the
image signal intensity is greater than the spectroscopy signal at
every point over the sample volume. Since the magnitude of the
signal is generally displayed in the imaging experiment, no
negative signals are produced.
I 7
In the above discussion, only homogeneous excitation with
surface coil reception has been considered. However, the
transmitter coil need not be homogeneous. For example, coaxial,
coplanar surface coils of different diameters have been used as
separate transmitter and receiver coils, where the larger diameter
coil is used for excitation and the smaller diameter coil for
reception[23]. Although a single, optimum flip angle is not
achieved over the entire sample, the flip angle distribution created
by excitation through the large coil is broader than that which
would be created through excitation by the small coil. In this case,
signal cancellation due to phase mismatches between the
transmitter and receiver still occur, although it is less severe than
the homogeneous excitaiton experiment[20,23,24], This scheme
has found useful applications, particularly in conjunction with
depth pulse sequences, in enhancing signal localization achievable
using the surface coil[17,20,23,25].
2.6 SENSITIVITY OF COIL
The sensitivity of the NMR receiver is quantified by the
signal -to-noise (S/N) ratio per unit sample volume obtained in the
NMR experiment [26]:
M 0{1 — e~T w y T l)
(1 — cosae-1*'7') ’
(9)
1 S
where Bjxy is the transverse field produced by unit current in the
receiver coil, W is the Larmor frequency, and R is the equivalent
series resistance representing the thermal noise generated in the
coil and sample. The noise associated with dielectric and inductive
(eddy current) losses in electrically conductive biological samples
usually predominates at high frequencies. This sample noise is
t
coupled to the receiver coil through the electric and magnetic
fields of the coil, and degrades the signal to noise ratio in the NMR
experiment. While dielectric losses can be reduced by using
circuits designed to minimize stray electric fields and by using a
Faraday shield to screen the electric field of the coil[27], little can
be done to eliminate inductive, losses. Rigorous analyses of the
effects of such loss mechanisms, as well as other factors affecting
receiver sensitivity, are treated in detail elsewhere[26-31].
The signal to noise ratio obtained for a volume element
located within the detection volume of a surface coil is related to
the Bjxy at that position. Along the axis of a surface coil B]xy is
related to the coil radius r as[29]
- e ~ T R rr’) "
(1 — cos a e -TR'Tl)
since. ( 1 0 )
where x is the axial distance from the coil Bixy is strongest at the
center of the surface coil, where it is inversely proportional to the
coil radius, and decreases rapidly with distance from the coil.
Sample volume elements close to the surface coil receiver (in
regions of high Bjxy) contribute strongly to the total signal as well
1 9
as to the total noise, while volume elements far from the coil
contribute markedly less to both. The effective resistance
associated with a conductive sample filling the detection volume of
a surface coil (small r) will have a larger sensitivity for a volume
element at the center of the coil than will a large surface coil.
However, the sensitivity of the smaller coil drops off more rapidly
with increasing distance from the coil than that of the larger
surface coil. Beyond some frequency-dependent depth, the
sensitivity of the surface coil becomes nearly independent of coil
radius[31], As a result of their strong Bjxy, surface coils have
much higher sensitivities than large volume homogeneous coils for
regions near the sample surface. However, due to the rapid fall-off
of Bjxy of the surface coil, large volume coils outperform surface
coils performance for regions deeper within the sample[29,30].
2.7 COIL LOCALIZATION
While the surface coil is well suited for signal detection from
a thick, superficial tissue layer, the curved iso-Bjxy contours and
intense B ixy near the coil make it difficult to detect signals from
deeper structures without contamination from the overlying
regions. Elimination of signal from superficial tissues is
particularly necessary for insitu spectroscopy of organs such as
liver and kidney which lie beneath a thick muscle layer. Selection
of a 180° flip angle at the sample surface reduced contributions
from the overlying region, but is insufficient to eliminate them.
Numerous investigators have resorted to surgical exposure of such
tissues to isolate signal detection to only the tissue of interest.
This compromises the otherwise non-invasive nature of NMR
spectroscopy. A number of elegant techniques have been
developed and applied to surface coil spectroscopy to improve the
inherent volume localization capabilities of the surface coil. These
techniques utilize gradients in either the static Bo field or the RF Bj
field to delineate a specific sample volume. Many excellent
reviews of localized spectroscopic methods have been published
(see for example[23, 32-35]).
Frequency selective RF pulses applied in the presence of
linear BO field gradients have been used to select a well defined
slice or volume of interest. Data are acquired in the absence of
applied gradients yielding high resolution spectra. Spins from
outside the volume of interest are either not excited, rapidly
dephased in the gradient fields, or subtracted from subsequent
data collections. Examples of such techniques include volume
selective excitation (VSE)[36-38], image selected in vivo
spectroscopy (ISIS)[39,40], and depth-resolved surface coil
spectroscopy (DRESS)[41.42]. Incremental pulsed field gradients
and subsequent multidimensional Fourier transformation have
been used to produce spectra from a series of spatially resolved
slices[43-45]. In the topical magnetic resonance (TMR)
technique[45,47], nonlinear gradients are used to profile the static
magnetic field to produce a homogeneous localized region from
2 1
which a high resolution spectrum is obtained. Difficulties
encountered in the localization schemes employing Bo gradients
include the possible generation of long lived eddy currents by the
switched field gradients and spatial smearing of the localized
volumes resulting from chemical shift dispersion (chemical shift
artifacts).
The natural B] gradients of the surface coil RF field have
been exploited to enhance signal localization. Rotating frame
imaging[48-50] and related techniques[51-53] acquire as series of
spectra in which the excitation pulse width is incremented,
thereby mapping the spatial distribution of the chemical species
according to Bi gradient contours. Composite pulses have been
used in depth selective sequences for enhancing spatial selectivity
and reducing off resonance phase distortion[54-57]. A family of
sequences known as depth pulses makes use of a series of phase
cycled pulses that result in spatial discrimination based on
sensitivity to the flip angle experienced[17,20,23,25]. Cascading of
depth-selective pulses further increases spatial selectivity. Fig. 10
illustrates the increased spatial discrimination achieved by a
simple ©; 20 [+x, ±y] depth pulse sequence (i.e. a pulse producing
flip angle 0 , in this case equal to 180° at the coil center, followed
by a phase cycled pulse producing flip angle 2 0) over that
achieved by a single pulse with a 180° flip angle at the coil center.
Since the shape of the localized volume selected by these
sequences is determined by the curved Bj contours of the surface
coil, the signal produced by the selected volume will still be
22
contaminated by the response of superficial tissues. Additionally,
alternating positive and negative signal contributions from a
sample in the high-flux regions close to the coil wire can be
significant. Both of these effects can be reduced by using Bo
gradients in addition to the depth selective sequences[ 17,41], or by
using separate surface-coil transmit and receive coils whose
excitation and detection volumes overlap[20, 25]. High flux signals
near the coil have also been suppressed by varying the excitation
pulse lengths in the depth pulse sequence or by including low flip
angle preparation pulses in the sequence[17, 59-61]. Severe off-
resonance effects and lengthy pulse trains can complicate Bj
gradient techniques.
An alternative approach to the elimination of superficial
signals that is not plagued by residual eddy currents, curved
excitation volumes, or high flux signal contributions, is the use of a
surface spoiling gradient in which a local (surface) gradient coil is
positioned directly over the unwanted tissue region. A highly
inhomogeneous Bo gradient field is produced which rapidly
dephases signals within a limited penetration depth [62,63].
Hence, these sample regions will contribute no net signal in
response to subsequent surface coil interrogation. This technique
has been demonstrated in vivo using simple single pulse sequences
and spin echoes [64]. Paramagnetically and ferromagnetically
generated spoiling gradients have also been used to eliminate
superficial signals [65-67].
23
2.8 FIELD INSENSITIVE PULSES
While the inhomogeneous Bjxy of the surface coil can be
advantageous in signal localization schemes, it can be an
impediment in other spectroscopic techniques, especially those
which depend on uniform signal excitation. Since an optimum flip
angle cannot be achieved over the entire sample with surface coil
excitation, the maximum signal-to-noise (S/N) obtainable in the
experiment is limited. To improve sensitivity and minimize
spectral distortions caused by the inhomogeneous Bjxy, phase
cycled composite pulses have been developed which are tolerant
of wide variations of Bjxy and resonance offsets[68]. In contrast to
the composite pulses used in localization schemes, which increase
the sensitivity of the signal to variations in Bjxy, these pulses
produce a significantly more uniform excitation profile over a wide
range of Bjxy values than achievable with simple pulses[54,70-72],
Adiabatic pulses, see (Fig.5), in which the amplitude and frequency
(or phase) of the pulse are modulated to achieve a more uniform
excitation response, also exhibit marked insensitivity to wide
variation in the Bjxy field[73-75].
In addition to increasing the signal-to-noise ratio obtainable
through more uniform excitation, field-insensitive pulses such as
the composite and adiabatic pulses have the added advantage that
they can be used with a single surface coil operating as both
transmitter and receiver. Hence, the excitation and reception
fields will have a constant phase relation over all space. These
24
pulses have been useful for signal excitation, inversion, and
refocusing in surface-coil experiments such as spin-echo formation,
imaging, localized spectroscopy, and relaxation-time
m easurem ents [7 6-80].
3 SURFACE COIL DESIGN PRACTICALITIES
3.1 INTRODUCTION
The basic surface-coil design consists of a circular loop of
wire composed of one or more turns, wound either in a cylindrical
or planar fashion. The coil is turned to the NMR frequency of
interest and matched to the impedance of the driving device,
typically 500. One standard tuning and matching scheme is the
capacitively-coupled resonant tank circuit shown in Fig.6a[13].
Degradation in circuit performance resulting from electrical
coupling to lossy conductive samples can be reduced by balancing
the NMR coil with respect to ground[81-83]. An example of a
capacitively-coupled balancing scheme is shown in Fig.6b. The
single-resonance surface-coil probe has been extended to
applications in multi-nuclear NMR spectroscopy through
development of multiple-frequency circuit-design technology[84-
94]. The use of resonant transmission lines and lumped reactive
25
elements in multiply-tuned circuits is discussed in detail
elsewhere[84-89].
Modifications of the circular surface-coil design have been
motivated by a desire to better accommodate the sample size or
shape, to enhance the signal localization, or to improve the
sensitivity of the coil.
NMR coils that are balanced with respect to ground (through
capacitive coupling[81] or inductive coupling[82] have been shown
to be effective in reducing dielectric losses in conductive samples.
Dielectric losses have also reduced by symmerically distributing
lumped capacitors around the coil to reduce the potential
difference between the coil and ground[19, 82]. Explicit grounding
at the coil center and electric shielding of the sample from the coil
are features of the "crossover" surface coil which minimize
dielectric losses in the sample[83].
An important class of coils used for local signal detection is
based on the loop-gap resonator design[98]. These low-loss, high-
sensitivity coils consist of wide conductor bands that are brought
to resonance with capacitive gaps. Two coplanar loops that are
laterally displaced and connected by a single gap from the planar-
pair loop-gap resonator[99]. This coil detects signals from a slab
shaped region close to the coil and is useful for imaging superficial
structures such as the temporomandibular joint. A further
advantage of this design when used as an imaging coil is that the
two loops are intrinsically isolated from an external homogeneous
field of arbitrary orientation, such as a homogeneous transmitter
26
field. This results from the fact that, for the proper circuit
resonance condition, an emf induced at the gap by flux linking one
of the loops will cancel that induced by flux linking the other loops.
The counter-rotating-current coil consists of two coaxial loops that
are axially displaced and support current flows in opposite
directions[100]. It, too, is intrinsically isolated from an external
field. Combinations of the planar-pair and counter-rotating-
current coils produce a resonant structure whose elements are
intrinsically isolated both from each other and from an external
transmitter field[l01 ]. Additionally, the signals induced in the two
coil components can be detected in quadrature and combined to
produced a net V2 improvement in the signal-to-noise ratio[102].
3.2 SHIELDING
The purpose of shielding is to prevent the NMR resonator
from behaving as an antenna, by partially or completely enclosing
it in a more or less perfectly conduction "cage", with the aim to
confine its field energy to some finite volume. This prevents
radiation losses, sensitivity to outside noise sources (such as radio
stations) and interaction with lossy or non-ideal outside materials.
The perfect shield acts as a "mirror", such that all currents flowing
in the antenna create oppositely flowing currents in the shield.
Since the total external flux created by the currents flowing in the
antenna structure thus cancels to zero (more precisely: the
27
external dipole characteristics is transformed into quadrupolar
characteristics with correspondingly steeper field decay), there is
no effective field and no energy stored outside the shield. This
effect is directly comparable to the working principle of gradient
shielding. Well-designed rf shielding structures allow high-
frequency currents to flow at minimal resistance, while low-
frequency currents, such as those induced by switching gradients
(eddy currents), should find a high-resistance path. Otherwise the
probe interior would be shielded from the gradient fields as well.
Therefore, if gradient fields have to penetrate, it is recommended,
to not use plain copper foil, but coarse copper mesh or preferably
"comb" structures that do not offer DC-current loops but
capacitively connected loops only.
3.3 COUPLING
The coupling mechanism is responsible for feeding
transmitter power to the probe to continuously replace the energy
lost in the sample and probe during a pulse, and to transfer signal
energy picked up by the probe to the receiver. Energy can be
transferred to the probe where it provides sensitivity. It can be
coupled electrically (capacitively) to positions where the probe's
internal electric fields are high (such as capacitances, whether as
lumped elements or implicit in the probe structure). Alternatively,
28
energy can be transferred by magnetic fields (inductively) at
positions of high intrinsic magnetic field.
Capacitive coupling to a single current loop such as a surface
coil is illustrated in (Fig.6). The driving voltage is applied to the
probe's capacitor through an adjustable matching capacitor. The
external matching and the internal tuning capacitor act as an
imaginary "voltage divider", where Ct determines the resonance
frequency, and the ratio C H /f C H + C m - 1) determines the partial
driving voltage and the coupling strength. More rigorous concepts
of coupling as an impedance matching process have been
developed [106,107].
It must be emphasized that tuning and matching are
basically independent functions. In many cases, the "voltage
divider" is set up with a variable matching capacitor which is
connected to a fixed capacitor of the probe, while probe tuning is
achieved by one or more variable capacities elsewhere in the
probe.
In the coupling scheme of (Fig.6a) one end of the coil is kept
to feed ground, while the other end assumes maximum electric
potential. The loss effect due to this coil potential can be strongly
reduced if a balanced feeding scheme is employed (Fig.6b). In this
symmetrical scheme the coil is uncoupled from ground and
assumes a virtual ground potential in its center, such that the
electric potentials at both coil ends and the corresponding energy
losses are reduced by a factor of two[107]. Virtual ground
positions can easily be detected by touching the coil, connected to a
29
sweep generator, with a finger tip. The higher the electric
potential, the more pronounced is the shift and widening of the coil
resonance in the wobble spectrum. Touching the coil at a virtual
ground position barely affects the appearance of the resonance.
Inductive coupling is less frequently used in combination
with surface coils, but it is instructive to demonstrate the principle
with this simple example (Fig.7). The primary surface coil
resonator, tuned to the desired resonance, is approached by a
coupling surface coil resonating in the vicinity of the main coil
resonance. The primary coil is not connected to ground; therefore
inductive coupling provides intrinsic balanced feeding. At some
distance, the sensitive volumes of both coils barely overlap, and
practically no energy is transferred from the driving coil to the
main resonator (Fig.7a). If the coupling coil approaches the main
coil, the magnetic flux common to both coils is increased and the
coupling becomes stronger. This behavior can be monitored if the
coupling loop is connected to a sweep generator. Closer coupling
corresponds to less power reflected from the coupling loop at the
resonance frequency of the primary circuit, as the power is
absorbed by the primary loop. Perfect match is achieved when all
power is absorbed by the primary coil, and no power is reflected
into the transmitter at the desired resonance frequency.
Alternatively, coupling strength can be increased by bringing the
resonance of the coupling coil closer to the main coil's resonance
frequency (Fig.7b), again increasing the flux from the coupling coil
to the main coil in the desired frequency range.
3 0
In both cases of capacitive or inductive coupling, the coupling
device can be viewed as a second resonating structure overlaid to,
and providing the energy loss by, the main resonator. With
respect to field homogeneity, this has no consequence if the
symmetry of the coupling device matches the main coil symmetry.
Probes that can be characterized by a single current loop (no
parallel paths in the magnetic field-producing conductors) are best
coupled capacitively, leaving the symmetry of the current
distribution in the inductors unaltered. Inductive coupling by a
surface coil-type mechanism inevitably overlays the surface coil
field characteristics onto the main coil field distribution. In low-
loss probe/sample combinations only weak coupling is necessary;
the coupling coil provides only a small percentage of the magnetic
field energy stored in the probe. Thus, the symmetry break is
negligible. High loss combinations require high energy flow
through the coupling device, and thus result in higher
inhomogeneity.
3.4 L-C RATIO
There exists an unlimited number of possibilities for the
construction of a 3 cm 19F surface coil at 1.5T: A single turn thick
wire coil in combination with a fairly high capacitance of more
than 30pF will lead to the desired resonance at 59.8 MHz, as well
3 1
as a 2 turn copper wire coil with a correspondingly lower
capacitance of approximately lOpF.
Two criteria lead to a narrower range of possibilities: In
principle, one wants the lowest possible inductivity of the coil in
order to achieve high currents at low voltages corresponding to
high magnetic fields at low electric fields. This requires high
capacitance values, which are also favorable with respect to
detuning due to changing the sample or patient movements, On
the other hand, one wants to minimize losses in the coil's
connection leads which cannot be arbitrarily shortened. If the
connection leads length is 1/3 of the total length of the coil wire,
roughly 1/3 of the magnetic energy does not contribute to the
desired sample interaction, but is dissipated in the coil. Therefore
one has to optimize towards short and thick probe wires, that are
still substantially longer than the connection leads. It must be
strongly emphasized that long connection leads (relative to the
total coil wire length) lead to serious degradation of probe
efficiency.
3.5 SENSITIVITY VS. HOMOGENEITY
A good resonator is characterized by both high sensitivity
and high field homogeneity. Homogeneity is optimized by evenly
distributed currents arranged with high symmetry. Examples are
the solenoid and the birdcage resonator, with high longitudinal and
32
pronounced rotational symmetry. The relationship between
symmetry classes and field homogeneity properties for all kinds of
magnetic field generating devices (applicable to the design of main
magnets as well as gradient coils) was elaborated in general
framework by Romeo and Hoult [110]. Sensitivity is optimized for
a given resonator and a given sample (assuming all electric loss
factors have been eliminated) if the sensitive volume of the
resonator is identical to the dimensions of the sample (or the
region-of-interest to be measured).
If a sample of cylindrical dimensions, such as a solution in an
NMR tube, is to be spectroscopically analyzed, the use of a
matching solenoid results in almost perfect homogeneity across the
sample at simultaneously maximum sensitivity. Conflicts between
sensitivity and homogeneity arise, if the sample is much larger
than the desired region of interest, e.g., if a transverse thin slice
image of the sample is to be acquired. A much longer solenoid
than the desired slice thickness shows good field homogeneity
across the slice, but it will pick up noise from parts of the sample
not contributing to the signal and degrade sensitivity. If, on the
other hand, the solenoid is shortened to the desired slice thickness,
its sensitivity is strongly enhanced at the cost of decreased
homogeneity.
Bottomley et al.[103] have elaborated that the sensitivity of a
birdcage resonator is increased by 2.6, if instead of the
length/radius of 4, which yields optimum homogeneity, a
length/radius ratio of 1.4 is used. This decrease in homogeneity is
33
less severe for imaging of transversal slices or spectroscopic voxel
localization than for longitudinal slices. Hoult[106] showed that the
transformation of a two turn surface coil into a Helmholtz coil
decreases sensitivity by 13.4% with a corresponding gain in
longitudinal field homogeneity. In both cases this change in
sensitivity is directly related to the change in the sensitive volume
of the coil. It is therefore recommended to design resonator
geometries specific to the desired samples, i.e., with
sensitive/homogeneous regions not extending substantially outside
the typically desired field-of-view dimensions. At this point one
has to consider if employing a large transmitter probe with
homogeneous transmit fields and a separate small reception probe
yielding optimal sensitivity would be preferable. This choice,
however, calls for measures of active or passive probe decoupling
[12].
SURFACE COIL EXPERIMENT
4.1 COIL REQUIREMENTS
The essential steps for obtaining values for 19F metabolite
concentration are:
34
1. To transmitte and receive at the 19F frequency and, also to
able to pick up 1H frequency signal for imaging,
2. Total depth of signal penetration to be selective via coil size,
3. To determine the relative position and orientation of the
chosen VOI (as used in a clinical MRS study) within a suitable
sensitivity map of the coil,
4. To integrate the sensitivity over this volume,
5. To use this information to scale the spectral peak
amplitudes, together with factors to account for coil
loading. The measured signal from a reference sample
in a fixed position may then be used to convert relative
peak amplitudes into metabolite concentrations.
6. To safeguard against excessive RF radiation of patients.
4.2 CIRCUIT DESCRIPTION
The double resonant design (Fig.8), is adapted from a single
frequency coil that makes use of semi-rigid coaxial line segments
to distribute capacitance symmetrically about the coil, thereby
reducing its peak rf electric "E" field by establishing a bilateral "E"
field balance about a virtual ground "0" central to and aligned with
the lead axis of the coil. This ground potential of the coaxial shield
eliminates coil tuning and match instabilities resulting from sheath
currents.
35
The centered ground minimizes "E" field coupling to the sample as
well. Lower peak voltages in the coil circuit result in an improved
peak power rating for this design. Viewed as two integrated loops,
this coil design is inherently double resonant. The high frequency
loop is composed of the solid conductor "c" capacitively coupled
with the coaxial line shields which in turn are connected at "d" to
complete the circle. This capacitive splitting of the high frequency
loop serves to increase the frequencies and diam eters achievable,
as well as to distribute more evenly the short wavelength rf
current for improved Bi field homogeneity. Variable capacitors "a"
and "b" provide for tuning and m atching of the high frequency
mode, and for electrical balance adjustm ent for the entire
structure. The low frequency resonance is established by the more
inductive loop consisting of the solid conductor connected to the
center conductors of the coaxial line elements of the coil. Tuning,
matching, and balance of the low frequency is performed by the
more conventional circuit shown. Although capacitance
adjustm ents to the two loops are somewhat interactive, two
resonances can be independently optim ized for m utual tuning,
matching, and balancing over a range of load conditions, This coil
can be divided into more segments to increase further diam eter
and/or frequency. The design is easily extended to volume coils of
the Helm holtz type.
EQUIPMENT USED
Invivo NMR 1.5T Magnetom
Siements M edical Systems Inc.
186 W ood Avenue South
Iselin, NJ 08830
N etw ork/Spectrum A nalyzer (HP 4396A)
H ew lett-Packard Co.
5161 Lankershim Blvd.
P.O. Box 3919
No. Hollywood, CA 91609
Signal Reflector Circut
ANZAC Adams-Russell Co. Inc.
80 Cambridge St.
Burlington, MA 01803
PARTS USED
Copper tubing O.D.=0.5 cm
Ceram ic Capacitors (fixed) (non-ferrom agnetic)
RF/M icrow ave Capacitors
Values used from lpF to 600pF
American Technical Ceramics Corp.
One Norden Lane
Huntington Station, NY 11746
Ceramic Capacitors (variable) (non-ferromagnetic)
RF/Microwave Capacitors
Values used lpF-15pF (variable)
lpF-30pF (variable)
Johanson Mfg. Corp.
Rockaway Valley Road
Boonton, NJ 07005
RF Cables (50 ohms) (non-ferromagnetic)
Alpha Wire Corp.
711 Lidgerwood Ave.
P.O. Box 711
Elizabeth, NJ 07207
4.3 COIL TESTING (BENCH WORK)
Set at a center frequency of 63.88 MHz with a range of + 20.0
MHz, HP 4396A (Network/Spectrum Analyzer) was used to tune
and match the surface coil. A refractance circuit was set up by
means of a 1.0 liter NaCl 9.0% doped solution used as a load to
measure the Q of the coil. All output was printed with a HP bubble
jet printer (Fig.9)-
38
4.4 COIL TESTING (IN-VIVO)
Experiments were performed with a loading phantom to
ensure that there was no change in the sensitivity profile of this
coil when loaded. When applying this technique to other coil
designs, it would be prudent to ensure that there is no effect on
the coil sensitivity profile. This method assumes a uniform
distribution of metabolites within the VOI and does not explicitly
include correction for losses during the localization process. These
effects will have to be considered subsequently.
All experiments were performed in a Siemens 1.5 T
Magnetom system. The coil map and spectroscopy measurements
were acquired with a 3 cm 19F/1H transmit/receive dual
frequency transmission line resonator (TLR) surface c o il(lll). A
proton marker ring containing water doped with NaCl and CuS04 to
give a T1 of approximately 260 ms was attached to determine the
precise coil position in all images, because an essential step in our
quantification process is an accurate geometrical location. The
pulse used for all our measurements was a single 10 ms adiabatic
rapid half-passage excitation pulse(112) with a bandwidth of 2.5
kHz. This ensured that the excitation was unaffected by the spatial
variation of the B i produced by the surface coil and therefore
39
essentially independent of position. Also, this pulse was used to
obtain the clinical data for which this technique developed.
4.5 COIL MAPPING
A coil map was obtained to determine the sensitivity
variation of the coil with a position assuming uniform excitation of
spins. For this reason the same large phantom was used to
produce an identical solution to that in the proton marker ring.
This phantom had to be large enough to extend beyond the extent
of the surface coil so that an accurate sensitivity coil map would be
obtained. The TLR was switched to mode (i.e., tuned for *H
nuclei) with the concentric proton marker ring attached (Fig. 10).
The 3D coil map was obtained by means of a 3D FLASH
gradient echo sequence and the pulse described earlier. A 256
matrix size was used with a repetition time of 2.0 secs, (which is in
accord with the specification that TR must be at least 5 times Tl).
4.6 VOI DETERMINATION WITH RESPECT TO THE COIL
To correct for the sensitivity of a surface coil to a given VOI,
it is necessary to know their relative spatial positions and angles .
For localized spectroscopy, the VOI position is usually defined on a
series of images relative to the coordinate system of the magnet.
4 0
To identify the coil position in both patient images and coil
sensitivity maps, a "proton marker ring" is rigidly mounted on the
RF coil. In MR images this appears as a pair of dots whose
separation varies according to where the image plane intersects
the ring. The position of the proton marker ring is determined for
both the patient images and the coil map data set using a curser to
define the approximate position of both of the marker ring dots in
each image slice. The program then performs a center of gravity
calculation to determine the precise center of the proton marker
ring dot and goes on to calculate the coordinates of the center of
the proton marker ring, the angle of the coil with respect to the z
axis and the offset in the z direction.
5 IN VIVO STUDIES
Patients are placed in a MRI unit and imaged to verify the
proper positioning of the tumor to be studied in relation to the
custom-designed 3-cm 19F surface coil. A background spectrum is
acquired to verify the absence of any prior 19F signals.
An external reference standard of 1,2-difluorobenzene (DFB),
placed inside the surface coil, is used for chemical shift verification
and for quantification.
The dose of 5-FU is administered, and serial spectra are
acquired over 4.7 min. periods using both localized chemical shift
imaging (CSI) and unlocalized (global) sequence. For the global
4 1
sequence, a total of 256 free induction decays (FIDs) are collected
in each acuisition period with a repetition time (TR) of 1 sec, a
pulse width of +2000 Hz, a vector size of 512, and an adiabatic
half-passage radiofrequency pulse to minimize differences of
detection due to distance from the surface coil[113].
Fig. X
The magnetic fields created by a circular wire, representing a
prototype NM R probe
6 FIG U R ES
4 2
2.0
X
ta
to
100
• 2 * 1 0 1
2
Y
a
2.0 -
X
0.5
100,
120
2 0 • 1 ■ 2
b z
Fig. 2
M agnitude of calculated (Bi)xy as a function of spatial coordinates
for a 1-turn, circular surface coil radii. . The (B i)xy intensity is
normalized relative to a value of 100 at the coil center.
v (a) The xy plane at z=0
(b) The xz plane at y -0
43
2.0
1.S
X 1.0
0.5
joa
0
110 n o
1 0 - 1 ■ 2
Y
Fig. 3
(a) The xy plane at z~0
(b) The xz plane at y=0
44
2.0
X 1.0
to
0.5
120 WO ■ S O ISO
■ 2 - 1 0 1 2
Fig. 4
The xz plane at y=0
45
S ign al , , S ig n a l
D ept.T P u lse
*50 -
100 -
Single P ulse
-100
2
Fig. 5
Adiabatic pulse profile
46
b
Fig. 6
(a) Capacitive tuning and matching, viewed as an imaginary
voltage divider network
(b) Symmetric matching capacities uncouple the coil from feed
ground and reduce capacitively induced sample losses
47
< u
u
c
a
tj
0 1
W e a k c o u p lin g
a;
0
!
fo
S tro n g co u p lin g
< u
(X.
0
* o
Frequency
Fig. 7
Coupling strength can be increased by spatially approaching the
main resonator with the coupling coil, to increase the flux common
to both coil, or alternatively by tuning the coupling coils resonace
frequency towards the main coils resonance to increase the
coupling coils' energy content at the desired frequency
48
Fig. 8
Double tuned surface coil with associated fixed and variable tuning
elements and 50£2 coaxial transmission line
ENTER
■■SPftN
CENTER
Fig.
3 cm coil profile with load
(a) at 63.9 MHz
(b) at 53.9 MHz.
Print out from N etw ok/Spectrum Analyzer
Water _ _
phantom
Coil
su p p ort
Proton
marker
rina
Coi!
m ounting
Patient table
Fig. 10
Experimental arrangem ent for acquisition of the coil map
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University of Southern California Dissertations and Theses

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Asset Metadata

Creator Kim, Hyun Kwon (author)

Core Title Flourine-19 NMR probe design for noninvasive tumoral pharmacokinetics

School School of Engineering

Degree Master of Science

Degree Program Biomedical Engineering

Degree Conferral Date 1995-04

Publisher University of Southern California (original), University of Southern California. Libraries (digital)

Tag engineering, biomedical,OAI-PMH Harvest

Format masters theses (aat)

Language English

Contributor Digitized by ProQuest (provenance)

Advisor Singh, Manbir (committee chair), Maarek, Jean-Michel. (committee member), Wolf, Walter (committee member)

Permanent Link (DOI) https://doi.org/10.25549/usctheses-c18-2314

Unique identifier UC11356844

Identifier 1376469.pdf (filename),usctheses-c18-2314 (legacy record id)

Legacy Identifier 1376469-0.pdf

Dmrecord 2314

Document Type Thesis

Format masters theses (aat)

Rights Kim, Kyun Kwon

Type texts

Source University of Southern California (contributing entity), University of Southern California Dissertations and Theses (collection)

Access Conditions The author retains rights to his/her dissertation, thesis or other graduate work according to U.S. copyright law. Electronic access is being provided by the USC Libraries in agreement with the au...

Repository Name University of Southern California Digital Library

Repository Location USC Digital Library, University of Southern California, University Park Campus, Los Angeles, California 90089, USA