j6
Hyperpolarized Noble Gas Magnetic Resonance
Imaging of Human Lungs in 150G Magnetic Field
by
Adelaide Zhang
Submitted to the Department of Electrical Engineering and Computer
Science
in partial fulfillment of the requirements for the degree of
Master of Engineering in Electrical Engineering and Computer Science
at the
MASSACHUSETTS INSTITUTE OF TECHNOLOGY
August
2 00
BARKER
©2000 Adelaide X. Zhang. All rights reserved.
The author hereby grants to M.I.T. permission to reproduce and
Distribute publicly paper and electronic copies of this thesis
And to grant others the right to do so.
I
MASSACHUSETT 8NSTITUTE
OF TECHNOLOGY
JUL 3 12002
LIBRARIES
Department of lectrical Engineering and Computer Science
August 14, 2000
A .
.
Certified by...
-........................................
Mitchell Albert
Thesis Supervisor
..........................................
Frederick Bowman
C ertified by.. /............
T)esis Supervisor
Accepted by..... .............
..
................
.
.
.......
Arthur C. Smith
Chairman, Department Committee on Graduate Student
Hyperpolarized Noble Gas magnetic Resonance Imaging of
Human Lungs in 150G Magnetic Field
by
Adelaide Zhang
Submitted to the Department of Electrical Engineering and Computer Science
on August 14, 2000, in partial fulfillment of the
requirements for the degree of
Master of Engineering in Electrical Engineering and Computer Science
Abstract
For the past 40 years, Magnetic Resonance Imaging (MRI) has been the most common
method used to obtain volume images of human organs. This technique involves the
detection of proton nuclear spin in cells at 1.5 Tesla (T) magnetic field strength.
However, conventional MRI is very costly and some organs in the human body, such as
the lungs, do not contain sufficiently high concentrations of protons to produce high
quality images. In 1994, Hyperpolarized Noble-Gas Magnetic Resonance Imaging (HPMRI) was introduced and is viewed today as the up and coming technology that will
solve this image scanning dilemma [1]. Its special features include a signal to noise ratio
(SNR) is independent of the static field strength and Noble gas hyperpolarization by
optical pumping which is 100,000 times larger than that is obtainable at thermal
equilibrium. Imaging has been done using 'He and 129 Xe at 1.5T magnetic field, but no
images have been produced at a low magnetic field, even though the magnetization of
HP-MRI is independent of the static field. To prove that HP-MRI theory can be used in
practice and that it has potential as a lower cost technology, human sized coils were
constructed for lung imaging using hyperpolarized noble gases at 150 Gauss (15 mT)
magnetic field. In order to operate the HP-MRI, a superconducting magnet at 1.5 T was
ramped down to 150 Gauss using a new set of broad-band electronics that can be used at
different Larmor frequencies. Comparison tests between different human sized coils
were done to optimize their function. Images of rat lung were produced in vivo using 3He
and 129 Xe and human lung images were produced using 3He at 150 Gauss to suggest that
the theoretical hypothesis is correct, and to opens possibilities for clinical HP-MRI in
Very Low-field.
Thesis supervisor: Mitchell Albert
Title: Project Supervisor
Thesis Supervisor: Frederick Bowman
Title: MIT Thesis Supervisor
2
Acknowledgments
I would like to thank my thesis supervisor at Brigham and Women's Hospital, Dr.
Mitch Albert, who was always there to advise me and to provide supports. Thanks to
Arvind Venkatesh, a Ph.D. student, who was so willing to help me to solve any technical
problems and who was always performed the experiments with me. Also thanks to my
lab-mates: Angela Tooker, Lyubov Kubatina, and Joey Mansour on helping the human
experiments and coil building. We made a great team! My special thanks to Ralph
Hashoian who taught me so much about coil technology and how to think like an
engineer.
I would also like to thank my MIT thesis supervisor, Dr. Frederick Bowman, who
was always there to help me with paper work, and who made sure that I turned in
everything on time.
My appreciation to my beautiful, and tall roommate, Hua-yin Yu, who helped to
polish up my thesis and gave me support through out my years at MIT. My cute, and
adorable friend, Moksha Ranasinghe, who stayed up with me to do work for countless
nights and offered me encouragement when I needed it. Finally, to my cool friend, Tina
Jan, who always gave me great and logical advises.
3
Contents
1 Introduction......................................................................................9
2 Theory and Background of HP-MRI.....................................................11
2.1 Magnetic Properties of Atomic......................................................11
2.2 MR Imaging Theory................................................................16
2.3 Hyperpolarized Noble Gas MRI..................................................18
2.3.1 Optical Pumping.........................................................19
2.3.2 Magnetization Generated by Hyperpolarization....................20
3 Setup of HP-MRI at the Very Low-field..................................................22
3.1 Superconducting Magnet.............................................................24
3.2 Gradient C oils..........................................................................26
3.2.1 Z Gradient Coils.........................................................26
3.2.2 X and Y Gradient Coils..................................................27
3.3 RF Transmit and Receive Coils..................................................28
3.4 RF Pre-Amplifier and Transmit and Receive (T/R) Switching System........32
3.5 Hyperpolarizing System for Noble Gases.........................................34
3.5.1 Xenon Gas Hyperpolarizing System...................................35
3.5.2 Helium Gas Hyperpolarizing System....................................36
3.6 Animal Gas Delivery System.........................................................37
4 Coil Development in the Very Low-field....................................................39
4.1 Basic Coil Circuit Components Calculation.......................................41
4.2 Method of Coil Building...........................................................44
5 Results and Discussion...........................................................................47
5.1 Coil Analysis in Very Low-field..................................................47
5.1.1 Coil Optimization with Number of Turns............................47
5.1.2 Optimizing Coil Shape for Lung Imaging..............................52
5.1.3 Conclusion on Optimizing Lung Coils..............................53
4
5.2 Phantom Images in Very Low-field................................................55
5.3 Very Low-field Lung Images of Rat in vivo.......................................56
5.4 Very Low-field Human Lung HP-MRI.............................................57
5.4.1 Background Noise..........................................................57
5.4.2 Human Lung Image........................................................62
5.5 Summ ary ................................................................................
63
5.6 Recommendations to Future Work...............................................63
Appendix.............................................................................................65
A ppendix A ...............................................................................
65
Appendix B ...............................................................................
66
References.........................................................................................67
5
List of Figures and Tables
Figure 2.1: Direction of the magnetic field generated by the charged particle depends on
the energy state level
Figure 2.2: Magnetic dipole moment generated by the unpaired protons.
Figure 2.3: Net magnetization can be generated when protons are placed in an external
magnetic field
Figure 2.4: In the absent of Bo-field, proton rotates along its axis, but when it is placed in
an external Bo-field, it starts to precess along the external field axis.
Figure 2.5: RF wave transmits into the body and exert a torque on the nuclear spin to
generate NMR signal, which later receives by the receive coil.
Figure 2.6: MRI signal generated in a homogeneous magnetic field, also called FID.
Figure 3.1: Overview of HP-MRI setup at Very Low-field
Figure 3.2: Cross section of a MRI superconducting magnet
Figure 3.3: Z Gradient Coil configuration
Figure 3.4: X Gradient Coil Configuration
Figure 3.5: Y Gradient Coil Configuration
Figure 3.6: The setup of transmit and receive coil which was used in all Very Low-Field
experiments
Figure 3.7: This figures shows the imaging area inside of the solenoid coil has
homogeneous magnetic field, but the field strength for surface coil is nonuniform
Figure 3.8: Overall diagram for the transmitting and receiving switch system
Figure 3.9: The quarter-wave circuit also called the pi circuit. The values of inductance
and capacitance depend on the Larmor frequency of the element.
Figure 3.10: Block diagram of the Hyperpolarizing System
6
Figure 3.11:
12 9
Xe Hyperpolarization chamber and trapping glassware used to purify
hyperpolarized xenon from helium and rubidium and to store the polarized gas
for long periods of time.
Figure 3.12: Animal Gas Delivery System
Figure 4.1: Helmholtz Coil
Figure 4.2: Circuit Diagram of the coil network
Figure 4.3: Simple Circuit model of coil
Figure 5.1: Human size proton coil inductance vs. number of turns
Figure 5.2: Normalized R and Q vs. surface coils
Figure 5.3: FID of Water Phantom using Human Size proton Surface Coils with different
turns
Figure 5.4: Possible lung coil designs in Very Low-field
Figure 5.5: FIDs of Helmholtz, planar, and surface coils
Figure 5.6: Spin echo water phantom image with Helmholtz Coil Dimension of the coil:
10.5"in diameter, 10" in height. Dimension of sample: 6" in diameter
Figure 5.7: Gradient echo helium phantom image. The bright spot in the center is from
the cell's Rubidium pull off finger. Dimension of Coil: 10.5" in diameter.
Dimension of Sample: 30 mm in diameter
Figure 5.8:
12 9Xe
lung image of rat in vivo
Figure 5.9: 3 He lung image of rat in vivo
Figure 5.10: (a) Basic noise data with noise peaks
(b) Hardware setup for the basic noise test
Figure 5.11: (a) Basic noise data after cross diodes reduced noise peaks
(b) Hardware setup with a pair of cross diodes
Figure 5.12: (a) Basic noise data without connecting all active electronics to the same
ground point.
(b) Basic noise data after connecting all active electroinics to the same
ground point.
Figure 5.13: Circuit diagram of low-pass filter. It is used to eliminate noise that was
generated by the gradient amplifier.
Figure 5.14: (a) Basic noise data before using gradient amplifier noise filter.
7
(b) Basic noise data after using gradient amplifier noise filter.
Figure 5.15: HP-MRI human lung image
Table 5.1: Comparison of surface coils with different turns
8
Chapter 1
Introduction
In today's medical field, Magnetic Resonance Imaging (MRI) has become the
primarily technique used to obtain high quality volume images of the inside of the human
body. MRI is based on the principle of Nuclear Magnetic Resonance (NMR), which is
the interaction between radio waves and atomic nuclei. The most commonly used atom
is hydrogen, which is ubiquitous throughout the human body, especially in the form of
water and fat.
However, water and fat are not very concentrated in some organs, which makes
imaging difficult. In order to overcome this problem, fast imaging methods and contrast
agents have been developed. Yet these techniques are unable to properly image the lung,
a gas-filled, water-free space. The idea of using other atomic nuclei as spin probes for
biological MRI studies has been proposed, but the concept is limited by the insufficient
concentrations that can be achieved in living organisms and the consequent high signalto-noise ratio (SNR). Noble gases such as xenon and helium also have the same problem,
but the nuclei of these gases can be hyperpolarized by optical pumping to achieve high
magnetization to overcome the SNR problem.
9
In obtaining lung images, patients are asked to inhale hyperpolarized (HP) helium
or xenon in order to enhance the NMR signal over one million times [4]. Since the NMR
signals are much higher than before, the static magnetic field can be decreased to reduce
the cost of the MRI scanner. The conventional 1.5 Tesla MRI scanner uses a high cost
super-conducting magnet. If HP-MRI is used, the magnetic field can be dropped from 1.5
T to 150 Gauss, which requires only a low cost permanent magnet for MR imaging. The
elimination of expensive superconducting magnets can lead to the development of vanbased imagers that could take this technology to patients whom are unable to travel to a
hospital. Another advantage of using permanent magnets is that MRI technology can be
used in the micro-gravity environments of spacecrafts and space stations to monitor the
physical states of astronauts [5].
The new Hyperpolarized Noble Gas Magnetic Resonance Imaging (HP-MRI)
technology, which utilizes a very low magnetic field, is currently being developed in the
laboratory of Mitchell Albert at the Brigham and Women's Hospital in Boston,
Massachusetts. In collaboration with Arvind Venkatesh, the project's goal is to build
transmitting and receiving coils for human lung HP-MR imaging in the 150 Gauss
magnetic field.
10
Chapter 2
Theory and Background of HP-MRI
This chapter reviews the magnetic properties of the atoms, basic magnetic
resonance imaging (MRI) theory, and the theories behind the hyperpolarized noble gas
magnetic resonance imaging process.
2.1 Magnetic Properties of Atomic Nuclei
Nuclear magnetic resonance (NMR), a technique that has been used for over 50
years to analyze chemicals, is the foundation on which MRI technology has been built.
The frequency of Electromagnetic waves and the energy received in MRI are much lower
than those of X-rays and of visible light. The energy of EM waves is directly
proportional to it frequency: E = h*v. Due to these properties of MRI, a type of
electromagnetic pulse called a radio frequency (RF) pulse is used to produce a signal.
NMR theory is based on Felix Bloch's discovery that spinning charged particles
create an electromagnetic field and that these particles have different energy levels. In
I1
MRI, hydrogen nuclei are the most common source of protons [6]. This is because
hydrogen is found in fat, which is in high concentrations in certain parts of the body, and
in water, which comprises approximately 70% of human body masses.
Following the Boltzmann statistics, charged particles have two energy states, -1/2
and +1/2, and spin about their axis to create a magnetic field (Figure 2.1). At room
temperature, the number of spins in the lower energy level, N+, slightly outnumbers the
number in the upper level, N-. This distribution can be represented by equation:
AE
-= e U
(Eqn. 2.1).
Where, E is the energy difference between the spin states, k is Boltzmann's constant,
1.3805x10-23 J/Kelvin, and T is the temperature in Kelvin. Because of the different
energy states, protons spin in opposite directions (north and south) from one another and
the resulting magnetic field is also generated in these two directions.
Direction of
magnetic field
Direction of spin
Direction of
magnetic field
Figure 2.1: Direction of the magnetic field generated by the
charged particle depends on the energy state level.
12
If there was an even number of protons, then every proton that spins pointing
north would be paired with another proton, which spins pointing south. The net magnetic
field created by this pair of protons is zero. If there is an odd number of protons, then
the unpaired proton causes a net magnetic field, known as a magnetic dipole moment
(MDM), which is in the same direction as the proton's spin. (Figure 2.2) The occurrence
of MDMs in some elements has lead scientists to utilize this property for imaging
purposes. A problem arises in the fact that the magnetic fields of many individual
protons cancel each other out, which results in a net field of zero. When an external
magnetic field (BO) is applied to the protons, they line up along the BO, with
approximately half pointing north and other half pointing south.
(No magnetic
field)
Paired protons
(Net magnetic
field)
Unpaired protons
Figure 2.2: Magnetic dipole moment generated by the unpaired protons.
13
However, about one in every million protons produces an extra northward spin,
which can accumulate to generate a net magnetization (M) in the direction of BO. (Figure
2.3) The MDM grows exponentially over time, where the time constant of the curve is
called the relaxation time, TI. M does not only depend on TI, but is also related to the
mobile proton density or the spin density, N(H). Therefore,
t
M =N(H)*(1-e
TI)
(Eqn 2.2).
:Net MF= 0
Bo off:
BoBon:
Figure 2.3: Net magnetization can be generated when
protons are placed in an external magnetic field.
When a proton spinning on its own axis is placed in a larger magnetic field, it
begins to wobble or move along the axis of BO as a result of gravity. (Figure 2.4) The
magnetic field exerts a force on the moving charge particle [7] can be described by
equation:
P = qv x BO
(Eqn 2.3),
14
Where v is the velocity of the charge, and P is the force. From the force equation, three
equations of motion can be derived,
m(
m(
x
dy
d X
)= qBO (-) -F(-)
r
dt
dt
dx_
dt
Y) =-qBO(x) -F(
2
dt
(Eqn 2.4)
(Eqn 2.5)
Y)
r
(Eqn 2.6).
m( d)= -F()
r
dt -
With the present of Bo, the equations of motion can be solved as:
2
2
M(d 2a =m( dd x )+im( d y2 )
m(2)
dt
dt2
dt-
F
da
-()a
dt
r
-iqBo (-)
(Eqn 2.7),
where a=x+iy, and i=V-1. Furthermore equation 2.6 can be derived into:
F
mk 2 +iqBok +-=0
r
(Eqn 2.8).
And the force exerted by BO is small compare to the central force, therefore,
-i( q )BO
(Eqn 2.9).
i
2m
rm
Thus
a = ecot (Q * el'st + R * e-1'wst)
(Eqn 2.10),
where Q and R are constants.
From this equation, the rate which proton precesses around the BO is given by the
Larmor Equation:
15
o=y*Bo
(Eqn 2.11).
where o is the angular precessional frequency of proton, normally expressed in units of
Hertz (Hz), and y is the gyromagnetic ratio, expressed in MHz/T. For hydrogen protons,
y(H)= 42.6 MHz/T.
Bo OFF
Bo ON
Figure 2.4: In the absent of Bo-field, proton rotates along
its axis, but when it is placed in an external Bo-field,
it starts to precess along the external field axis.
2.2 MR Imaging Theory
RF (radio frequency) waves and the resonant frequency of protons, also know as
the Larmor frequency, is delivered from a RF coil to the human body to excite the
hydrogen nuclei. (Figure 2.5) Coils are electrical devices generally composed of multiple
loops of wires that either generate magnetic fields or detect the changes in magnetic
fields. Typically from a stationary point of view, a proton spins around the magnetic
field axis at the Larmor frequency. However, if the spinning system is viewed from the
16
frequency equal or close to the Larmor
perspective of a coordinate system rotating at a
becomes stationary [8]. In such a
frequency, the perspective of the rotating proton
is easier to follow during and after the
coordinate system, the macroscopic magnetization
excitation of a RF wave.
RF
Tran1smit
Signal generated
Receive
exert
Figure 2.5: RF wave transmits into the body and
NMR
a torque on the nuclear spin to generate
signal, which later receives by the receive coil.
with Larmor frequency are
Commonly, high intensity and short-term RF waves,
body. RF pulses exert a torque on
pulsed into the coil to excite the protons in the human
away from the direction of the external
the magnetization (M), which causes the M to tip
to precess around BO with the Larmor
field, BO. However, the magnetization is forced
an oscillating magnetic field, part
frequency. Since this rotating magnetization represents
is picked up by a receiving coil also
of the energy that is associated with this oscillation
is received from a homogeneous
can be considered as an RF antenna. The signal that
induction decay (FID), which looks like
sample in a homogeneous field is called the free
17
a damped oscillation of a sine wave. (Figure 2.6) When a non-uniform external field is
applied to the sample, a MR image can be obtained.
I>L~
If
Received
signal
Time
Figure 2.6: MRI signal generated in a homogeneous
magnetic field, also called FID.
2.3 Hyperpolarized Noble Gas MRI
Conventional MRI is a water-proton-based technique that can effectively identify
abnormalities in most soft tissues. However, it is not as effective in the lungs and other
lipid bilayer membranes regions where proton based cells do not dominate the area.
Albert and his group proposed the first solution to this problem: instead of using
conventional MRI methods, hyper-polarized noble gases MRI (HP-MRI) can be used to
solve the dilemma.
18
2.3.1 Optical Pumping
HP-MRI requires that a patient inhale hyperpolarized xenon (129Xe) or helium
(3He) gas to fill his/her lungs. The image of the lung region is then created from the
signal that is obtained from the resonating nuclei of the isotope
12 9Xe
or 3He, rather than
that of protons. Normal thermal-equilibrium polarization of noble gases is only 0.04% of
the proton in the human body, which is insufficient for imaging. However, when a laser
optical pumping process is used to polarize the noble gases, the magnetization is
enhanced by a factor of 105 [1].
This is achieved when hyperpolarized I2 9Xe or 3 He
nuclei are produced by spin exchange with optically pumped rubidium (Rb), alkali-metal,
atoms.
Rb is an alkali metal which vaporizes at 85C. In an external magnetic field, Rb
splits into different electron spin states under the excitation of a titanium sapphire laser.
Depending on the direction of the circularly polarized laser beam, Rb moves into either a
plus or minus state with respect to ground [10]. Ground state electrons can be excited out
of these states by absorbing photons of 794.7nm wavelength. The polarization (P) of
alkali-metal with nuclear spin I=0 can be represented by:
P=NN1
-N
22
1
1 and N 1 +NiN_ =1
2'2
22
(Eqn 2.12).
2 2
The optical pumping exchange rate between two states can be described by equation:
dN(± })L2)_=) [ FsDy
2
dt
]N(-+) T FSDN(+-L)
(Eqn. 2.13).
19
Where y, is total rate per atom of pumping out of one state into another, and FsD is the
possible relaxation of electron spin polarization [12].
If polarized gaseous Rb is brought together with
12 9Xe
or 3 He, the collisions
between Rb and 129Xe or 3He atoms result in the transfer of angular momentum from Rb
valence electrons to the
129Xe
or 3He nuclei. The steady-state of the polarization after the
spin exchange can be expressed as:
P = N
N
N
(Eqn. 2.14)
N is the total number of spins. The spin exchange process increases the noble gas spin
population by twenty five percent over the equilibrium state, and therefore enhances the
NMR signal up to 105 times the thermal equilibrium value. If the polarized gas is kept in
a chilled magnetized charm, the polarization can last for hours, even days.
2.3.2 Magnetization Generated by Hyperpolarization
The magnetization of hyperpolarized gas can be derived from the general
magnetization equation:
M
-N y
(Eqn. 2.15).
4kT
Now, substitute the polarization equation into the equation 2.14 to form the
magnetization of hyperpolarized gas [13],
M Hyperpolarized -
y
2
(Eqn. 2.16)
20
The substitution of P is derived the Boltzmann equation,
N-
e(
k
1+
- AE
N+
(Eqn. 2.17).
kT
Let's manipulate the equation 2.16 by subtracting 1 from both sides, then the new
equality becomes:
N_-N
N+
-AE
kT
(Eqn. 2.18),
then dividing both sides by N and multiplying N+ to produce:
N_ -N
N
-AE N
kT N
(Eqn. 2.19),
Substitute h o> for AE, and
N for N+, by assuming the spins in one state comprises
approximately half of the total spin population,
N_-N,
N
--
2
kT N
(Eqn. 2.20)
Final, replacing the left side of equation by P, we obtain the equality of
P=
h
2kT
(Eqn. 2.21).
Equation 2.16 shows that magnetization of hyperpolarized gas is independent of the
external magnetic field strength, which means Very Low-field MR image is theoretically
possible. In the later chapter, this concept was also proven experimentally.
21
Chapter 3
Setup of HP-MRI at the Very Low-field
The schematic representation of the Very Low-field system with which
hyperpolarized 129Xe and
3He
images were acquired is shown in figure 3.1. The
prototype Very Low-field MRI imager was a 60 cm diameter superconducting magnet
with a magnetic field ramped from 1.5 T down to 15 mT (150 Gauss). Within the magnet
were Techron gradient coils for producing a gradient in the BO in the X, Y, and Z
directions with maximum strength of 0.8 Gauss/cm. Inside the magnet core, a shield
made from a 0.1 mm thick copper sheet was installed to reduce electrostatic noise. In
order to acquire MRI signals, a self-designed RF coil was placed inside the scanner to
transmit RF pulses and to receive NMR signals. Depending on the type of experiment,
different RF coils were used. The scan room was surrounded by an RF shield, which was
made from copper sheets. The shield prevented the surrounding RF pulses to enter the
scan room. It especially blocked various RF signals from television and radio stations
which can by detected by the imager.
22
Figure 3.1: Overview of HP-MRI setup at Very Low-field
The magnet, gradient coils, and RF coil were interfaced with a Resonance
Instruments Maran Ultra imaging spectrometer, which was operated from a Gateway
computer that controlled all components on the imager. The spectrometer signaled the
pulse software to send sine wave pulses at the
129
Xe and 3 He frequencies. The software
program caused the shaping of the RF pulses into apodized sinc pulses. After the RF
source sent out the pulse, the amplifier augmented the pulses' power from milli-Watts to
kilo-Watts before it reached the RF coil. The computer also controlled the gradient pulse
program, which set the shape and amplitude of each of the three gradient fields. A
23
gradient amplifier increased the power of these pulses to a level that was sufficient to
drive the gradient coils.
Commands to the MRI system were given to the computer through a control
console. Imaging sequences were selected and customized from the console according to
each experiment. The image and free induction decay (FID) signals were displayed on
the console, and further evaluations and analyses were done using the MATLAB
software program.
Besides the normal setup that is required for all MRI systems, HP-MRI also
requires a different gas hyperpolarizing system for the
12 9Xe
and 3 He experiments. For
animal experiments, Sprague Dawley white rats are used. A gas delivery system is set up
to help the rats inhale the hyperpolarized gas into their lungs. For human experiments,
gas is delivered in a Tedlar bag to the subject.
In the following sections of this chapter, the function of individual hardware,
hyperpolarizaing system, and the gas delivery system of the Very Low-field MRI will be
described in greater details.
3.1 Superconducting Magnet
The magnet is the component of the magnetic resonance imaging system which
created the static field for the imaging. Depicts the actual magnet used in these
experiments, most conventional clinical magnets are 1.5 Tesla superconducting magnets.
However, the one we used was ramped down to 150 Gauss. A superconducting magnet
24
is an electromagnet made of superconducting wire. This wire has a resistance
approximately equal to zero when it is immersed in liquid helium that is cooled to a
temperature close to absolute zero (-273.15 C or 0 K). Once a current induced in the coil,
it will continue to flow as long as the coil is kept at liquid helium temperatures. Some
losses might occur over time due to very small resistance that might be exhibited by the
coil. These losses are on the order of a ppm of the main magnetic field per year.
Figure 3.2 shows a cross sectional view of a superconducting magnet. The length
of superconducting wire in the magnet is typically several miles long. The wire coil is
kept at a temperature of 4.2 K by immersion in liquid helium in a large dewar. The
dewar is surrounded by a liquid nitrogen dewar at a temperature of 77.4 K which acts as a
thermal buffer between the exterior temperature (298K) and the liquid helium
temperature.
Vacuum
Liquid Helium
-E Liquid Nitrogen
MEE Container &Support
ME Superconducting Coil
Figure 3.2: Cross section of a MRI superconducting magnet
25
The helium and liquid nitrogen used to cool the superconducting magnet and to
keep it running are costly. However, with only a 150 Gauss magnetic field strength, the
expensive superconducting magnet can be replaced with the relatively cheap permanent
magnet. This is one of the greatest advantages of Very Low-field MRI.
3.2 Gradient Coils
The gradient coils produced the linear gradient change in the homogeneous static
magnetic field, BO, which decoded spatial information from the signal and localized it in a
given space. The variation in the gradient magnetic field is several orders of magnitude
smaller than that of the static magnetic field, but it is significant enough to allow spatial
encoding. Unlike imaging magnets, the gradient was kept at room temperature and the
configuration of the coils created the desired gradient in the X, Y, and Z directions.
3.2.1 Z Gradient Coils
Even though the imaging magnet was ramped down to 150 Gauss, the Very Lowfield system was still standardized like other MRI systems. A gradient in BO in the Z
direction is achieved with an Anti-Helmholtz type of coil (Figure 3.3). Current in the two
coils flow in opposite directions, creating a magnetic field gradient between the two coils.
The B field at one coil adds to the BO field while the B field at the center of the other coil
subtracts from the BO field.
26
Z Gradient Coil
Gz
B
B
X
zY
Figure 3.3: Z Gradient Coil configuration
3.2.2 X and Y Gradient Coils
The X and Y gradients in the Bo field are created by a pair of eight-shaped coils.
In the X axis, the direction of the current through the eight-shaped coils create a gradient
in BO in the X direction (Figure 3.4). The Y axis figure-eight coils generate a similar
gradient in Bo along the Y direction as the X-gradient coils (Figure3.5).
X Gradient Coil
Gx
x
Y
B
Figure 3.4 X Gradient Coil Configuration
27
Y Gradient Coil
X
Figure 3.5 Y Gradient Coil Configuration
3.3 RF Transmit and Receive Coils
RF coils used on humans to obtain MRI signals are also called transmit/receive
coils. The coils can be divided into three general categories: receive coils, transmit coils,
and both transmit and receive coils. A transmit coil is used to send RF pulses to interact
with protons inside of the human body. A receive coil is used as an antenna to pick up
the signal that is sent out from the body. For all the experiments contacted in the HypX
lab, the same coil acts as both transmitter and receiver to reduce the interference in the
magnetic field. (Figure 3.6) The RF source transmits pulsed waves to the RF coils, which
creates the magnetic field B, that is perpendicular to the static BO field. The coils also
detect the transverse magnetization as it precesses in the XY plane.
28
T/R
Switch
ftM~fl::
Transmission
~~~Line
Receiving
Line
Figure 3.6: The setup of transmit and receive coil which
was used in all Very Low-Field experiments
A transmit only coil is used to create the B, field and a separate receive only coil
is used in conjunction with it to detect or receive the signal from the spins in the imaged
object. However, the separate transmit and receive coils have to be placed inside of the
BO field together, which means the two coils will interfere with each other and alter the
behavior of the B, therefore distorting the signal from nuclear spin. This is the main
reasons why a single coil is used in experiment to transmit and receive RF waves. The
ones used in the experiments described are the transmit and receive coils. These types of
coils serve as the transmitter of the B1 field and receiver of the RF energy from the
imaged object.
All RF coils must resonate, and efficiently store energy at the Larmor frequency
in order to obtain the small NMR signal that is produced by the spin in the atoms.
Therefore the coils are resonated with capacitor circuits. Conventional MRI also uses
inductors in conjunction with the capacitors in the circuit. However, in the 150 Gauss
field, the Larmor frequencies of proton, helium, and xenon are all below 1 MHz which is
much lower than the Larmor frequencies at 1.5 Tesla.
As we know the resonant
frequency, o, is related to inductors (L) and capacitor (C) circuit by
29
LC
(Eqn 3.1).
For RF coil building, C is the total capacitance in the circuit. In conventional MRI, L is
the total inductance for the coil and the inductor, whereas in the Very Low-field, L only
represents the coil inductance. Inductors are not recommended for use in the circuit for
150 G magnet field. This is because L is inversely proportional to o, which is very small,
and therefore for the same capacitor values, L is relatively large for the Very Low-field.
Adding more inductors to the circuit would only increase the inductance and the
resistance in the coil, which leads to a decrease in the coil's energy storing efficiency.
Further descriptions of coil building and development will be described in chapter four.
One of the requirements of an RF coil is that the produced B, field must be
perpendicular to the Bo magnetic field. Therefore, solenoid coils were used for the animal
experiments, and surface coils were used for the human experiments. The magnetic field
created by the solenoid coil is given by the equation,
B =#JOiN
(Eqn 3.2).
The magnetic field created by the surface coil is given by the equation,
B
-
0 i 0N
21r
(Eqn 3.3).
where go is the permeablity constant, i is the current flowing through the wire, N is the
number of loops of wire that constitute the coil, and r is the distance from the coil. By
comparing these two equations, it is clear that the magnetic field in the solenoid is
homogenous, and that the field in a surface coil decreases with the distance, r (Figure
30
must be placed
3.7). In lung imaging, solenoid coils are preferred, but the solenoid
field. Therefore a
horizontally in the external magnet in order to create the perpendicular
the external magnet
human cannot fit in the coil, because a human cannot fit in
horizontally.
B0
B0
A
T/R
Switch
Solenoid
Coil
Figure 3.7: This figures shows the imaging area inside
of the solenoid coil has homogeneous magnetic field,
but the field strength for surface coil is non-uniform.
useful only to certain
Each coil generates its own magnetic field pattern which is
in the Very Low-field lungs
part of the body. Only solenoid and surface coils are used
of coils, such as birdcage
imaging, but convention field generation utilizes other types
coil of choice for imaging the
coils and single turn solenoid coils. The birdcage coil is the
field around head. The
head and the brain, because the coil can create a homogeneous
turn solenoid is similar to the
shape of the magnetic field that is generated by the single
for the detection of
shape of the breasts, therefore this type of coil is often used
abnormality in the breasts.
31
The 150 Gauss coil building technique is based on the equations and formulas that
are used in 1.5T, but alterations have been made. Development of coils in low field will
be discussed in detail in Chapter Four.
3.4 RF Pre-Amplifier and Transmit and Receive (TR)
Switching System
The Resonance Broadband RF Pre-amplifier and TR switch system generate
frequencies from 0.1MHz to 30MHz. Pulse sequence controls regulate the amplitude and
duration of the pulse (Figure 3.8). The pulse is then transmitted into a circuit that
switches between the transmit and receive states.
Receiver
COIL
ProbeL
C
C2
O
Transmitter
G
Figure 3.8: Overall diagram for the transmitting and receiving switch system
32
The RF wave is sent down to the transmission port in the pre-amp system and
then through a pair of cross diodes to the RF coil. The cross diodes are used to reduce the
noise from the transmission line when the MRI spin signal is received from the coil. A
quarter wave circuit set at the desired noble gas frequency is connected to the coil to
serve as an TR switch (Figure 3.9). If one end of the quarter wave (1/4) circuits is open or
closed, then the other end of the circuit is being inverted to closed or open. At a 50Q
load on one end, the circuit will also see a 50K at the other end.
L
Output
Input
Figure 3.9: The quarter-wave circuit also called the pi circuit.
The values of inductance and capacitance depend on
the Larmor frequency of the element.
When the RF power is transmitted into the coil, the circuit path is closed at one
end of the quarter wave, and the circuit connected to the pre-amplifier at the other end is
open. When the circuit is in the receive mode, the quarter wave sees 50K2 at the coil end.
This results in the transmission of the NMR signal through the k/4 circuit to the first
stage MITEQ pre-amplifier. The NMR signal is then filtered through a band-pass filter at
33
the desired noble gas nuclear frequency before it is sent to the pre-amplifier (Appendix
A).
The band-pass filter serves an important role in the circuit by eliminating noise
that is added to the system by other sources, and allowing only the desired frequency to
pass through. In the circuit, a passive filter is preferred over an active filter. A passive
filter contains only passive elements such as capacitors and inductors. However an active
filter needs to be activated with a power source. In general, the noise level of the power
source is as high as the acquired NMR signal, making the active filter undesirable for use
in the circuit.
3.5 The Hyperpolarizing system for Noble Gases
The 3 He and
129Xe
gas flow-through system, also known as the hyperpolarizing
system, was designed with a 150 Gauss portable HP-MRI in mind for the future. The
entire system was built on a cart with pneumatic casters for accessibility [14]. The
hyperpolarizing system takes mixed gas from a compressed gas tank and delivers it into a
purifying manifold via " stainless steel tubing. After purification, the elutriated gas
flows into the hyperpolarization chamber to be polarized by a laser and is then collected
in a glass cell for final use. When the experiment is ready to proceed, the collected
polarized and purified gas is transferred in a Tedlar bag for human inhalation, or to a gas
delivery system for animal use (Figure 3.10).
34
Gas
m ure
_Manifold!
ufcon
mixtuepuifictionchamber
Glassware/
hyperpolarizing
Hyperpolarized
gas
Figure 3.10: Block diagram of the Hyperpolarizing System
3.5.1 Xenon Gas Hyperpolarizing System
In the case of the 129Xe gas flow-through system, the gas mixture consists of
1% xenon, 98% helium and 1% other impurities. The helium in the gas mixture acts as a
catalyst to increase the collision rate between rubidium and xenon during optical
pumping. The gas mixture flows into the purifying manifold and then enters the
hyperpolarizing chamber, where gas is stored in a glass cell that is coated with rubidium
alkali-metal. When the cell is heated to 85C in a glass oven, the rubidium vaporizes in
the cell and mixes with the
129Xe
gas. Then a laser is shined on the cell to induce the two
gases to shift in their energy states and begin the optical pumping process (Appendix B).
After the gas mixture is polarized, it flows into capture/transfer glass cells to separate the
pure hyperpolarized 129Xe gas. The first transfer glass cell uses cold water to condense
and capture all of the rubidium and helium atoms, and only allows hyperpolarized
35
129 Xe
to move to the second transfer glass cell. The pure hyperpolarized ' 29Xe is condensed and
collected by using liquid nitrogen to freeze the cell. (Figure 3.11) This method is able to
keep the polarization of the gas for days. When the gas is needed, the liquid nitrogen is
removed and the polarized gas starts to flow into the Tedlar bag or the animal delivery
system.
Hyperpolarization
Chamber
To Tedlar
Oven
First
Transfer/Capture
Glass-ware
Second
Transfer/Capture
Glass-ware
Figure 3.11: 12 9Xe Hyperpolarization chamber and trapping
glassware used to purify hyperpolarized xenon from helium and
rubidium and to store the polarized gas for long periods of time.
3.5.2 Helium Hyperpolarizing System
The method used to polarize 3He follows four steps similar to that used to polarize
129 Xe
(figure 3.13) except with slight differences. The helium mixture contains 99%
helium, and 1 % impurities. The purifying manifold eliminates most impurities that are
introduced during the manufacturing process. Unlike the
3He
129Xe
polarization process, the
system does not require the transfer/capture glass-cell. The polarized gas does not
need to be frozen or separated from other molecules. Instead, the only glass cell that is
36
needed is the inside of the oven where gas can be stored and polarized by the laser. The
optical pumping process is the same as the ones described for the I29Xe system.
(Appendix B) The
3He
must be pumped until the experiment is ready to proceed, when it
can be transferred into a Tedlar bag.
3.6 Animal Gas Delivery System
The animal gas delivery system simulates the gas breathing process for the
Dawley rats. Computer controls the ventilation, gas injection and triggers scanner
through LabVIEW software. The SAR-830 series ventilator regulates the gas that go in
and out of the animal by connected to a series of tubes and valves that adjoined the
hyperpolarizing system and the animal [9]. (Figure 3.12)
SAR-830 serdes Ventilator
PC with TimIng Board & Labvlew
PC conls
%VnNadon,Gas
Nnjbon
0-d
TRACHEAL
PRECISION
Frigae
prue
s
N
plu "
RESTRICTOR
TB
PNEUMAIC
UAT
Figure 3.12: Animal Gas Delivery System
37
The gas is stored in a Tedlar bag contained in a rigid acrylic box. By pressuring
the box, the gas flows out of the bag and enters the breathing circuit. The gas injection to
the rat is synchronized to the computer controlled animal ventilation which connected to
animal's airway through a set of non-metallic valves. The animal ventilator gates a
controlled gas flow into the animal during the inspiration phase of the respiration. The
airflow is set by a rotameter-type regulator, and the timing is controlled by the LabVIEW
software program. Expiration is passive and occurs when the animal's airway is opened
to the atmosphere via the solenoid valves.
38
Chapter 4
Coil Development in the Very Low-field
From the MRI theory, the Larmor frequency is proportional to the external
magnetic field strength. In the 150 Gauss Very Low-field, the Larmor frequencies for the
different elements' nuclear spins are 634.87 KHz for protons, 483.64 KHz for 3He, and
175.60 KHz for
129Xe.
In order to achieve resonant frequencies lower than 1 MHz and to
generate a large signal to noise ratio (SNR), the RF coils need to have a high quality
factor,
Q.
QoL
The Q factor for the coil is describe by the relationship:
(Eqn 4.1)
R
where R is the loss in the coil, coo is the Larmor frequency, and L is the inductance of the
coil. In order to have a large
Q,
wo is given for 150 Gauss, and for a small R, L has to be
large. The only way to achieve large inductance, multiple loops of litz wires are used to
build the RF coil. Litz wire is proven to be better than copper wire by Y.J. Yang at Korea
University. In his study, it showed that Litz wire gives approximately 2.6dB (30%) better
39
SNR for frequencies around 200 KHz, but at 1 MHz, the difference between two wires is
only 0.5dB (5%)[15].
After deciding what type of wire to use for building the RF coil, the next major is
decision lies in the design of that coil that should be used for lung imaging. Surface coils
are commonly used in MRI to obtain excellent high-resolution images. The surface coil
design, which is based on the reception field, is only coupled to the region of interest and
served to decrease the coupling to the environment which reduces the noise level.
Surface coil is ideal for body surface imaging, such as the spine, but the lungs are set
deep in the chest cavity and therefore difficult to image with the surface coil. To solve
this problem, a pair of surface coils known as Helmholtz coils are used to create a
homogeneous magnetic field in the medial plane of the body where the lungs are located
(Figure 4.1).
Transmit
-~-
----
Solenoid
Coil
Receive
Figure 4.1: Helmholtz Coil
40
The first step in building a Helmholtz coil is to construct two surface coils which
are matched and tuned to resonate at the desired Larmor frequency. When a coil is
matched, it means the input impedance is 50Q, so no power is reflected back to the coil.
After completing the task of tuning and matching the two individual surface coils, the
final Helmholtz coil can be build upon the two surface coils. The following sections will
discuss the details of coil building and make comparisons between different coils.
4.1 Basic Coil Circuit Components Calculation
To build an effective coil, it is important to choose the correct value for the
capacitors in the coil circuit. In this section, circuit analyses will be used to calculate the
values for the capacitors [16].
L02
C
Figure 4.2:Circuit Diagram of the coil network
41
Figure 4.2 represents the general structure of surface coil circuit. L is the coil
inductance, C, is the lumped capacitance to establish the input port, C2 is the total
distributed capacitance, R is total coil loss, and Zi, is the input impedance of the circuit.
From circuit analysis, Zi,, can be represented by:
1
R+
Zi
=
II_(1--
2
LC 2 )
(Eqn 4.2).
COC 2
1+
-
-oLC,
+ jwRC,
C2
At the desired resonance frequency, wo,
00 =
1
__
4.3),
IC(Eqn
LCC2
C1 + C2
where the reactance of L and the sum of C1 + C2 are equal. Therefore the input impedance
can be reduced to:
Zn(cw
=
-
1
(wOC
1)
2
R
+
1
jo 0 C,
(Eqn 4.4).
A simple circuit of a series of capacitors and resistors (Figure 4.3) can model this
equation.
42
zin
Figure 4.3: Simple Circuit model of coil
After finding the capacitance, the tuned coil needs to be matched to 50Q, RO.
When a coil is tuned to the Larmor frequency, the reactance of the circuit tends to be
more capacitive, and the resistance of the coil depends on the values of the capacitors.
The next step is to replace Zi, with the desired 50E2 impedance, and to solve equation 4.4
for the value of C, and to solve equation 4.2 for the C2 value.
R
CI
=
OJ~ L
(Eqn 4.5),
and
1-R
(Eqn 4.6).
C2 =
0z L
43
4.2 Method of Coil building
Before the values of capacitors can be calculated, there is a systematic way of
finding the resistance of the coil and then using the theoretical capacitor values to tune
and match the coil. However, in practice, the theoretical value is only close to the actual
value, and therefore the procedure of coil building is written below:
Step One: Measure the inductance of the coil by connecting it across the reflection port
of the network analyzer. Set the analyzer to the sl measurement and read the
inductance value from the Smith chart at the resonate frequency of interest.
Step Two: After obtaining the inductance value, use the relationship c>=2rf=1/V(L*C)to
find the resonant capacitance value wheref is the Larmor frequency, L is the
inductance, and C is the capacitance. Place the capacitor in parallel with the coil to
obtain the desired resonant frequency. Use a s21 measurement on the network
analyzer to tune the coil to the Larmor frequency. If the resonant frequency does not
match withf, then adjust the capacitor value untilf is obtained.
Step Three: Obtain quality (Q) factor experimentally, and using the equation Q=(w)*L)/R
to calculate R, the resistance of coil. In order to measure the
Q factor atf, first use the
s21 measurement to capture the curve which represents the coil behavior. On the
analyzer display,
Q represents the 3dB bandwidth
of the coil in the s21 measurement.
Setf as the reference point by fixing the marker to be zero at that point. Then, use the
bandwidth function to find the 3dB points on the curve. The
the right corner of the screen. Then use the
44
Q value is
Q value to calculate
R.
displayed on
Step Four: Use the equations, Cparaiei= V(R/50)/(dw*L) and Cseries
=
(I-V(R/50))/(d*L),
which were derived in section 4.1, to calculate theoretical values for capacitors in
parallel and series to the coil. Place those capacitors in the circuit and obtain the
resonant frequency from the sIl measurement.
Step Five: If the measured frequency does not equal tof, then adjust the capacitor value
by adding or removing capacitors, and by adding adjustable capacitors into the circuit
until f is observed. Use the s 11 measurement and the smith chart to perform the final
tuning and matching of the circuit.
Step Six: After the RF coil is tuned and matched on the work bench, the final
adjustments can be made in the simulator (the big shield) if the coil is built for an
animal experiment. Otherwise, for the human experiment, the final adjustments have
to be made inside of the MRI scanner. The simulator emulates the interior
environment of MRI scanner, which couple to the coil and shield all other RF sources
in that room that would have affected the tuning and matching processes.
Step Seven: For a human sized coil, an extra set of capacitors is added between the coil
and the ground to stop the oscillation in the coil.
This procedure should be followed to build the first two individual surface coils,
which will be used to build the Helmholtz coil. Wind both coils in the same direction,
and make two leads at the end. Place the coils on top of each other, then solder the
bottom lead of the top coil with the top lead of the bottom coil. After forming this
connection, place a stick between the two coils to simulate the gap that will be created by
the human chest area. During the experiment, one of the coils will be placed on top of
45
the chest, and the other on the back of the chest. The lungs are positioned in the middle
of the Helmholtz coil to experience the homogeneous field that is created by the coil.
After setting up the Helmholtz coil on the workbench, follow the procedure again
to tune and match for the Helmholtz coil. However, when a human is placed between the
coils, the frequency and the impedance can change dramatically. It is necessary to
actually place the person between the coils and put him/her inside of the magnet to readjust the capacitors' value until the Larmor frequency is achieved.
46
Chapter 5
Results and Discussion
5.1 Coil Analysis in the Very Low-field
In the study of Very Low-field HP-MRI spectroscopy, coil design plays an
important role in obtaining a high signal to noise ratio during the imaging process.
Analyses of different types of surface coils and different number of turns that optimize
the coil efficiency were conducted to help determine which coil design should be use for
the human lung imaging experiment.
5.1.1 Surface Coil Optimization with Number of Turns
In order to understand how the number of loops of wire is affecting the properties
of the coil, 1-turn, 5-turn, 10-turn, 20 turn and 30-turn human size (10.5inches in
diameter) surface coils were built. To keep the consistency of the experiment and the
convenience of testing the SNR in the scanner, these coils are tuned and matched to the
47
proton frequency, 634.87 KHz, instead of to the frequency of 3He or I 29Xe. The proton
sample does not need to be hyperpolarized, therefore the magnetization is constant for
every experiment as long as the same sample is used. If 3He or 129Xe is used, the
polarization depends on the quality of the optical pumping, which means that the
magnetization varies from experiment to experiment. Therefore an additional unknown,
other than the number of loops used for the surface coil, affects the SNR. In the table 5.1,
inductance, quality factor, calculated capacitance, measured capacitance, and resistance
of the surface coil network at proton frequency are given.
Coil Number
Number of
1 turn
2
5 turns
3
10 turns
4
20 turns
5
30 turns
1.35pH
19.43 gH
80 pH
266 pH
570 pH
1
Turns
Inductance
(L)
_
Quality
6
10
31
205
308
40.3 nF
1.25 nF
429 pF
160 pF
114 pF
6.2 nF
2.0 nF
356 pF
76 pF
52 pF
50 nF
2.953 nF
710 pF
200 pF
57 pF
11.5 nF
0.76 nF
820 pF
100 pF
39 pF
N/A
N/A
802 pF
180 pF
N/A
0.90 Q
7.75 Q
10.29 Q
5.176 Q
7.38 £2
Factor (Q)
Calculated
Cparallel
Calculated
Cseries
Actual
Cparallel
Actual
Cseries1
Actual
Cseries2
Resistance of
Coil
Network
Table 5.1: Comparison of surface coils with different turns
Figure 5.1 shows the values of inductance verses the number of turns for all
surface coils. The relationship between inductance and the number of turns follows the
physics equation:
48
i
(Eqn. 5.1),
where L is the inductance, N is the number of turns, 0P is the flux, and i is the current.
Inductance is positively proportional to the number of turns in a coil. An increase L will
lead to an increase in inductance.
Figure 5.1 Human Size Proton Coil Inductance vs.
Number of Turns
600500-
£
0
400-
C.)
E
a)C.)
Cu
300200-
C.)
V
C
1000-
-
0
10
20
30
40
Number of Turns in Coil
It is necessary to determine which number of turns creates the most affective coil
from the parameters that we measured and calculated. An efficient coil should have low
resistance (R) and a large inductance value to produce a large Q value. In figure 5.2, the
Q and R values are normalized to 1 individually, with 1 being the most desired value. To
normalize Q, the larger the quality factor, the closer it is to 1. In normalizing R, the
smaller the resistance, the closer it is to 1. In the graph, two normalized values are
stacked on top of each other for evaluation. Coil #5 (30 turns) has the highest normalized
49
value, whereas coil #4 (20-turns) is more evenly distributed between the normalized
Q
and R. It is apparent that coil #5 and coil #4 are better than coils #1, #2, and #3.
Figure 5.2: Normalized R and Q vs. Surface Coils
1.4 1.2 -
0
cc
G)
N
E
0
z
1 -
0.8
--
-
MNormalized
R
Normalized
Q
0.60.40.202
4
3
5
Coil Number
To test the efficiency of the surface coils, a water phantom was used to obtain
FID, from which SNR can be calculated. In figure 5.3 we can see that the 1-turn surface
coil gives a relatively small signal, which has been distorted by background. The FIDs of
5-turn and 10-turn coils look much clearer than the FID obtained by thel-turn coil, but
still an enormous amount of background noise is riding on the signals. The FIDs
obtained with the 20-turn and 30-turn coils are sufficient for MR imaging. However, the
20-turn coil has the highest SNR and lowest noise riding on the FID.
50
Figure 5.3 FID of Water Phantom using Human Size
proton Surface Coils with different turns
IJ
-
N
E--
--
5-turns, SNR: 22, File: 991030-18
1-turn, SNR: 7.50, File: 991030-24
-M
-
-E
-l~
-
20-turns, SNR: 129, File: 000320-01
10-turns, SNR: 21.7, File: 991030-13
I
-
-ii
-s
30-turns, SNR: 41.9, File: 000320-04
51
H
5.1.2 Optimizing Coil Design for Lung Imaging
After determining the optimal number of turns to be used for a coil, the next step
was to obtain the best shape for lung imaging. In the previous section, it was mentioned
that coil design is unique to different part of the human body. In the case of lung
imaging, the possible coil designs are surface, planar, and Helmholtz coils (Figure 5.4).
The surface coil's wires are wound on top of each other and it looks like a short solenoid
with a large diameter. The planar coil has its wires wound spirally on a flat surface. The
Helmholtz coil is two surface coils on top of each other.
Surface
coil
Planar
coil
Helmholtz
coil
Figure 5.4: Possible lung coil designs in Very Low-field
The qualities of these three coils were tested with a water phantom. FIDs were
obtained to compare the SNR difference between the coils (Figure 5.5).
All FIDs were
the results of 90-degree flip angle pulses with the shim turned on to fine tune the
homogeneity of the magnetic field. The 90-degree flip angle refers to the RF power that
is required to preccess the charged particles that spin along the Z-axis to X-axis plane.
52
By comparing the measured FIDs and calculated SNRs, it was found that the Helmholtz
coil obtained a sufficiently high signal that noise could not be detected beyond of the
FID. Besides the high signal to noise ratio, the FID obtained by Helmholtz coil decayed
much more slowly and uniformly than the ones obtained by the planar and surface coils.
The slow and uniform decay of FID indicated that it had a long relaxation time and hence
a long lasting signal which can give a clear image.
5.1.3 Conclusion on Optimizing Lung Coils
The experimental and theoretical results show that a 20-turn Helmholtz design is
the most desired coil to be used for lung imaging in the Very Low-field.
The theoretical calculation has shown that it is better to have coils which can
create homogeneous magnetic fields in the region of imaging. Helmholtz coil satisfies
this criterion conceptually and experimentally. Through the FID tests that I conducted,
the numbers imply that the Helmholtz coil is the choice for lung imaging in the Very
Low-field.
Apart from the shape of the coil, the number of turns of the coil was proven to
play an important role in SNR. The inductance (L) of the coil increases with the number
of turns. The total loss of the coil network (r) is inversely proportional to the quality
factor (Q). A coil with large Q, large L, and small r will have a high SNR. In our case,
the bench test showed that a 30-turn coil has a large Q, but its total network loss is greater
than that of a 20-turn coil. However, the FIDs that were obtained using these two coils
showed that a 20-turn coil has a SNR of 129, and a 30-turn coil a SNR of 41.9. By
comparing FID results the 20-turn coil was chosen for the Helmholtz design.
53
surface coils
Figure 5.5 FIDs of Helmholtz, planar, and
Helmholtz coil, SNR: 363, File 000421-02
i
HMM
EMEoE
Eim
Planar coil, SNR: 87, File 000312-01
Surface coil, SNR: 134, File 000227-88
54
5.2 Phantom Images in Very Low-field
Proton FID was used as a reference for Very Low-field Lung imaging. Before the
experiments with the hyperpolarized gas MR lung imaging was carried out, a water
phantom was imaged with the Helmholtz coil to prove the proper functioning of the coil.
(Figure 5.6) After successfully imaged the water phantom, a helium phantom was
imaged with the same coil (Figure5.7). These tests concluded that both the coil and
polarization of helium gas are ready for human lung imaging.
Figure 5.6: Spin echo water phantom image with Helmholtz Coil,
Dimension of the coil: 10.5"in diameter, 10" in height
Dimension of sample: 6" in diameter
Figure 5.7: Gradient echo helium phantom image. The bright
spot in the center is from the cell's Rubidium pull off finger.
Dimension of Coil: 10.5" in diameter
Dimension of Sample: 30 mm in diameter
55
5.3 Very Low-field Lung Images of Rat in vivo
While the human sized lung coil was being designed and constructed, the HP-MR
lung imaging was carried out on a rat in vivo with 12 9Xe and 3He gas. After obtaining the
very first image used
129Xe
gas, He gas was also used for imaging. (Figure 5.8, 5.9)
Figure 5.8:
12 9Xe
lung image of rat in vivo
Figure 5.9: 3He lung image of rat in vivo
56
By comparing the two images in figure 5.7 and 5.8, it can be seen that the image
obtained by using
3He
gas has a higher signal level than the one obtained by 12 9Xe gas.
This phenomenon can be due to three main factors. First, the hyperpolarization process
of 3He is much better understood then 129 Xe in our lab, even though both processes have
not achieved the maximum polarization. Another factor is that the gyromagnetic ratio for
129 Xe is 2.8 times smaller than 3He, therefore even with maximum polarization, 129Xe
polarization would still be lower than 3He by a factor of 2-4. Finally, the abundance of
129Xe
isotope is only 26% of xenon at natural abundance, where 3He is at 100% isotopic
enrichment.
By observing the comparison between
129Xe
and 3 He produced images, 3He was
chosen to be used for human lung HP-MRI for the initial run.
5.4 Very Low-field Human Lung HP-MRI
5.4.1 Background Noise
In conventional MRI, the two significant sources of noise are the biological
sample and the coil network. Coil noise is proportional to o>P up to 1MHz, since we are
operating in the Very Low-field regime, this effect can be ignored. The sample noise is
proportional to o, therefore it would have more of an effect in high-field than the lowfield situations [2]. Therefore, sample noise dominates in the high-field, and coil network
noise dominates in the low-field.
The major difficulty that we encountered during HP human lung imaging that did
not occur during animal experiments was the background noise picked up by the RF coil.
The background noise increased with the size of the coil. In comparison, the size of the
57
coils used for humans is much larger than the one used for animals. Coil shielding,
which was used in the animal experiments to reduce noise, was not used in the human
experiment due to the coupling effect.
The strategy we used to reduce background noise was to eliminate paths which
might introduce noise into the HP-MRI hardware system. Figure 5.10 (a) shows a basic
noise level test with large noise peaks. Often, the signal level of noise induced from the
TX line is much larger than the NMR signal obtained from the lungs. The hardware
connection for the test was the RF source connected to the RF port, 50Q connected to the
coil port, and receive port connected to the pre-amplifier (Figure 5.10 (b)). In order to
eliminate the noise peaks, a pair of cross diodes was interfaced between the transmission
(TX) line and the TR switch (Figure 5.11 (b)). This component blocks all noise from the
transmission line during the receiving mode. Figure 5.11(a) shows the cross diodes have
reduced the noise peak from the background noise.
Figure 5.10 (a): Basic noise data with noise peaks
58
Pre-Amp
Receiver
I
RF Source
T/R Switch
Coil port
-
500 Resistor
Figure 5.10 (b): Hardware setup for the basic noise level test
V
Figure 5.11 (a): Basic noise data after cross diodes reduced noise peaks
Pre-Amp
Receiver
R
Coil port
RF Source
T/R Switch
I
0
Resistor
Figure 5.11 (b): The hardware setup with a pair of cross diodes
59
Other than adding the cross diodes, all active components such as gradient
amplifier, MITEQ amplifier and the Resonance Pre-amplifier were grounded together to
the superconducting magnet. Connecting all grounds together generated a common
reference point for the system which reduced noise entering the system. Comparing part
(a) and (b) of Figure5.12, the noise peaks in the background noise were reduced by
connecting all electronics to the same ground.
1.1'
1
..
.. 1
111.1111
1 III.I j
,11. IJ I.1.1
fill 11.11
1
11
'fl
(a)
(b)
Figure 5.12: (a) Basic noise data without connecting all
active electronics to the same ground point. (b) Basic noise
data after connecting all active electroinics to the same ground point.
60
Finally, with Dileep's help, a passive low-pass noise filter was built for the
gradient amplifiers to reduce noise that was generated by the power source. (Figure 5.13)
The noise source produced by the gradient amplifier did not affect the SNR for the
animal coil, but distorted signal level for the human coil. (Figure 5.14 (a)) Background
noise was measured again after altering the low-pass filter. Figure 5.14 (b) shows the
elimination of the noise peaks.
Gradient
Amplifier
L
CGradient
Coil
Figure 5.13: Circuit diagram of low-pass filter.
It is used to eliminate noise that was generated by the gradient amplifier.
(b)
(a)
Figure 5.14: (a) Basic noise data before using gradient amplifier noise filter.
(b) Basic noise data after using gradient amplifier noise filter.
61
5.4.2 Human lung image
In HP-MRI, since the magnetization of the atom is set by its polarization, which is
2
independent of the external field, the signal scales with BO instead of B0 at low field and
scales with BO at high field. The ratio between SNR(HP)/SNR(water) scales with 1/B,
therefore HP-MRI has significant advantage over conventional MRI at Very Low-field.
Since the noise factor of HP-MRI only depends on the coil network, this problem
can be reduced by using low loss superconducting RF coils with unloaded Q factor
greater than 50,000 [3]. However, this technique does not help conventional MRI,
because the sample noise dominates [18].
Figure 5.15: HP-MRI human lung image
From the theoretical point of view, HP-MRI is a certain possibility in the Very
Low-Field, but no one has proved it experimentally. After testing the coil and the
hyperpolarizing system, and reducing background noise, the very first HP-MRI human
62
lung image was taken with Mitch Albert as the subject. (Figure 5.15) In the image, the
lobes of the lung can be seen clearly, however, the 10" in diameter coil is not big enough
to cover the apex and the base of the lung.
The successful imaging in the Very Low-field shows its possible clinical use for
people with lung disease in the future.
5.5 Summary
A 20-turn Helmholtz coil was built for obtaining human lung images. To design
the coil, the number of turns was optimized to 20-turns. Among surface, planar, and
Helmholtz coil shapes, Helmholtz was chosen to be used for its high SNR. Besides
having an efficient human lung coil, noise sources were eliminated from the HP-MRI
hardware system. A set of cross diodes were placed between the T/R switch and the
transmitting line to reduce the noise entering the system during the receiving mode. All
active electronics were connected to the same ground reference point to reduce the
oscillation between each connection. Final alteration of the system was done by adding a
low-pass noise filter between the gradient amplifier and the gradient coil. The filter
reduced the noise in the same range as the Helium Larmor frequency.
5.6 Recommendations to Future Work
Further work needs to be carried out to improve the Very Low-field HP-MRI
system to obtain higher quality images. The 10" diameter Helmholtz coil is not large
enough to image the full length of the lungs. Figure 5.14 depicts a human lung image,
63
where it is can be seen that the base of the lungs is missing. The use of a larger Helmholtz
coil may allow more complete lung imaging. A circular shaped coil may not be the best
design, therefore it is necessary to investigate the utility of other shapes, such as that of a
rectangular coil.
The experiments described used a 30-turn coil, which may not be optimal.
Further studies should be conducted to determine the ideal number of turns. In order to
understand how inductance, quality factors, and loss resistance in the coil network affect
the SNR, 40-turn, 60-turn and 100-turn coils should be studied.
When a new Very Low-field HP-MRI system is assembled, all active electronics
should be placed close to the scanning magnet. A shielding system worked effectively in
the animal experiments and should be considered for the human studies. Coil shielding
was attempted for human experiments, but it coupled to the coil greatly so that the SNR
was reduced. However, increasing the distance between the coil and the shielding might
cause an augmentation in the SNR.
64
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2~- *4
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i
VEER a
e
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Appendix B: Optical Pumping
The figure shows the optical pumping setup and the procedure is described below.
Heater
Laser
Temperature
Control
Box
Gas Cell
Ga Cell
Glass
Oven
l.Setup the 3He or 129Xe cell and place it in the glass oven.
2.Aligned the laser beam and the glass oven with the external magnetic field.
3.Turn on the Omega Temperature Control Box and make sure the heater sensor is taped
to the cell. Set the temperature to 150C for 3He, and 1 10C for 12 9Xe.
4.Turn on the air compressor then the Variac heater.
(WEAR EYE PROTECTION)
5.When the cell temperature reaches 80C, turn on four channels of laser beams on.
3He
CELL: WAIT FOR 2 HOURS 129Xe CELL: WAIT FOR 1.5 HOURS
6.After the waiting period, turn off the Variac heater. When the temperature stops
dropping, then turn off two channels of laser beams.
7.When the temperature reaches 80C, turn off laser, and put the hyperpolarized cell
inside of cold water.
8.The hyperpolarized cell is ready to be used.
66
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