Hardware for Performing Hyperpolarized Helium Imaging on
a Clinical MR Imager
by
Angela C. Tooker
Submitted to the Department of Electrical Engineering and Computer Science
in Partial Fulfillment of the Requirements for the
Degrees of Bachelor of Science in Electrical Engineering and Computer Science
and Master of Engineering in Electrical Engineering and Computer Science
at the MASSACHUSETTS INSTITUTE OF TECHNOLOGY
January 24, 2002
MAsSACHUSETTS INSTITUTE
OF TECHNOLOGY
JUL 3 1 2002
© 2002 Angela C. Tooker. 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.
LIBRARIES
Author
Department of Electrical Engineering and Computer Science
January 24, 2002
Certified b3
Mitchell Albert
Thesis Supervisor
Certified by_
Frederick Bowman
Thesis Supervisor
Certified by
Chairman, Department Committee on Gn,
. Smith
Ae Theses
Hardware for Performing Hyperpolarized Helium Imaging on
a Clinical MR Imager
by
Angela Tooker
Submitted to the Department of Electrical Engineering and Computer Science
January 24, 2002
In Partial Fulfillment of the Requirements for the Degrees of
Bachelor of Science in Electrical Engineering and Computer Science
and Master of Engineering in Electrical Engineering and Computer Science
Abstract
For the past forty years, magnetic resonance imaging (MRI) has been one of the most
common methods for obtaining volume images of human organs. However, some organs
in the human body, such as the lungs, are resistant to these techniques, due to their lack of
protons. In 1994, hyperpolarized noble gas MRI was introduced, and today is viewed as
the best solution to the MR imaging dilemma. The different resonance frequencies of
proton and helium, however, present an obstacle for hyperpolarized helium imaging.
Conventional MR scanners are proton-based and, thus, are incapable of transmitting and
receiving at frequencies other than that of proton. Broadband MR scanners, which are
capable of transmitting and receiving at a variety of frequencies, are expensive and are
not readily available for all nuclei. A heterodyne system is an inexpensive system that
can be added to conventional MR systems to enable them to image a variety of nuclei.
This system transforms the output of the MR system, at the frequency of proton, to the
frequency of the nuclei of interest. The resulting signal obtained from the subject is
transformed back to the frequency of proton by the heterodyne system, thereby enabling
the MR system to create an image of the subject. With this system, hyperpolarized
helium images of human lungs were obtained using a variety of different protocols. This
heterodyne system opens up the possibility of new types of imaging, based on nuclei
other than proton, on conventional MR scanners.
Thesis Supervisor: Mitchell Albert
Title: Project Supervisor
Thesis Supervisor: Frederick Bowman
Title: MIT Thesis Supervisor
2
Acknowledgements
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 provide support. Special thanks to
Professor Bob Lenkinski for providing the basic design for the heterodyne system, and
for helping in the inital setup and testing of the system. Thanks to my lab-mates, Arvind
Venkatesh and Adelaide Zhang, who taught me so much and helped during the initial
setup and testing of the heterodyne system. Many thanks to Ken Hickson, the GE Field
Engineer, for helping me debug the gating problem and for his continued help with the
intricacies of making our equipment talk to the Signa equipment. Many thanks to Ralph
Hashoian who taught me so much about RF coils and mixers, and who always had
wonderful suggestions for solving the problems with our coils, our T/R switch and preamplifier, and the Signa.
I would like to thank my MIT thesis supervisor, Dr. Frederick Bowman, for his
support and help with the seemingly endless paperwork.
Thanks to Dr. Kwan Soo Hong, who always performed the experiments with me.
Special thanks to Erin McKinstry for her help with the experiments and for her
suggestions regarding my thesis. Finally, many thanks to my parents for all their support
during my years at MIT and for helping to proofread my thesis.
3
Contents
1 Introduction.....................................................................................................................8
2 Magnetic Resonance Theory...................................................................................
10
2.1 Atoms and Magnetic Dipole Moments.......................................................
10
2.2 Precession and the Resonance (Larmor) Frequency....................................
13
2.3 Bulk Magnetization......................................................................................14
16
2.4 Radio Frequency Coils and Pulses..............................................................
2.5 Magnetic Resonance Imaging and Gradients..............................................18
3 Hyperpolarized Noble Gas MRI...............................................................................20
3.1 Introduction................................................................................................
. 20
3.2 Method for Hyperpolarizing 3He................................................................
21
3.3 Advantages of Hyperpolarized Noble Gas MRI.........................................
21
3.4 Disadvantages of Hyperpolarized Noble Gas MRI.....................................22
4 Heterodyne System Design...........................................................................................24
4.1 B ackground.................................................................................................
24
4.2 Overview of the Heterodyne System..........................................................
25
4.3 Transmit Channel........................................................................................
26
4.4 Radio Frequency Amplifier........................................................................
29
4.5 Transmit/Receive Switch, Pre-Amplifier, and Coil.....................................29
4.6 Receive C hannel..........................................................................................
31
5 Heterodyne System Results and Discussion...............................................................33
5.1 Hyperpolarized 3He Human Lung MR Images............................................33
5.2 Voltage Lost Considerations.......................................................................
35
5.3 Gradient R eversal..........................................................................................
38
5.4 Gating and the Radio Frequency Amplifier................................................
40
4
6 C onclusions and Future W ork.................................................................................
42
6.1 Conclusions.................................................................................................
42
6.2 Recom m endations for Future W ork............................................................
43
R eferences.......................................................................................................-
5
...... ....-44
List of Figures
Figure 2.1 A spinning charged nucleus creates a magnetic field, similar to that created by
a bar magnet.
Figure 2.2 The direction of the magnetic field depends directly on the direction in which
the nucleus is spinning.
Figure 2.3 (A) The magnetic fields for paired protons cancel, resulting in no net
magnetic field.
(B) Unpaired protons, on the other hand, create a net magnetic field.
Figure 2.4 With no external magnetic field, the magnetic dipole moments from
individual nuclei cancel, resulting in a zero net magnetic field.
Figure 2.5 In the presence of an external magnetic field, B0 , the axes of the magnetic
dipoles align with the field, creating net magnetization.
Figure 2.6 (A) With no external magnetic field, the nucleus spins about its dipole axis.
(B) In the presence of an external magnetic field, the nucleus spins about its
dipole axis and precesses about the external magnetic field axis.
Figure 2.7 Motion of the magnetization vector, after a RF pulse has been applied, in both
(A) stationary and (B) rotating frames of reference.
Figure 2.8 Decaying sine wave of the received signal (FID).
Figure 2.9 When a linear gradient is applied to a body, the effective applied magnetic
field at different points in the body change. This, in turn, changes the
resonance frequency of those points.
Figure 4.1 Overview of the heterodyne system (shaded in gray) and the connections to
the MR system, RF Amplifier, T/R switch, pre-amplifier, and coil.
Figure 4.2 The transmit channel of the heterodyne system.
Figure 4.3 The receive channel of the heterodyne system.
Figure 5.1 Proton MR image of the lungs in the coronal plane.
6
Figure 5.2 3 He image of the lungs in the coronal plane, obtained using the heterodyne
system.
Figure 5.3 3He image of the lungs in the axial plane, obtained using the heterodyne
system.
Figure 5.4 3He image of the lung airways obtained using the heterodyne system.
Figure 5.5 Voltage output by the MR system and the transmit channel of the heterodyne
system.
Figure 5.6 Proton frequencies for different slices of a body in a 1.5 T external magnetic
field, with an applied linear gradient.
Figure 5.7 Helium frequencies for different slices of a body in a 1.5 T external magnetic
field, with an applied linear gradient. Notice the reversal of the applied
linear gradient.
Figure 5.8 Proton frequencies, and their corresponding helium frequencies for different
slices of a body in a 1.5 T external magnetic field, with an applied linear
gradient. Notice, each proton slice corresponds to the opposite slice for
helium.
7
Chapter 1
Introduction
Conventional magnetic resonance imaging (MRI) techniques are proton based.
By exciting the protons in water and measuring the resulting output signals, images of a
variety of structures within the body can be obtained. However, these techniques provide
little useful information about the lungs because the inhomogeneous air-tissue interfaces
in the lungs make it difficult to acquire the signal fast enough. Consequently, other
techniques are needed to make MR images of the lungs. One of the most promising
techniques
involves
the use
of hyperpolarized
noble
gas
[1,2].
After the
hyperpolarization process, the 3He or 129Xe has a signal 105 times greater than its
equilibrium signal, making it easily detectable by MRI. Using a specialized MR scanner,
hyperpolarized helium MR images of the lungs that are considerably more detailed than
conventional proton images can be obtained. Hyperpolarized helium images can be used
to discern information about the lungs for possible diagnoses, tracking the time-course of
pulmonary diseases, such as cystic fibrosis, asthma, and emphysema [2-6], and staging
the effectiveness of various treatments.
8
Proton and helium have different resonance frequencies.
Conventional MR
systems are only able to transmit and receive at the resonance frequency necessary to
excite protons and, hence, cannot be used to excite other nuclei, such as helium. There
are specialized MR systems that can be used to image nuclei other than proton. Their
high cost, however, prevents most institutions from obtaining these systems, and for
some nuclei, these specialized MR systems are not available.
The goal for this project is to develop an inexpensive heterodyne system that can
be added to an MR system, enabling it to transmit and receive at a variety of frequencies.
This system can be used to image a variety of nuclei on any conventional MR system.
The particular application of interest for this project is hyperpolarized helium MR of the
lungs. With the heterodyne system, conventional MR scanners can be used to obtain
hyperpolaried helium MR images of the lungs, showing considerable detail from the lung
periphery and airway structures.
9
Chapter 2
Magnetic Resonance Theory
2.1 Atoms and Magnetic Dipole Moments
The nucleus of an atom, comprised of protons and neutrons, has a positive charge.
Each proton within the nucleus spins, creating an electromagnetic field, like a typical bar
magnet (Figure 2.1) [7-11].
-- Greater magnetic field
<- Spinning nucleus
with charge
Bar Magnet
Figure 2.1 A spinning charged nucleus creates a magnetic field, similar to that
created by a bar magnet.
The hydrogen nucleus, containing a single proton, has a spin quantum number of
Hence, there are two possible energy states, denoted
directions in which the protons may spin.
10
-2
'/2.
and + 2 , and two possible
Due to the different spinning directions
possible for the protons, magnetic fields are created in opposite directions (Figure 2.2),
which, by analogy to the bar magnet, are often referred to as north and south.
Direction of
magnetic field
Direction of spin
Direction of
magnetic field
Figure 2.2 The direction of the magnetic field depends
directly on the direction in which the nucleus is spinning.
If there are even number of protons in the nucleus, then each proton with a
magnetic field pointing north has a paired proton with a magnetic field pointing south.
These anti-parallel magnetic fields cancel out, creating a zero net magnetic field (Figure
2.3A).
On the other hand, if there are an odd number of protons, there will be one
unpaired proton. This unpaired proton, pointing either north or south, creates a non-zero
net magnetic field, or a magnetic dipole moment (MDM) represented by the vector p
(Figure 2.3B).
(Net magnetic
field)
(No magnetic
field)
Paired protons
Unpaired protons
B
A
Figure 2.3 (A) The magnetic fields for paired protons cancel, resulting in no net
magnetic field. (B) Unpaired protons, on the other hand, create a net magnetic field.
11
Each individual proton, in a collection of protons, spins about its own axis and has
its own magnetic field, or magnetic dipole moment (MDM) [7-11]. Each axis points in a
random direction and the resulting sum of the individual MDMs is zero (i.e. there is no
net magnetic field) (Figure 2.4).
:Net MF= 0
Bo off:
Figure 2.4 With no external magnetic field, the magnetic dipole moments
from individual nuclei cancel, resulting in a zero net magnetic field.
If this collection of protons is placed in an external magnetic field, B0 , the axes of the
individual MDMs will align along the direction of B 0 . Some will point in the same
(parallel) direction as h 0 and some will point in the opposite (anti-parallel) direction
(Figure 2.5).
Bo on:
4+4+44
Figure 2.5 In the presence of an external magnetic field, B 0, the axes of
the magnetic dipoles align with the field, creating net magnetization.
Initially, exactly half of the MDMs are in the same direction as B 0 and half are in the
opposite direction.
Over time, however, the number of MDMs pointing in the same
direction as B. increases, creating a net magnetization, M, in the direction of B0 . The
net magnetization grows exponentially with time:
12
(Eqn. 2.1)
M = N(H)(1 - e~"I,)
where T is the relaxation time and N(H) is the density of mobile protons in the sample.
Eventually, all of the dipole axes of the protons will be parallel to B0 .
2.2 Precession and the Resonance (Larmor) Frequency
If a proton is placed in an external magnetic field, its own small MDM causes it to
both spin about its own MDM axis and precess about the axis of the external magnetic
field (Figure 2.6B) [7-11].
Bo OFF
Bo ON
B
A
Figure 2.6 (A) With no external magnetic field, the nucleus spins about its
dipole axis. (B) In the presence of an external magnetic field, the nucleus spins
about its dipole axis and precesses about the external magnetic field axis.
The angular frequency of this precession is called the Larmor, or resonance, frequency
(co); it is proportional to both the magnitude of the external magnetic field (B0 ) and the
gyromagnetic ratio (y ) of the nucleus
(Eqn. 2.2)
co = B0 y .
13
Since the gyromagnetic ratio is element specific, different atoms will have different
resonance frequencies even when the magnetic field strengths are the same.
For
example, the gyromagnetic ratio of hydrogen is y, = 42.6MHz / Tesla, whereas the
gyromagnetic ratio of helium is YHe = 32.4MHz / Tesla . Therefore, at a magnetic field
strength of BO = 1.5Tesla, the resonance frequency of hydrogen is O H=
the resonance frequency of helium is coH,4
63.87MHz and
= 48.65MHz.
2.3 Bulk Magnetization
Due to its natural abundance in the human body, hydrogen is the basis of the
signal in conventional magnetic resonance imaging (MRI).
When a human body is
placed in an external magnetic field (B 0 ), each individual proton within the sample aligns
its MDM (i ) axis either parallel or anti-parallel to the external magnetic field (see
Figure 2.5) [7-11]. Hence, a magnetization vector M can be defined which is the sum of
the individual MDMs within the sample:
NS
AI = I
(Eqn. 2.3)
I,
n2=1
where NS is the total number of spinning protons in the body and i, is the MDM for the
nth proton. From quantum theory, there are two possible spin states for proton. The
Boltzmann relationship relates the population difference to the energy difference in these
two spin states:
N = exp
N
,E
(Eqn. 2.4)
KT,
14
where N, is the number of spins pointing upward, NI is the number of spins pointing
downward, AE is the energy difference between the two spin states, K is the Boltzmann
constant (1.38 x 10-23 J/K), and T, is the absolute temperature of the body. Since in
practice,
(Eqn. 2.5)
AE << KT
the ratio can be rewritten
I ~1+ A
N
KT
(Eqn. 2.6)
where y is the gyromagnetic ratio, h is Planck's constant (6.6 x 1034 J-s), and B0 is the
magnitude of the external magnetic field. Using the relations
NS= N +NU
( qn. . )
NS ,~ 2N-2Nt
where NS is the total number of spins, the above ratio in Eqn. 2.6 can be rewritten as
NT-
N
(Eqn. 2.8)
NhB
2kTS
This indicates there is a small excess of spins in the upward direction. This difference
occurs since a spin will more likely be in the lower-energy state (with an upward pointing
spin) than in the higher-energy state (with a downward pointing spin).
This small
population difference results in a non-zero magnetization vector, M . The magnitude of
the magnetization vector is related to this population difference:
(NT - NJy
2
(Eqn. 2.9)
Hence, the magnitude of the bulk magnetization is
15
=
_
2h2B
0 N~
2 2
NsO
(E q n . 2 .1 0 )
4KI:,
Although the individual protons precess around the external magnetic field,, B,
the bulk magnetization vector, M, does not precess around BO . The individual protons
are out of phase with each other and, hence, when their individual MDMs are added
together, there is a large component in the direction of BO, but the phase differences
cancel each other. Therefore, the bulk magnetization vector does not precess.
2.4 Radio-Frequency Pulses and Coils
A radio-frequency (RF) pulse is an electromagnetic wave that is applied to the
sample via a RF coil. Coils are electrical devices composed of multiple loops of wires
that can generate and/or detect magnetic fields [7-12].
Typically RF pulses of high
intensity and short duration are applied to the body at the resonance frequency of proton,
in a direction perpendicular to the external magnetic field. The RF pulse causes the bulk
magnetization vector to tip away from the external magnetic field and into the plane
containing the rotating RF pulse. The magnetization vector tips because the protons with
a resonance frequency equal to that of the RF pulse absorb the energy in the RF pulse and
are, thus, excited to a higher energy state. (Nuclei with different resonance frequencies
do not absorb any energy and, therefore, are not excited.) While it is tipping away from
equilibrium, and as it is relaxing back (once the RF pulse is turned off), the magnetization
vector precesses around the external magnetic field, B0 , at the resonance frequency.
When viewed from a stationary frame of reference, the protons precess around the
16
external magnetic field axis at the resonance frequency. If, instead, the spinning protons
were viewed in a frame of reference rotating at the resonance frequency, they appear to
be stationary. Hence, when viewing the precessing magnetization vector in the stationary
frame of reference, the magnetization vector traces out a spiral (Figure 2.7A). On the
other hand, if the precessing magnetization vector is viewed in the rotating frame of
reference, the magnetization vector traces an arc into the plane containing the RF pulse
(Figure 2.7B).
Z
Z
Spiral observed from
Simple arc visualized from
a point outside the
a rotating reference point
coordinate system
X
x
RF
B
A
Figure 2.7 Motion of the magnetization vector, after a RF pulse has been
applied, in both (A) stationary and (B) rotating frames of reference.
This rotating magnetization is an oscillating magnetic field that can be detected by the
coil. The signal detected is a damped oscillating sine wave called the free induction
decay (FID) signal (Figure 2.8).
Received
signal
Time
Figure 2.8 Decaying sine wave of the received signal (FID).
17
2.5 Magnetic Resonance Imaging and Gradients
To create a magnetic resonance image (MRI), gradients are used to obtain
information about the volumes of interest [7,8,10]. If no gradients are applied, the entire
body would see the same magnetic field, B0 . Hence, the FID received by the coil as a
result of the RF pulse would be from the entire body. Gradients are used to vary the
external magnetic field applied to the subject, thereby changing the resonance frequencies
of different slices within the body. Consider a body in an external magnetic field, B0,
shown in Figure 2.9.
-
1.55T1.5T=Bo
Figure 2.9 When a linear gradient is applied to a body, the
effective applied magnetic field at different points in the
body change. This, in turn, changes the resonance frequency
of those points.
To define various slices within the sample, a linear gradient is applied. As shown in
Figure 2.9, this applied gradient changes the effective magnetic field seen at each point in
the body. Since the resonance frequency varies with the applied magnetic field, this
causes a change in the resonance frequency at each point. Therefore, if a RF pulse is
applied at a particular resonance frequency, only one corresponding line of the body is
excited. By applying a RF pulse at a range of frequencies (a bandwidth), many lines
within the body, i.e. a "slice," can be excited. As the bandwidth increases, a larger slice
18
within the body will be excited by the RF pulse. To excite other slices, RF pulses at
different frequencies need to be applied. Once the gradient has been applied, any FID
resulting from the application of an RF pulse with a specified bandwidth will contain
information from only one slice of the body.
By applying gradients in varying
configurations, using all three spatial directions, specific volumes within the sample can
be specified. Any signal obtained after the RF pulse will be due only to the nuclei in the
specific region being excited. The strength of the received signal is directly proportional
to the number of protons excited in the specific volume. An MR image of the body can
be obtained using various transforms once the positions and number of the protons are
known.
19
Chapter 3
Hyperpolarized Noble Gas MRI
3.1 Introduction
The field of hyperpolarized noble gas MRI is rapidly growing. One limitation of
conventional proton-based MRI is that regions lacking water protons, such as the lungs or
the lipid bilayer membranes of the brain, cannot be imaged. Other techniques are needed;
hyperpolarized noble gas MRI is one technique that can be used to image these
inhomogeneous regions of the lungs [2,13]. Hyperpolarized 3He images of the lungs, in
both humans, rats, dogs, and pigs, have already been obtained.
In addition,
hyperpolarized 129Xe images have been obtained of the lung and brain structures in rats
and humans.
Current hyperpolarized noble gas MRI requires either the inhalation of the gas by
the subject or the injection of the gas in the form of microspheres [2,13,14]. The MR
image is then created from the signal obtained after applying a pulse at the proper
resonance frequency.
Thermal equilibrium polarization of noble gases is only about
0.04%, which is insufficient for imaging.
20
However, using a laser optical pumping
process, the polarization can be increased by a factor of 105, making MR imaging of
noble gases possible in the human body.
3.2 Method for Hyperpolarizing 3He
A laser optical pumping process is used to hyperpolarize 3He. In this method, the
helium nuclei undergo a spin exchange with optically pumped rubidium (Rb) atoms
[2,13,15-17]. Rb is an alkali metal that vaporizes at 85 'C. When vaporized and placed
in an external magnetic field, Rb splits into different electron spin states under the
excitation of a diode laser.
Ground state electrons in Rb atoms can be excited by
absorbing photons of light with a wavelength of 794.7 nm. If these polarized Rb atoms
are then brought together with 3He atoms, the collisions between the atoms will result in
the transfer of angular momentum from Rb valence electrons to the 3He nuclei. This spin
exchange process increases the noble gas spin population by about 25% over the
equilibrium state and, therefore, enhances the resulting MR signal by up to 105 times its
thermal equilibrium value, making MR imaging possible.
A similar process can be used to hyperpolarize other noble gases, such as
12 9Xe.
To achieve a polarization of about 10-20% requires approximately seven hours for 'He.
If the hyperpolarized noble gas is cryogenically cooled and kept in a magnetic field, the
polarization can last for hours, even days.
3.3 Advantages of Hyperpolarized Noble Gas MRI
There are several advantages to using hyperpolarized noble gases for MRI. The
MR signal from hyperpolarized noble gases can be up to ten times larger than the signal
21
from a corresponding volume of protons. Hence, a new range of exploitable responses
and contrast, as well as increases in space and time resolutions, can be obtained
[1,2,13,15,16].
In addition, since the polarization process in hyperpolarized MRI is
independent of magnetic field strength, less expensive, low-field (< 0.1 Tesla) magnets
can be used.
Currently, people with metal implants, such as aneurysm clips or
pacemakers, cannot undergo MRI. With low-field magnets, it might be possible for these
people to avail themselves of this technology [18].
Another advantage of hyperpolarized noble gas MRI is that it is possible to
visualize biological systems invisible to proton-based MRI techniques.
For example,
conventional MRI detects the signal from the water proton molecules in the body.
However, the lung-space is a nearly water-free environment. As such, details about the
lungs are not visible in MR images. Hyperpolarized helium and xenon, however, can be
used to discern information about the lungs in MR images [2,13], for possible diagnoses,
tracking the time-course of pulmonary diseases, such as cystic fibrosis, asthma, and
emphysema, and staging the effectiveness of various treatments [2-6].
lipid-filled
regions
of the brain cannot be
imaged with
Likewise, the
conventional
MRI.
Hyperpolarized xenon, which rapidly dissolves in lipids, can be used to image these
regions of the brain [1]. Such information could be used in the diagnosis, treatment, and
staging of various white-matter diseases, such as multiple sclerosis.
3.4 Disadvantages of Hyperpolarized Noble Gas MRI
Despite these
promising
hyperpolarized noble gas MRI.
advantages,
there
are
two
disadvantages
for
First, hyperpolarizing helium or xenon is a time-
22
consuming, costly procedure. The hyperpolarization process can take upwards of seven
hours and there is a limit to the amount of gas that can be hyperpolarized at one time.
The second disadvantage for hyperpolarized noble gas MRI is that only
specialized MR systems can be used, due to the differences in the resonance frequencies
of proton and 3He and 129Xe. The conventional MR systems that are routinely used are
incapable of transmitting and/or receiving at a frequency other than that of proton. Only
some MR systems, equipped with specialized hardware can transmit and receive at a
variety of different frequencies.
The large cost or lack of availability prevents most
institutions from adding to their conventional MR systems the specialized hardware
necessary for performing MR imaging of other nuclei. Thus, before hyperpolarized noble
gas MRI can be used routinely, an inexpensive method for imaging other nuclei on
conventional MR scanners must be found. The heterodyne system is an inexpensive
means that enables conventional MR scanners to image other nuclei, such as helium.
23
Chapter 4
Heterodyne System Design
4.1 Background
Conventional MR uses the hydrogen proton as the nucleus for imaging. However,
it is possible to use nuclei other than proton, such as atomic elements with intrinsic spins,
like the noble gases, helium and xenon, for MR imaging. In order to image using these
nuclei, specialized MR systems with broadband hardware are required. MR systems,
with broadband hardware, are capable of transmitting and receiving a variety of
frequencies and are used to image different nuclei. In contrast, conventional MR systems
have narrowband hardware and are only capable of transmitting and receiving at the
frequency of proton; thus, these systems cannot be used for imaging other nuclei.
Specialized MR systems with broadband hardware are very expensive.
Thus, most
institutions have narrowband MR systems. Given that most MR scanners in both clinical
and research environments are narrowband and that imaging of other nuclei, such as
helium and xenon, is rapidly growing, an important question is how can conventional MR
scanners be inexpensively converted to allow imaging of these other nuclei?
24
An inexpensive solution to this problem is the heterodyne system [19].
A
heterodyne system can be added to conventional MR systems allowing imaging to be
performed on a variety of different nuclei, including hyperpolarized helium.
4.2 Overview of the Heterodyne System
There are two main components to the heterodyne system: the transmit channel
and the receive channel.
(The heterodyne system was designed and tested for
hyperpolarized helium MRI at 1.5 T. With only minor modifications, however, this
system can be used for a variety of different nuclei at different magnetic field strengths.)
The primary task of the transmit channel is to convert the signal transmitted by the MR
system from 63.87 MHz (the resonance frequency for proton at 1.5 T) to 48.65 MHz (the
resonance frequency for helium at 1.5 T). For the receive channel, the primary task is to
convert the 48.65 MHz helium signal received from the coil to the 63.87 MHz proton
signal that is detectable by the MR system.
Figure 4.1 contains a diagram of the
heterodyne system (shaded in gray) and its connection to the MR system. The RadioFrequency (RF) amplifier, Transmit/Receive (T/R) switch, pre-amplifier, and coil are not
part of the heterodyne system; they are required, however, for imaging of all nuclei on all
MR systems.
25
MR SYSTEM +
Broadband
RF Amplifier
T/R Switch
COIL
FPre-Amplifier
Figure 4.1 Overview of the heterodyne system (shaded in gray) and the
connections to the MR system, RF Amplifier, T/R switch, pre-amplifier, and coil.
4.3 Transmit Channel
The main goal of the transmit channel is to convert the 63.87 MHz proton signal
transmitted by the MR system to the 48.65 MHz signal needed to excite the helium
nuclei. A frequency mixer and a filter can be combined together to convert the original
proton signal. (See Figure 4.2 for a diagram of the transmit channel.) The frequency
mixer has two inputs and outputs their convolution. The two inputs to the mixer are: 1)
the 63.87 MHz signal output by the MR system and 2) the local oscillator, which is a
pure sine wave at 112.52 MHz. The local oscillator is a Hewlett-Packard (Palo Alto,
California) 8648B Signal Generator. The voltage output by the local oscillator is 300
mVrms.
The frequency mixer is a Level 7 ZP-3 double-balanced mixer from Mini-
Circuits (Brooklyn, New York). The resulting output of the frequency mixer is the
convolution of the two input signals, i.e. the sum of two sine waves one at
112.52MHz - 63.87MHz = 48.65MHz
26
(Eqn. 4.1)
and one at
112.52MHz + 63.87MHz = 176.39MHz.
(Eqn. 4.2)
Thus, by changing the frequency of the local oscillator, the sine waves output by the
frequency mixer changes accordingly. The heterodyne system, therefore, can be used to
image other nuclei, by changing the frequency of the local oscillator.
Mini-Circuits ZP-3
Level 7 Frequency Mixer
Transmitted Proton
Signal (63.87 MHz)
from MR System
Local Oscillator (112.52 MHz)
Hewlett-Packard 8648B
Signal Generator
Mini-Circuits BLP-90
Low-Pass Filter
Transmitted Helium
Signal (48.65 MHz) to
RF Amplifier
Figure 4.2 The transmit channel of the heterodyne system.
The signal output by the MR system is not continuous; rather, it is a pulsed signal.
The frequency mixer does not require a continuous signal input, and outputs a signal only
when receiving input. Hence, the output of the mixer has the same pulsed shape as the
original, only the frequency of oscillation changes.
The same local oscillator is used in both the transmit and receive channels. It is
important that the local oscillator used in both channels has the same frequency. To
eliminate the need for another signal generator, a power splitter is used to create the
27
signals needed for both channels. A power splitter is used to split the output of the local
oscillator into two identical signals, each with the same frequency as the original and
approximately half the amplitude of the original. (The power splitter is a ZFSC-2-2 2Way-00 Power Splitter/Combiner from Mini-Circuits.)
If, instead, the local oscillator outputs a sine wave with a frequency of 15.23
MHz, the frequency mixer would output the sum of two sine waves: one at 48.65 MHz
and the other at 79.09 MHz. Either local oscillator frequency, 112.52 MHz or 15.23
MHz, could be used in the heterodyne system, since both produce the desired resonance
frequency of helium, 48.65 MHz. The local oscillator frequency is set to 112.52 MHz to
enable better filtering. To avoid damaging the coil and other electrical components, as
well as for the safety of the patient, it is necessary that only the desired frequency, 48.65
MHz, leave the heterodyne system. The other frequency output by the frequency mixer
must be eliminated. The frequency difference between the two sine waves is 127.74
MHz for the local oscillator frequency of 112.52 MHz and only 30.44 MHz for the local
oscillator frequency of 15.23 MHz. It is easier to filter a frequency that is 127 MHz away
from the desired frequency than a frequency that is only 30 MHz away from the desired
frequency. Hence, the higher local oscillator frequency, 112.52 MHz, was chosen.
A low-pass filter is used to filter the sine wave at 176.39 MHz, leaving only the
desired 48.65 MHz signal. A BLP-90 Low-Pass Filter (Mini-Circuits) is used. This filter
has a passband from DC to 81 MHz (i.e. only frequencies in the 0 - 81 MHz range pass
through un-attenuated). Hence, upon leaving the filter, the sine wave at 176.39 MHz has
negligible amplitude, while the sine wave at 48.65 MHz has no decrease in amplitude. (If
other nuclei are used, the output of the frequency mixer will contain different frequencies
28
and, thus, a different filter will be required.) Therefore, with a frequency mixer and a
filter, the transmit channel of the heterodyne system converts the proton signal output by
the MR system to the desired signal at the frequency of helium.
4.4 Radio-Frequency Amplifier
Once a signal of the desired frequency has been obtained by the transmit channel,
it needs to be amplified before it is sent to the coil so that the pulse will be strong enough
to excite the nuclei. Conventional MR systems use narrowband Radio-Frequency (RF)
amplifiers. They are only capable of amplifying a signal at the frequency of proton and
actually attenuate signals at other frequencies. Thus, a separate RF Amplifier is needed
to amplify the signal at the frequency of helium. While a narrowband RF amplifier that
amplifies the frequency of interest, in this case helium, can be used, a broadband RF
amplifier is a better choice as it amplifies a wide range of frequencies and, thus, can be
used for a variety of different nuclei.
(Broadband RF amplifiers, while often more
expensive than narrowband RF amplifiers, are much less expensive than specialized
broadband MR systems.) Hence, with the broadband MRI-2000 Linear Pulse Amplifier
(ENI) used in this system, the 48.65 MHz signal transmitted to the coil will be large
enough to excite the helium nuclei.
4.5 Transmit/Receive Switch, Pre-Amplifier, and Coil
A Transmit/Receive (T/R) switch is used to transmit the signal from the RF
amplifier to the coil, as well as to receive the signal from the coil. The T/R switch must
be tuned to operate at the frequency of interest, in this case, the frequency of helium,
29
48.65 MHz. The MR system adds a 15 V bias to the transmitted signal traveling from the
RF amplifier to the T/R switch. This bias tells the T/R switch that the signal must be
transmitted to the coil. The T/R switch removes the DC bias from the signal and then
transmits it to the coil
The coil must be tuned to operate at the frequency of interest, 48.65 MHz. The
coil employed is a flexible, quadrature coil built by Midwest RF, LLC. The coil wraps
around the torso of the subject so that hyperpolarized helium images of the lung can be
obtained. This coil was specially built for lung imaging; however, any type of coil, for
any application, could be used with the heterodyne system.
The signal detected by the coil from the excited helium nuclei travels back to the
T/R switch.
Since this signal has no DC bias, the T/R switch recognizes it as a
"received" signal and transmits it to the pre-amplifier. The received signals are generally
very small, so the pre-amplifier is used to increase the gain of the signal. Like the T/R
switch and the coil, the pre-amplifier must be tuned to the proper frequency, 48.65 MHz.
The output of the pre-amplifier, at 48.65 MHz, is then sent to the receive channel of the
heterodyne system.
The T/R switch, pre-amplifier, and coil are not part of the heterodyne system,
although they are required for hyperpolarized helium imaging, or imaging of any other
nuclei, on all MR systems.
30
4.6 Receive Channel
Mini-Circuits ZP-3
Level 7 Frequency Mixer
Received
Helium Signal
(48.65 MHz)
from Coil
Local Oscillator
(112.52 MHz) HewlettPackard 8648B Signal
Generator
Received Proton
Signal (63.87 MHz)
to MR System
Figure 4.3 The receive channel of the heterodyne system.
The receive channel is similar to the transmit channel (see Figure 4.3). The signal
received from the coil is at the frequency of helium, 48.65 MHz. However, conventional
MR systems are incapable of detecting this frequency. Thus, the receive channel must
convert the 48.65 MHz helium signal received from the coil (via the T/R switch and preamplifier) back to the frequency of proton, 63.87 MHz. This is done in a manner similar
to that for the transmit channel. Another ZP-3 Level 7 double-balanced frequency mixer
from Mini-Circuits is used (along with the same local oscillator as in the transmit
channel) to convert the received signal at 48.65 MHz to the frequency of proton, 63.87
MHz. The output of the mixer will be the sum of two sine waves: one at
112.52MHz - 48.65MHz = 63.87MHz
(Eqn. 4.3)
and one at
112.52MHz+ 48.65MHz =161.17MHz.
31
(Eqn. 4.4)
Since the MR system is only capable of receiving signals at the frequency of proton, no
external filter is necessary to filter the 161.17 MHz sine wave from the mixer output.
Since only the frequency of oscillation of the received signal is changed, the MR system
can render an image in the same manner as for proton.
The receive channel has extra, protective circuitry, as shown in Figure 4.3. The
T/R switch requires a 15 V bias in order to transmit the signal to the coil. The inductor
and capacitors are used to prevent this DC bias from damaging the frequency mixer and
the MR system, and to allow the oscillating received signal to reach the frequency mixer.
With this circuitry, the receive channel of the heterodyne system is able to convert the
received helium signal, at 48.65 MHz, to a signal at the frequency of proton, 63.87 MHz,
that the MR system then transforms into an image.
32
Chapter 5
Heterodyne System Results and Discussion
5.1 Hyperpolarized "He Human Lung MR Images
Conventional proton MR images of the lungs provide limited information. As
shown in Figure 5.1, the lungs appear as a black void in proton MR images; no
information regarding the lung periphery or airway structure is visible. (The white areas
within the black voids of the lungs are the blood vessels in the lungs.)
Figure 5.1 Proton MR image of
the lungs in the coronal plane.
3
In contrast, hyperpolarized He MR images of the lungs show considerably more detail.
Figure 5.2 shows a 3He MR image of one slice of a subject's lungs obtained on the GE
33
Signa LX 1.5 T MRI system (GE Medical Systems, Milwaukee, WI), running Software
Revision 8.4, equipped with the heterodyne system described in Chapter 4. With the
attached heterodyne system, the pulse sequences and options available on the Signa MRI
system can be used.
Figure 5.3 3He image of the lungs
in the axial plane, obtained using
the heterodyne system.
Figure 5.2 3 He image of the lungs
in the coronal plane, obtained
using the heterodyne system.
Figure 5.3 shows one slice of a 3He image of a subject's lungs in the axial plane, as
opposed to the image in Figure 5.2, which is in the coronal plane. The heterodyne system
does not limit the planes in which the images can be obtained. In the images shown in
Figures 5.2 and 5.3, only the lung periphery is visible; in contrast, lung airways are also
visible in the image shown in Figure 5.4. A pulse sequence different from the one used
for the images in Figures 5.2 and 5.3 was used to obtain the image in Figure 5.4. Thus,
the various pulse sequences and imaging options available on the MR scanner for
conventional proton imaging can also be used with the heterodyne system.
34
Figure 5.4 3 He image of the lung airways
obtained using the heterodyne system.
As demonstrated with these images, the heterodyne system enables conventional MR
scanners to image other nuclei. The pulse sequences and imaging options available on
the MR scanner are available for use with the heterodyne system to perform a variety of
different types of imaging for various clinical and experiment protocols.
With
hyperpolarized helium, these options can be used to visualize lung periphery and airway
structures.
5.2 Voltage Loss Considerations
If the power reaching the coil is not sufficient to excite the nuclei, no image will
be obtained. Hence, it is important to minimize the power lost by the heterodyne system.
Due to the signal conversions performed by the heterodyne system, there is the possibility
was
for signal loss. The voltage output by both the MR system and the transmit channel
measured. The voltage output by the MR system can be altered in a variety of different
ways; for these measurements, the voltage output is varied by changing the transmit gain
35
(TG), i.e. the power in the RF pulses transmitted to the coil.
Larger TG values
correspond to larger power outputs. As can be seen in Figure 5.5, the voltage output by
the low-pass filter in the heterodyne system is significantly smaller than the voltage
output by the MR system. For large inputs (i.e. large TG values), more than 50% of the
input signal is lost. For smaller inputs, the signal loss is smaller; however, it is never
zero.
Voltage Output
300
E 200
20-+-
E 100
MR System
+
-
Transnit Channel
0
0
200
100
Figure 5.5 Voltage output by the MR system and the
transmit channel of the heterodyne system.
The voltage loss is due to the frequency mixer. The particular frequency mixers
used in the heterodyne system are double-balanced mixers [12].
One of the
characteristics of double-balanced mixers is a conversion loss, i.e. as the convolution of
the input signals is performed, voltage is lost between the input and output signals. 50%
of the voltage can be lost, depending on the input signal and the local oscillator. To
ensure minimal conversion loss by the frequency mixers, the voltage from the input
signal (the signal from the MR system) must be less than the voltage from the local
oscillator. Thus, as the input voltage from the MR system increases (i.e. as the TG
increases), the voltage lost by the frequency mixer increases (as seen in Figure 5.5). This
implies that for a given input voltage from the MR system, less voltage will be lost
36
during conversion for larger local oscillator voltages. These measurements were repeated
for other local oscillator voltages. For larger voltage outputs from the local oscillator, the
voltage lost by the frequency mixer for a given input signal is less than for smaller local
oscillator voltages. However, even for small input signals from the MR system and large
local oscillator voltages, there is always some voltage loss.
Although there is a voltage loss in the transmit channel of the heterodyne system,
it is not detrimental to obtaining hyperpolarized
3 He
images with the GE Signa MR
system and the quadrature coil. One possible reason is, despite the loss in the transmit
channel, the RF amplifier provides enough amplification of the signal so that when it is
transmitted to the coil it is still sufficient to excite the nuclei.
There are several possible methods for eliminating this voltage loss. Instead of
using double-balanced frequency mixers in the heterodyne system, active frequency
mixers can be used [12]. Active frequency mixers require a DC voltage input; with this
extra input, the resulting conversion loss by the frequency mixer is negligible. Hence,
there would be no voltage lost between the input and output signals of the active
frequency mixer. Another possible solution is the addition of an amplification stage.
This would consist of either 1) an amplifier, distinct from the RF amplifier, on the output
of the low-pass filter that would amplify the output signal to offset any signal loss by the
frequency mixer or 2) a different RF amplifier that would provide more amplification.
While all three solutions would solve the voltage loss problem in the heterodyne system,
the active frequency mixer is the best solution. The addition of an amplifier, either at the
output of the low-pass filter or via a different RF amplifier, would result in an amplified
signal, containing amplified noise. The active frequency mixer, in addition to being less
37
expensive than the amplifier, would provide amplification of the signal, without the
simultaneous amplification of the noise. (Similar losses are seen in the receive channel,
and the solutions outlined above can also be used to minimize these losses.)
5.3 Gradient Reversal
As explained in Section 2.5, MR imaging requires the use of gradients. With
these gradients, spatial information about the object being imaged can be obtained
[7,8,10]. The gradients change the applied magnetic field for different slices within the
object. These changes in magnetic field cause changes in the resonance frequency of the
nuclei in those slices. Figure 5.6 shows an applied gradient in a 1.5 T external magnetic
field and the resulting changes in the resonance frequencies of proton for the different
slices. RF pulses are transmitted by the MR system at the different frequencies for each
slice.
Proton Frequencies
for Different Slices
Applied Linear
Gradient
66.87 MHz
65.87 MHz
64.87 MHz
63.87 MHz
62.87 MHz
External
Magnetic Field
BO = 1.5 T
61.87 MHz
60.87 MHz
Figure 5.6 Proton frequencies for different slices of a body in a
1.5 T external magnetic field, with an applied linear gradient.
38
The transmit channel in the heterodyne system converts the frequency for proton to the
frequency for helium. Figure 5.7 shows the resulting helium frequencies output by the
transmit channel for different slices described in Figure 5.6.
As can be seen, this
corresponds to a reversal of the applied gradient field.
Helium Frequencies
for Different Slices
Applied Linear
Gradient
112.52 MHz - 66.87 MHz = 45.65 MHz
112.52 MHz - 65.87 MHz = 46.65 MHz
112.52 MHz - 64.87 MHz = 47.65 MHz
112.52 MHz - 63.87 MHz = 48.65 MHz
External
Magnetic Field
Bo = 1.5 T
112.52 MHz - 62.87 MHz = 49.65 MHz
112.52 MHz - 61.87 MHz = 50.65 MHz
112.52 MHz - 60.87 MHz = 51.65 MHz
Figure 5.7 Helium frequencies for different slices of a body in a 1.5 T
external magnetic field, with an applied linear gradient. Notice the reversal
of the applied linear gradient.
Helium Frequencies for
Corresponding Slices
Proton Frequencies
for Different Slices
Applied Linear
Gradient
66.87 MHz
45.65 MHz
65.87 MHz
46.65 MHz
64.87 MHz
47.65 MHz
63.87 MHz
48.65 MHz
62.87 MHz
49.65 MHz
61.87 MHz
50.65 MHz
60.87 MHz
t
External
Magnetic Field
Bo = 1.5 T
51.65 MHz
X
Figure 5.8 Proton frequencies, and their corresponding helium frequencies for
different slices of a body in a 1.5 T external magnetic field, with an applied linear
gradient. Notice, each proton slice corresponds to the opposite slice for helium.
39
However, since the gradients are applied as shown in Figure 5.6, the RF pulse from the
MR system selecting a particular slice with the proton frequency would be transformed
by the transmit channel of the heterodyne system to an RF pulse selecting the opposite
slice with the helium frequency (as shown in Figure 5.8). Hence, the transmit channel of
the heterodyne system creates an effective gradient reversal.
It is important that the slices in the body be selected and labeled correctly. With
the current implementation of the heterodyne system, the slice selected by the MR system
will not be the slice that is actually imaged. Two possible means for correcting this
problem are: first, the gradients on the MR system can be reversed so that the slices
selected with the proton frequency correspond to the same slices with the helium
frequency. Second, if the local oscillator is changed to the other frequency, as described
in Section 4.3, (15.23 MHz instead of 112.52 MHz), when the transmit channel converts
the frequencies, there will be no apparent gradient reversal. With this second method,
each slice in proton corresponds to the exact slice in helium. (The reason for choosing
the local oscillator frequency of 112.52 MHz is explained in Section 4.3.)
5.4 Gating and the Radio-Frequency Amplifier
Both the RF amplifier used with the heterodyne system and the narrowband RF
amplifier used by the Signa MRI system are pulsed amplifiers. They do not continually
transmit a signal; only when triggered by an external input (called the gating input) will
they transmit the RF pulse. This gating signal turns on before the RF pulse is transmitted
by the MR system and turns off only once the pulse has been transmitted. The gating
signal is a square wave output by the MR system. When an RF pulse is being transmitted
40
by the MR system, the gating signal outputs 5 V, and when no RF pulse is transmitted,
the gating signal outputs 0 V. Hence, the narrowband RF amplifier that is part of the MR
system transmits when it receives a 5 V gating input and does not transmit when it
receives a 0 V gating input. The broadband RF amplifier used with the heterodyne
system, on the other hand, transmits when it receives a 0 V gating input and does not
transmit when it receives a 5 V gating input. Therefore, unless the gating signal output
by the MR system is altered, the broadband RF amplifier used with the heterodyne
system will always be off when the RF pulse is transmitted. This means no signal will be
transmitted to the coil and, thus, no image can be obtained. With the use of an inverter,
the gating signal was transformed. The gating signal was input to an inverter. Thus,
when the gating signal output 0 V (signaling that no RF pulse is being transmitted), the
inverter output 5 V to the RF amplifier, turning it off. Likewise, when the gating signal
output 5 V (signaling that an RF pulse is being transmitted) the inverter outputs 0 V,
thereby turning the RF amplifier on. With this inverter, the broadband RF amplifier is
turned on when then RF pulse is transmitted and, thus, images can be obtained.
41
Chapter 6
Conclusions and Future Work
6.1 Conclusions
Conventional MR techniques use proton as the nucleus for imaging. However,
these techniques provide little information about the proton-free regions of the body, such
as the lungs. Techniques using other nuclei, such as hyperpolarized helium, have shown
much promise for imaging these regions.
Conventional MR scanners, however, are
incapable of imaging nuclei other than proton. With the addition of the heterodyne
system described in Chapter 4, conventional proton MR scanners can image other nuclei,
such as helium. Although broadband MR scanners can also image other nuclei, the
heterodyne system is considerably less expensive and can be added to any existing
narrowband MR scanner. As shown in Figures 5.2 - 5.4, the addition of the heterodyne
system does not limit the functionality of the MR scanner, with regards to potential
imaging planes, pulse sequences, and imaging options. This heterodyne system opens up
the possibility of a variety of new types of imaging, based on nuclei other than proton, on
conventional MR scanners.
42
6.2 Recommendations for Future Work
Future work needs to be done with this system to improve image quality and the
signal-to-noise ratio (SNR).
Active frequency mixers, instead of double-balanced
frequency mixers, should be investigated to minimize the signal loss in the heterodyne
system, thereby improving image quality. In addition, since the heterodyne system is an
inexpensive alternative to broadband MR systems, the image quality and SNR obtained
with the heterodyne system on conventional proton MR scanners should be compared to
that obtainable on broadband MR scanners.
This heterodyne system will work for a variety of different nuclei, different
magnetic field strengths, and on any MR scanner. Thus, further studies can be conducted
with different nuclei, magnetic field strengths, and MR systems.
43
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