mri relaxation measurements in rats and mice · mri relaxation measurements in rats and mice...
TRANSCRIPT
MRI relaxation measurements in rats
and mice
Master’s Thesis, 30hp
Jesper Lindberg
Spring 2008
Supervisor:
Michael Horn
University of Gothenburg
Department of radiation physics
Abstract
The work presented focus on the measurement of proton relaxation times, T1 and T2for organs in mice and the brain in rat at 7 T. The knowledge of T1 and T2 relaxationtimes are necessary for optimizing contrast in imaging methods like functional MRI andperfusion imaging. Relaxation times can be an important way to investigate propertiesof tissues and separate normal tissue from diseased, e.g. cancerous tissue.
The pulse scheme used for the relaxation measurements is IR-TrueFISP, a sequencewhich allows measuring both T1 and T2 within the same scan.
Initial experiments were done in a phantom to investigate the behavior of the pulsesequence. Parameters tested were number of averages, number of frames, ROI size, multi-slice acquisition, intervariability and slice orientation dependencies. Conclusions from thephantom measurements were that ROIs should be larger then 0.27 cm2 and single slicemeasurements are preferred.
Relaxation measurements for mice were performed in the strain C57/BL6. Measure-ments were performed in kidneys, front leg muscle and liver. Relaxation times were(Organ:T1:T2 in ms) kidneys:1559:54, front leg muscle:1190:39 and liver:985:60. Furthermeasurements were performed to investigate the ability of the method to discriminatewhite and brown adipose tissue by relaxation time. Differences could be seen, however,further measurements are needed to establish the method for discrimination of the fatdepots.
Measurements in the rat brain were performed in the strain Sprague Dawley. Re-gions measured were cerebellum, medulla oblongata, cerebral cortex, hippocampus, thala-mus and corpus callosum. Relaxation times were (location:T1:T2 in ms) in cerebel-lum:1343:80, medulla oblongata:1560:68, cerebral cortex (right side):1246:76, cerebral cor-tex (left side):1635:78, hippocampus (right side):1249:171 and hippocampus (left side):1072:254.
Problems that occurred and in some extend affected the results were the lack of trig-gering, eddy current compensation, and magnetic inhomogeneities.
Contents
1 Introduction 4
2 MR theory 4
2.1 Behavior in an external magnetic field . . . . . . . . . . . . . . . . . . . . . . . 4
2.2 Measuring the net magnetization . . . . . . . . . . . . . . . . . . . . . . . . . . 6
2.3 Relaxation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6
2.4 Inversion recovery technique . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7
2.5 Gradient, slice selection and image encoding . . . . . . . . . . . . . . . . . . . . 7
2.6 Echo creation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8
2.6.1 Gradient echo . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8
2.6.2 Spin echo . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9
2.7 Relaxation measurements with IR-True-FISP . . . . . . . . . . . . . . . . . . . 9
3 Materials and Methods 10
3.1 Phantom . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10
3.1.1 Number of averages . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10
3.1.2 Number of frames . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11
3.1.3 Size of the ROI . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11
3.1.4 Single vs. multislice . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11
3.1.5 Intervariability . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11
3.1.6 Slice orientation dependencies . . . . . . . . . . . . . . . . . . . . . . . . 12
3.2 Mouse . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12
3.2.1 Organs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12
3.2.2 Brown and white adipose tissue . . . . . . . . . . . . . . . . . . . . . . . 13
3.3 Rat . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13
4 Results 14
4.1 Phantom . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 14
4.1.1 Number of averages . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 14
4.1.2 Number of frames . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 15
4.1.3 Size of the ROI . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16
4.1.4 Single vs. multislice . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16
4.1.5 Intervariability . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 17
4.1.6 Slice orientation dependencies . . . . . . . . . . . . . . . . . . . . . . . . 18
4.2 Mouse . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 19
4.2.1 Organs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 19
4.2.2 Brown and white adipose tissue . . . . . . . . . . . . . . . . . . . . . . . 19
4.3 Rat . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 19
5 Discussion 20
5.1 Sequence evaluation with phantom experiments . . . . . . . . . . . . . . . . . . 20
5.2 Relaxation measurements in mice . . . . . . . . . . . . . . . . . . . . . . . . . . 22
5.2.1 Relaxation measurements in organs and anatomical structures . . . . . . 22
5.2.2 Brown and white adipose tissue . . . . . . . . . . . . . . . . . . . . . . . 23
5.3 Relaxation measurements in rat brain . . . . . . . . . . . . . . . . . . . . . . . 23
Acknowledgement 25
References 26
Appendix 28
1 Introduction
The big advantage of MRI compared to other imaging techniques is the fact that it’s non-
invasive, opposite to x-ray no ionizing radiation is used. MR is a great tool for investigating
living systems, this means that animal studies can decrease the number of animals, no need
for dissection when a MR image could give the same information. MR has the ability to
make images of soft tissues, both structure and function. Images can be acquired with high
resolution and even in 3D. Another benefit is that the image slices can be placed in any
direction. [1]
Measurement of relaxation times in vivo is an important way to investigate properties of
tissues, one way to separate normal tissue from cancerous tissue. [2] The knowledge of T1 and
T2 relaxation times are necessary for optimizing contrast in imaging methods like functional
MRI and perfusion imaging. [3]
The aim of this study was to measure relaxation times (T1 and T2) of the organs in mice
and the brain in rats. We investigated as well if there are any differences in relaxation times
of white- and brown adipose tissue on mice.
2 MR theory
MRI, magnetic resonance imaging is derived from NMR, nuclear magnetic resonance, hence
the properties of the nucleus in the magnetic field is the importance. Nuclei used in MRI are
proton (1H), carbon (13C), and phosphorus (31P). The nuclei must have a non-homogenous
charge distribution.
All fundamental particles posses a property called spin, which can be explained as a ro-
tation of the nucleus around its own axis. Due to spin and the non-homogenous charge
distribution (i.e. a moving charge) the nucleus produces a magnetic moment.
MRI signal depends on the number of nucleus and their magnetic moment, therefore the
best choice (for regular imaging) is the proton, 1H, because of its high abundance and its high
(compared to other nucleus) magnetic moment. [4, 5, 6]
2.1 Behavior in an external magnetic field
The proton behaves like a small bar magnet when it’s placed in a static magnetic field, B0.
It aligns almost parallel to the direction of the field (i.e. the same way a compass aligns to
the magnetic field of the earth). Because of the spin and the laws of quantum mechanics
the behavior is a bit different from a compass needle. The proton does not align perfectly to
the field. Due to the torque it experience from the magnetic field it will precess around the
4
field axis, as shown in figure 1. The frequency of precession (known as the Larmor frequency)
can be derived from both classical and quantum mechanics. The Larmor frequency is given
in equation 1, where ω is the Larmor frequency, γ is the gyromagnetic ratio and B0 is the
amplitude of the Magnetic field (for derivation see appendix).
Figure 1: The proton aligns parallel or anti-parallel and precess in an magnetic field. Image retrievedfrom http://www.easymeasure.co.uk/ [7]
ω = 2πν = γ · B0 (1)
In an external magnetic field the numbers of possible values for the angular momentum are
2I +1. The 1H have a spin quantum number, I = 12 . That means it can have two orientations
(energy levels) in the field, spin up or spin down (i.e. parallel and anti-parallel to B0). These
two energy levels have slightly different energies. Population of the two levels is dependent
on both magnetic field and temperature. At 37 ◦C and 1.5 T there is a ratio of 1.000004, so
slightly more protons with spin up. The average of all proton magnetization will give a net
magnetization aligned exactly along the B0 field.
It’s possible to induce transitions between the energy levels. This can be achieved by
an additional oscillating magnetic field, B1 perpendicular to the static field, generated by a
radiofrequency (RF) pulse. The frequency of the RF pulse must match the energy difference
between the two states of the proton. The relation between energy and frequency, de Broglies
equation is given in equation 2.
∆ε = � · ω (2)
Where ∆ε is the energy difference, � is Planck’s constant divided by 2π and ω is the
associated frequency. [4, 5, 6]
5
2.2 Measuring the net magnetization
The net magnetization is flipped perpendicular to B0 field, this makes it possible to measure
the magnetization. Applying an RF pulse will make the magnetic moment of the protons to
change state, due to the higher number of spin-up protons more of them will change state and
the net magnetization will decrease. With the right strength and duration of the RF pulse the
net magnetization will be zero in the longitude view, this is called a 90◦ pulse. Fortunately
protons get into phase with each other during the RF pulse, otherwise the net magnetization
would be zero in the transverse plane too. So the RF pulse makes the net magnetization to
flip down from the z-axis and make the protons precess in phase.
Magnetization in the transverse plane will generate a current in the receiver coil (according
to Faradays law). Over time the signal decays while the net magnetization return to its
equilibrium state, as shown in figure 2. The induced signal is called the free induction decay
(FID). [4, 5, 6]
Figure 2: The transverse magnetization induces a current in the receiver coil and the signal is collected.And overtime the signal decays. Image retrieved from http://www.easymeasure.co.uk/ [7]
2.3 Relaxation
There are two mechanisms that make the net magnetization return to its equilibrium value,
spin-lattice relaxation, spin-spin relaxation.
The return of magnetization in longitude direction (for example after a 90◦ pulse) is called
spin-lattice-, longitudinal- or T1 relaxation. Spin-lattice relaxation is loss of energy from the
spinning nuclei to the surroundings (lattice). This relaxation is characterized by the time
constant, T1. T1 is defined as the time between a complete 90◦ pulse and the relaxation back
to 63 % (1-1/e) of its original value. This means that protons with different surroundings will
relax with different T1 times and give a contrast between different compositions. T1 has a
great dependency of the magnetic field strength. [4, 5, 6]
The return of the transverse magnetization is called spin-spin-, transverse- or T2 relaxation.
The name spin-spin comes from the exchange of energy between the nucleuses. Spin-spin
relaxation is loss of phase coherence, this is due to inhomogeneities of the magnetic field.
6
This inhomogenities descends from both the protons itself and other magnetic active nuclei
(molecular interactions) and from variations in the B0 field. The time constant, T2* (T2 star)
that characterize the spin-spin relaxation is defined as the time between a complete 90◦ pulse
and a decay to 37 % (1/e) of its original value. The time constant T2 have the same definition
as T2* aside from that it not includes the B0 variations. T2* is always shorter or equal T2.
[4, 5, 6, 8]
2.4 Inversion recovery technique
In the beginning of the pulse sequence the net magnetization is flipped by a 180◦ pulse. The
net magnetization will immediately start to recover and after a time TI (time of inversion) a
90 ◦ pulse is applied, see figure 3. IR (inversion recovery) is used to measure the T1 relaxation
time or to enhance contrast in T1-weighted images. Furthermore, it can also be used for fat
suppression, when the TI is equal the time of relaxation between the 180◦ and 90◦ for fat there
will not be any net magnetization to measure. [4, 6]
Figure 3: This image shows an inversion recovery sequence with spin echo. At first the RF pulse flipsthe net magnetization by 180◦, after a time TI the 90◦ pulse flips the net magnetization again. Thenext RF pulse is needed for the echo creation.
2.5 Gradient, slice selection and image encoding
In order to produce small variations in B0, gradient coils are implemented in a MRI system
(in three spatial dimensions). These variations are used to locate where the acquired signal
originates.
The slice selection gradient, Gz is a linear gradient applied orthogonal to the desired slice.
When the slice selection gradient is turned on, the effective magnetic field at the position z is:
Beff = B0 + Gzz (3)
7
And the resonance frequency is then:
νeff =γ
2π(B0 + Gzz) (4)
Thus the resonance frequency of the nuclei is dependent of the position along the z-axis.
Hence the slice thickness is determined by the bandwidth of the RF-pulse. The slice thickness
is given in equation 5. [4, 6]
∆z =2π∆ν
γGz(5)
Next gradient to be applied is the phase-encode gradient, Gx, which is oriented perpen-
dicular to the slice gradient. When the gradient is turned on, protons positioned where the
gradient strength is highest (i.e. at the highest magnetic field), will precess faster then the
protons on the other side and the protons at the middle precesses with the same frequency
as before. After the gradient is turned off all protons within the slice returned to the original
larmor frequency, but because of the gradient they have now shifted phases in different extent.
Protons in the high magnetic region is a little ahead the ones in the middle and protons in
the lower magnetic region are a little bit behind. This means they are positioned in columns
precessing with different phases. [5]
The final gradient is the frequency-encode gradient, Gy, (also known as the read gradient)
this one is applied perpendicular to both slice and phase gradient, to form a three dimensional
coordinate system. The gradient is turned on while the signal is collected at the receiver coil.
Now a single line in frequency domain (k-space) is collected. To collect all lines needed for
the image it’s necessary to start over again with a new phase-encode gradient followed by the
read gradient which is repeated until all lines are collected.
After acquiring of data of k-space in the frequency domain, a Fourier transformation is
performed to yield data in time domain. [5]
2.6 Echo creation
The signal collected in the receiver is not the real FID, as the gradients need time for switching,
thus an echo of the FID is acquired. The creation of this echo can be done in two ways, spin
echo or gradient echo.
2.6.1 Gradient echo
After the initial RF pulse a negative gradient is applied, this will dephase the transverse
magnetization. Next a positive gradient is applied and due to this the net magnetization will
8
rephase until maximum and then dephase again. The gradient echo does not compensate for
dephasing according to field inhomogeneities or spin-spin relaxation. In other word, the decay
of the echo is dependent on T2*.
2.6.2 Spin echo
The spin echo sequence starts with an RF pulse, making the net magnetization to flip 90◦.
Initially the spins are all in phase and are now allowed to dephase. After a time TE/2 (time
of echo divided by two) a 180◦ pulse is applied. This pulse will reverse the phase angle of the
spins, it will make the spins refocus back and form an echo (i.e. dephase backwards). The
echo will appear at a time equal the time between the initial RF pulse and the 180◦ pulse (i.e.
at the time TE). Field inhomogeneities have less effect in spin-echo, assuming the proton is
in the same position at all times of the measurement.
2.7 Relaxation measurements with IR-True-FISP
FISP is an acronym for Fast Imaging with Steady-state Precession. When all gradients used
in the sequence are rewound the sequence is called True-FISP, also known as fully rewound
gradient echo. Steady state imaging means that net magnetization reaches a steady state after
some RF pulses, i.e. the decrease of net magnetization due to excitation is the same as the
relaxation. When the sequence starts with an inversion recovery module we get IR-TrueFISP.
[6, 3]
IR-TrueFISP has the possibilities to measure both T1 and T2 in the same scan. The signal
intensity of an IR-True-FISP sequence is given in equation 6, where Sstst is the steady state
signal, for large inversion times, INV is the ratio of initial (S0) and steady state signal, Sstst.
T1* is the apparent T1 relaxation, T1* describes the recovery back to steady state. [3]
S(t) = Sstst ·(1 − INV · e(−t/T1∗)
)(6)
Schmitt et al. [9, 10] showed that the relaxation times can be calculated with equations
7 and 8. The T1* can be calculated by fitting equation 6 to a series of IR-TrueFISP images
with different TI times.
T1 = T1∗ cosα
2(INV − 1) (7)
T2 = T1∗ sin2 α
2
(1 − cos α
2
INV − 1
)−1
(8)
9
3 Materials and Methods
All measurements were done on a Bruker MR system, (Bruker BioSpin, Ettlingen, Germany)
located at Center for Mouse Physiology and Bio-Imaging (CPI), with a magnetic flux density
of 7 T. Acquisition of images was done with Paravision 4.0 (Bruker BioSpin). Slice thickness
of 1mm was selected to avoid partial volume artifacts. Minimum pixel size is dependent on
organ size.
Determination of SNR (signal to noise ratio) was done according to equation 9, where
Smean is the mean signal in the ROI (region of interest) placed on the measured object and
Nmean is mean signal of the ROI placed outside of the measured object.
SNR =Smean
Nmean(9)
Pulse scheme used for relaxation measurements were IR-TrueFisp (FISP-T1+T2-map),
with help of the macro IrTrueFispMap (supplied by Bruker). The macro generates a phase
corrected series of images. On the generated images a ROI is placed, with use of the ISA tool
(Image Sequence Analysis) a curve fit is performed and the values of T1 and T2 are given for
the selected ROI.
3.1 Phantom
For transmitting radiofrequency pulses and receiving MR data a volume coil, inner diameter
of 72 mm was used (Bruker BioSpin, Ettlingen, Germany). All measurements on the phantom
were done to optimize the measurement time, investigate the behavior of the sequence and to
learn how to handle the system. Scan time is preferred to be as short as possible in animals
due to anesthesia. A homemade phantom was used, built of five small plastic tubes containing
water with different concentrations of agarose and gadolinium, see figure 4. Scan repetition
time must be kept higher than 5 times T1, because the signal needs to recover. A couple of
parameters (i.e. averages, frames, ROI size, multiple slices, intervariability and slice selection)
were altered to see there effect on the measured value. The different variation schemes are
given below (3.1.1 to 3.1.6).
3.1.1 Number of averages
A sequence with TE: 1.5 ms TR: 3 ms Scan Repetition time: 10 s and matrix 128x128 was
run multiple times with different number of averages (1, 2, 3, 4 and 8). The scan was repeated
three times with different order (i.e sequential, reversed and mixed) there was also a delay
between the scans of 1 h, to avoid effects from heating up the phantom. T1 and T2 were
10
Figure 4: The phantom used for sequence evaluation.
measured for three different ROIs. More averages will lead to better SNR.
3.1.2 Number of frames
To test how many frames (i.e. number of points on the curve) needed for a sufficient curve fit
a sequence with TE: 1.5 ms, TR: 3.0 ms, scan repetition time: 10 s, number of averages: 4,
size of matrix: 128x128 was run three times with different number of frames (20, 40 and 60).
T1 and T2 were measured for three different ROIs.
3.1.3 Size of the ROI
To investigate if the size of the ROI has any influence in the result one image with TE: 1.5
ms, TR: 3.0 ms, scan repetition time: 10 s, number of averages: 1, size of matrix: 128x128
was acquired. T1 and T2 were measured 8 times with different size of the ROI (in cm2 0.0005,
0.0022, 0.0088, 0.0352, 0.1100, 0.27, 0.42, 0.64), see figure 5
3.1.4 Single vs. multislice
In the Bruker manual [3] it’s recommended to run only a single slice at the time. Testing how
much it affected the results a series of measurements were done with different number of slices
(1, 3, 4, 5, 20) and also with and without gap in between them. T1 and T2 were measured at
the same position and in the most middle slice.
3.1.5 Intervariability
To test if the results are stable over time a scan with TE: 1.5 ms, TR: 3.0 ms, scan repetition
time: 10 s, 4 averages and size of matrix 128x128 was repeated 10 times with 1 hour in between
the measurements. Relaxation times were measured at two positions.
11
Figure 5: Different sizes of ROIs were tested. The ROIs were placed centered to each other in thesame position of the phantom.
3.1.6 Slice orientation dependencies
To ensure that the results were not dependent on slice orientation a scan with TE: 1.5 ms, TR:
3.0 ms, scan repetition time: 10 s, 4 averages and size of matrix 128x128 was repeated 3 times
with different slice orientations (axial, sagittal and coronal). The phantom used was different
from the other measurements, this one contained water, gadolinium and sodium chloride in a
single compartment.
3.2 Mouse
The mouse strain used was C57/BL6. Initial measurements were done on dead mice, to learn
the anatomy. Relaxation measurements were done in vivo on living mice.
As the system could not be triggered in senseful mode, measurements of relaxation times
in organs which are moved during respiration and/or cardiac cycle could not be established.
3.2.1 Organs
First a multi slice multi echo scan was acquired for orientation purposes. Thereafter FISP
images (TE: 1.5 scan repetition time: 10 s, averages 4 and matrix 128x128) for each individual
organ (kidney, liver and front leg muscle) were acquired.
12
3.2.2 Brown and white adipose tissue
The aim of this measurement was to investigate if there are any differences in relaxation times
between brown and white adipose tissue. The differences should be used to discriminate fat
depots for questions within obesity research. The coil used was a 50 mm quadrature volume
coil (Rapid Biomedical, Rimpar, Germany), used for both transmit and receive. The mouse
were in supine position with head first. Brown adipose tissue is located in the back between
the scapulae, so supine position reduces motion artifacts.[11] A MSME (multi slice multi
echo) sequence were used to locate the fat deposit and then a 3 slice, 4 averages FISP image
with TE of 1.7 ms and a scan repetition time of 10 s, matrix size 64x96 and a resolution of
0.0235x0,0302 cm/pixel was acquired. Relaxation times were measured at different positions
within the fat deposit, see figure 6.
Figure 6: How the ROIs where placed for the differentiation of brown and white adipose tissue.
3.3 Rat
Aim of this experiment was to measure relaxation times of different regions of the rat brain,
i.e. (cerebellum, medulla oblongata, thalamus, hippocampus, corpus callosum and cerebral
cortex). Rat strain used was Sprague Dawley and all measurements were done in vivo. Coils
used were a 72 mm volume coil for transmission and a quadrature rat brain coil (Rapid Bio-
medical, Rimpar, Germany) for receiving. First a 20-slice, high resolution (in plane resolution
13
of 50x68 µm/pixel), T2-weighted, GE image was acquired for orientation purposes. FISP
images were acquired at appropriate positions with guidance from the high resolution image
and Paxinos and Watsons brain atlas [12].
4 Results
4.1 Phantom
4.1.1 Number of averages
Number of averages affect the SNR as well as the relaxation times as shown in Figure 7 for
T1 and in Figure 8 for T2. On T1 there is no big variation at medium (5.9 %) and high (3.8
%) SNR, in the low SNR an increase (10.2 %) of T1 time is seen after 2 averages.
There were a big variation of the T2 values at the low SNR (35.6 %). At the medium SNR
the variation were 12.7 % between the highest and lowest value. The high SNR value had a
variation of 10.7 %. The values of SNR is shown in Table 1. The increase of SNR between 1
and 8 averages were approximate 60 % for all ROIs.
Figure 7: T1 relaxation time as a function of number of averages.
Table 1: Signal to noise ratio are changing with number of averages.Averages Low SNR Medium SNR High SNR
1 1.4 4.1 5.22 1.8 5.7 7.13 2.2 7.8 94 2.5 8.6 10.68 3.2 11.0 13.4
14
Figure 8: T2 relaxation time as a function of number of averages.
4.1.2 Number of frames
Result of the number of frames are shown in figures 9 and 10. The T1 at high SNR is almost
a horizontal line, medium SNR is quite flat with a 11.3 % increase at 60 frames and the low
SNR have a big decrease (55.3 %) with an increased number of frames. T2 times at high and
medium SNR follow the same trend, an increase between 20 and 40 frames and a decrease
between 40 and 60 frames. The T2 values in the low SNR areas are increasing (28.7 %) with
number of frames.
Figure 9: T1-relaxation time as a function of number of frames.
15
Figure 10: T2-relaxation time as a function of number of frames.
4.1.3 Size of the ROI
The relaxation time as a function of ROI size is shown in figure 11. Some variations (3.9 %
for T1 and 12.2 % for T2) at the smallest ROIs and with larger ROIs a small decrease (8.7 %)
of the T1 times and an increase (17.4 %) of the T2 times.
Figure 11: T1 and T2 relaxation times as a fuction of ROI size.
4.1.4 Single vs. multislice
Results of measurements with different number of slices are shown in figures 12 and 13. The
high SNR T1 is decreasing with more slices and narrower gaps. Medium and low SNR show
a decrease between 1 and 5 slices with the large gap and then an increase with 5 slices (0.05
16
mm gap) and 4 and 20 without gaps. The T2 time is decreasing with more slices and narrow
gaps for the low SNR, while the high SNR have an increase of time between 1 and 5 slices
with the large gap and decrease with reduction of gap size or number of slices. No values for
the medium SNR were achieved due to too long relaxation time.
Figure 12: T1-relaxation time as a function of number of slices.
Figure 13: T2-relaxation time as a function of number of slices.
4.1.5 Intervariability
Results from the intervariability measurement are shown in figure 14. T1 for the first ROI
is stable within 4.7 % between scan number 1 and 7, between scan 1 and 10 the variation is
15.7 %. T1 for the second ROI is stable within 0.5 % between the first 7 scans and have a
variation of 12.2 % between scan 1 and 10. There was no times calculated for T2 on the first
ROI until scan number 9 and the variation between scan 9 and 10 are 9.1 %. T2 times for
the first ROI was stable within 1.3 % between scan 1 and 7 and between scan 1 and 10 the
17
variation was 37.6 %.
Figure 14: Intervariability of relaxation time measurement.
4.1.6 Slice orientation dependencies
Results from the slice orientation dependencies measurement are shown in figure 15. There
were only small variations between axial, sagittal and coronal slices in T1 (1.6 %). The T2
values between axial and sagital has a variation of only 0.4 % but between sagital and coronal
the variation were 9.7 %.
Figure 15: Slice orientations dependecies on relaxation times.
18
4.2 Mouse
4.2.1 Organs
The measured relaxation times for the organs in the mouse are given in Table 2. The shortest
T1 value is seen in the liver and the highest in the kidneys. The shortest T2 value is seen in
the front leg muscle and the highest in the liver.
Table 2: Relaxation times of organs in the mouse.Location T1 [ms] T2 [ms]Kidneys 1259 54Front leg muscle 1190 39Liver 985 60
4.2.2 Brown and white adipose tissue
T1 values for the different regions are shown in figure 16. As we see the lowest value are
recorded in the far back of the mouse, and increasing with more lateral location.
Figure 16: Relaxation times for the white and brown adipose tissue, locations of the ROIs can beseen in figure 6 in chapter 3.2.2.
4.3 Rat
The results from the relaxation measurements of the rat brain are given in Table 3. As can
be seen the relaxation times of thalamus and medulla oblongata are almost equal, difference
of T1 by 1.9 % and T2 by 0.15 %. Corpus callosum has the shortest T1 time, and the second
19
highest T2 time. Hippocampus and cerebral cortex have different values for the right and left
side.
Table 3: Relaxation times of several locations in the rat brain.Location T1 [ms] T2 [ms]Cerebellum 1343 80Corpus Callosum 1080 128Thalamus 1530 68Medulla oblongata 1560 68Cerebral cortex R/L 1246/1635 76/78Hippocampus R/L 1249/1072 171/254
5 Discussion
5.1 Sequence evaluation with phantom experiments
Careful evaluation of the pulse sequence is necessary to ensure valid results when used in
animals. Properties tested were the number of frames (i.e. data points for the fit curve),
size of ROI, number of averages, single vs. multislice acquisition, intervariability and slice
orientation dependencies.
Many data points are needed for a reliable curve fit, therefore a large number of frames is
recommended. A higher number of data points will lead to more exact values of the fit in case
the fit model is chosen correct for the data. In return, a wrong fit model or problems with
the measurements can be identified easier with a higher number of data points. Fewer frames
will give a higher standard deviation of the fit. As the test with varying the frames were all
done at the same total scan time, we decided for 60 frames, aiming for high accuracy. Initial
tests (data not shown) with a lower number of frames, thus less data points, provided still
acceptable curves for a fit. However, as a complete relaxation has to be reached in a FISP
sequence prior to acquisition of the next point in k-space, the shortage in the sequence did
not yield any advantage.
Organs and other anatomical structures can limit the size of ROIs, as they are not always
big enough for placing a sufficient large ROI. Due to small size of the ROI, the average and
the results might be influenced. There is no doubt that large ROIs, covering more tissue will
give an average with lower errors than small ROIs. The largest possible ROI that cover the
organ is therefore recommended, while taking care to average only uniform structures of the
organ. The results shows that ROIs larger then 0.27 cm2 are sufficient in size (i.e. no more
changes when increasing the area) and will provide valid results. If the ROI is smaller than
20
0.27 cm2, the results are prone to error. Mainly, low accuracy does not provide good data for
the fit and standard error of the fit is increasing dramatically. In the phantom the difference
between a ROI of size 0.035 cm2 and 0.42 cm2 were 5.8 % for the T1 and 9.1 % for the T2.
Organs in mice which are big enough to be covered by a ROI of sufficient size are liver, kidney,
muscles and fat. In principle, even the brain has a sufficient size to be covered by a single
voxel, however, the brain consist of several compartments, i.e. grey and white matter, CSF,
which are known to show different relaxation times. [13]
Variation of the signal to noise ratio, SNR (i.e. number of averages) did affect measured
results for the relaxation time, the largest variations were at the low SNR regions. Some images
indicated that there were susceptibility problems in the high-SNR ROI, probably caused by
an air bubble. The air bubble is seen as signal void in the upper compartment of the phantom
in Figure 5. This can explain the variation of relaxation times in our measurement. In theory
the increase of SNR should go by the square root of number of averages. Comparison between
measurement and theory are shown in Figure 17. Values achieved in the measurements were
aligned quite well to the theory. The standard value used in the sequence is four averages,
thus if the time was not critical in animal measurements 4 averages were used.
Figure 17: This figure shows a comparison between measured SNR and theoretical values.
Some early thoughts about the relaxation measurements were to cover the whole animal
with relaxation measurements and then pick out the values of the most interesting areas
afterwards. The resulting scan time was unacceptable long (>3 h) so complete body scans
could not be performed. Furthermore single vs. multislice measurements indicated that the
values were not identical when changing the number of slices and the gap in-between. This
could be an effect of inaccurate slice selections, and is in concert with Brukers recommendation
for single slice usage of FISP-T1-T2 protocol.
The results should not be time dependent, thus intervariability measurements were per-
21
formed. The results are constant until scan number 8. As the measurements were performed
automatically over night, something could have happened beyond operator control. Repetition
of the experiment was not possible due to time constraints.
Slice orientation should not affect the results. Almost all values were identical, and inde-
pendent of slice orientation. Remarkably, the coronal T2 value does not align to the others.
Because of the usage of a one-compartment phantom this is unlikely caused by susceptibility.
Earlier results (not shown) were dependent on slice selection, this was before a correction and
adjustment of eddy current compensation were performed. Eddy currents can induce field
which is different for each slice orientation. Corrections and tests of the eddy currents are of
great importance for reliable values in the in vivo situation.
5.2 Relaxation measurements in mice
Measurements on living animals are a lot different from measuring a phantom, as the animal
needs to be monitored and anesthetized. A living animal is moving, due to both cardiac
function and respiration and without triggering there might be some problems. To avoid
motion artefacts frequency- and phase encoding directions must be properly configured.
The protocol provided by Bruker has the possibility to use a triggered mode of the data
acquisition. Unfortunately, the trigger is setup such that an initiating event, e.g. heart beat
or respiration phase, is taken as the trigger point for the whole sequence. I.e. once the trigger
point is reached, the whole measurement is acquired and the motion is not frozen by a triggered
data acquisition. Besides the question what this type of trigger should be useful for, the T1-
T2 measurements could be reprogrammed as SSFP (steady state free precession) sequence.
We have placed this proposal at the manufacturer’s software department. Programming from
scratch would by far exceed the scope of this work.
5.2.1 Relaxation measurements in organs and anatomical structures
As mentioned above, the ROIs to be investigated have to have a certain size in order to get
reproducible values. In the mouse, where body weight is around 20 gram in an adult mouse,
organs are small and are likely to fall beyond the defined size of the ROI. In addition, the
inadequate trigger handling, as mentioned in chapter 5.2, does not allow meaningful data
acquisition of moving organs, i.e. the heart and the thorax, as well as structures next to them.
T1 increases in the order of liver, muscle and kidney and T2 values increases in the order
of muscle, kidney and liver. One might speculate where differences in relaxation times could
be originated. As specific weight of the different organs is nearly constant [14], cell density can
be ruled out. However, different levels of blood in the various organs could be one source for
22
the differences. [15] Blood, especially when oxygenated contains substantial amounts of Fe3+,
which is paramagnetic. Thus, relaxation times should be shorter with higher regional blood
content of the organs. As the order of T1 relaxation times aligns very well to the regional
blood volume, we assume that Fe3+ content contributes to the differences in the relaxation
times. However, other factors might influence the relaxation times as well.
5.2.2 Brown and white adipose tissue
Non-invasive in-vivo differentiation of brown and white adipose tissue would be helpful in
obesity research, a these fat pools act differently. [16, 17] Brown adipose tissue is supplied
with substantial blood flow, which causes the coloration of the fat. In contrast, white adipose
tissue has limited blood flow. As mentioned in 5.2.1, presence of oxygenated blood which
contains paramagnetic Fe3+, will shorten the relaxation times. During this work, we tested,
whether differences in T1 and T2 relaxation might be helpful to discriminate brown and
white adipose tissue. Motion artifacts due to respiration rule out a clear differentiation. As
stated in 5.2, a SSFP sequence could help to eliminate the problem. Currently, we can state
that differences of 12 % can be seen, however, the differences are not statistically significant.
Further work with an improved sequence should clarify the potential role of MRI in this field
of obesity research.
5.3 Relaxation measurements in rat brain
T2-weighted high resolution images of the rat brain were acquired. Thereafter, positions and
different structures in the rat brain could be assigned to the data of an anatomical brain
atlas. As the size of the ROI is a limiting factor, we could place ROIs only in medulla
oblongata, thalamus, cerebellum, cerebral cortex, hippocampus and corpus callosum, while
regional variations within the different locations could not be evaluated. The values for medulla
oblongata and thalamus are identical (1560/68 ms and 1530/68 ms). At the recent Annual
Scientific Meeting of the ISMRM (International Society for Magnetic Resonance in Medicine),
T1 values of marmoset at a field strength of 7 T were presented by Bock et al.. T1 values
in white and grey matter were 1350 ms and 1750 ms respectively. These values correlate
very well with our own results in the cerebellum (1343 ms). To our surprise, values measured
in the Corpus callosum (1080 ms) are lower, although both structures belong to the white
matter. A difference was seen for structures in the grey matter as well, as medulla oblongata,
thalamus and cerebral cortex show lower values. As brain structure and myelination differ
between species, differences in the observed values in comparison to marmoset could be based
on differences in morphology.
23
We have seen left-right variation in relaxation times of hippocampus and cerebral cortex.
The differences are in the order of 14 % and 24 % for hippocampus and cerebral cortex, respec-
tively and cannot be neglected. As we exclude a physiological background of the differences,
they might arise from differences in the phase information acquired during the FISP mea-
surements. Even with a reconstruction which reports magnitude values, which should exclude
the phase problem, the values obtained for the left and right hemisphere were not identical.
Field inhomogeneities are one possible reason for the inaccuracy of the phase information.
With non-uniform field strength over the rat brain, precession of the spins will be different
and relaxation times are likely to be influenced. These difficulties show the need for a critical
evaluation of the data acquired by the FISP sequence.
In our measurements, the ratio of the T1 relaxation times between grey and white matter
are approximately 1.15. The ratio is quite low in comparison to values for the human brain
at a field strength of 1.5 T [18], where the ratio is about 2. Thus, we can conclude that the
high field is not beneficial for the imaging contrast of grey and white matter. Discrimination
of grey and white matter is easier at lower field strength, however, the advantages of the high
field strength including higher resolution should not be forgotten.
24
Acknowledgement
I would like to give my deepest gratitude to my supervisor Michel Horn.
I will also thank Maria Enge for letting me borrow some of her fat mice, and to Erik Schéle
for the lesson and dissection of brown adipose tissue.
Measurements where done at Center for Mouse Physiology and Bio-Imaging (CPI), a core
facility of the University of Gothenburg.
25
References
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26
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27
APPENDIX
Derivation of the Larmor frequency, classical mechanics.
The protons magnetic moment is given in equation A-1, where �µ is the magnetic moment, γ
is the gyromagnetic ratio and �B is the magnetic field.
�µ = γ · �B (A-1)
If an external magnetic field is applied, the magnetic moment experiences a torque and pre-
cesses around the magnetic field axis.
∆µ = µ sin θ · ω∆t (A-2)
Where �µ = µ sin θ and small angles are assumed.
Angular momentum, equation of motion:
d�S
dt= �µ × �B (A-3)
Combine equation A-1 and A-3.
d�µ
dt= γ�µ× �B (A-4)
From equation A-4:
∆µ = γ∣∣∣�µ × �B
∣∣∣ · ∆t = γµ · B sin θ · ∆t (A-5)
Set equation A-2 equal to A-5
µ sin θ · ω∆t = γµ · B sin θ · ∆t (A-6)
We now have the Larmor equation:
ω = γ · B (A-7)
28
Derivation of the Larmor frequency, quantum mechanics.
The protons magnetic moment:
�µ = γ�I = γ · �S (A-8)
In an external magnetic field there is 2I +1 possible values for the angular momentum. Thus,1H have two possible enery levels ±1
2 .
Energy states:
ε = �µ · �B = γ · � · �I · �B (A-9)
Energy difference between the two states:
∆ε =[(
12
)−
(−1
2
)]γ · � · B = γ� · B (A-10)
De Broglie’s wave equation:
∆ε = � · ω (A-11)
Set equation A-10 equal to A-11:
γ�B = �ω (A-12)
And now we have the larmor equation:
ω = γ · B (A-13)
29