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Efficient Fourier transformation of unstructured meshes

and application to MRI simulation

by

Alejandro Martinez

A thesis submitted in conformity with the requirementsfor the degree of Master of Mechanical Engineering

Graduate Department of Mechanical and Industrial Engineering

University of Toronto

Copyright c© 2014 by Alejandro Martinez

Abstract

Efficient Fourier transformation of unstructured meshes and application to MRI

simulation

Alejandro Martinez

Master of Mechanical Engineering

Graduate Department of Mechanical and Industrial Engineering

University of Toronto

2014

This thesis demonstrates different ways to compute the MRI signal in finite element

magnetic resonance imaging or FE-MRI. For linear elements the MRI signal can be found

analytically. The error metric for linear elements to calculate the signal with a fixed error

in k-space using an adaptive Gaussian quadrature is derived using finite element function

derivatives and is then extended for use in quadratic elements. The feasibility of using

the numerical steepest descent (NSD) method, which is designed to compute oscillatory

integrals like the MRI signal equation, was also studied. The method was found to

compute the MRI signal faster than adaptive Gauss quadrature for quadratic surface

elements using typical MRI imaging parameters. The feasibility of NSD is discussed for

volumetric quadratic elements and for different mesh geometries.

ii

Dedication

For my love, Helena Violet Kelly, with gratitude for your love and support

iii

Acknowledgements

I would like to thank prof. Steinman. and prof. Huybrechs and the other members of

the examination committee prof. Nachman and prof. Mandelis for their support and

discussions. I would also like to thank my family for all their love and support.

iv

Contents

1 Introduction 1

1.1 Motivation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1

1.2 Introduction to MRI theory and simulation . . . . . . . . . . . . . . . . . 1

1.2.1 Signal generation . . . . . . . . . . . . . . . . . . . . . . . . . . . 3

1.2.2 Slice selection . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7

1.2.3 Spatial encoding . . . . . . . . . . . . . . . . . . . . . . . . . . . 8

1.2.4 Signal acquisition and reconstruction . . . . . . . . . . . . . . . . 10

1.2.5 Image resolution . . . . . . . . . . . . . . . . . . . . . . . . . . . 10

1.2.6 Image artefacts . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12

1.3 Literature review . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16

1.3.1 Analytic solutions . . . . . . . . . . . . . . . . . . . . . . . . . . . 17

1.3.2 Simulators using magnetization maps . . . . . . . . . . . . . . . . 20

1.3.3 Simulators using T1, T2 and proton density maps . . . . . . . . . 20

1.3.4 Finite element based simulators . . . . . . . . . . . . . . . . . . . 21

1.3.5 Bloch equation simulators . . . . . . . . . . . . . . . . . . . . . . 23

1.4 MRI acquisition within the FE-MRI library . . . . . . . . . . . . . . . . 23

1.5 Research objectives . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 23

2 Adaptive Gauss quadrature for FE-MRI elements 26

2.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 26

v

2.2 Gauss quadrature to compute oscillatory integrals . . . . . . . . . . . . . 29

2.3 Computing oscillatory integrals in higher dimensions . . . . . . . . . . . 30

2.4 Error estimation for tri3 elements . . . . . . . . . . . . . . . . . . . . . . 31

2.5 Error estimation for tri6 elements . . . . . . . . . . . . . . . . . . . . . . 36

2.6 Error estimation for tet4 elements . . . . . . . . . . . . . . . . . . . . . . 38

2.7 Error estimation for tet10 elements . . . . . . . . . . . . . . . . . . . . . 40

3 Algorithm details and implementation 41

3.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 41

3.2 FE-MRI software architecture . . . . . . . . . . . . . . . . . . . . . . . . 42

3.3 K-space acquisition algorithm . . . . . . . . . . . . . . . . . . . . . . . . 43

3.4 Cycles per element quadrature table . . . . . . . . . . . . . . . . . . . . . 46

3.5 Calculation of quadrature rules . . . . . . . . . . . . . . . . . . . . . . . 48

3.6 Slice profile simulation for volumetric meshes . . . . . . . . . . . . . . . . 48

3.7 Additions and improvements to current FE-MRI functions . . . . . . . . 49

4 Results and discussion 52

4.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 52

4.2 Validation of optimal algorithm for tri3 elements . . . . . . . . . . . . . . 52

4.3 Validation of optimal algorithm for tri6 elements . . . . . . . . . . . . . . 56

4.4 Validation of optimal algorithm for tet4 elements . . . . . . . . . . . . . 56

4.5 Validation of optimal algorithm for tet10 elements . . . . . . . . . . . . . 58

4.6 Computation time comparison between linear and quadratic surface dis-

cretizations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 59

4.7 Simulation of MRI acquisition of a carotid artery . . . . . . . . . . . . . 61

5 Feasibility of NSD for faster signal calculation 62

5.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 62

5.1.1 Approximate evaluation of oscillatory integrals . . . . . . . . . . . 62

vi

5.1.2 Adaptive quadrature . . . . . . . . . . . . . . . . . . . . . . . . . 63

5.1.3 Filon method . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 63

5.1.4 Levin method . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 64

5.1.5 Numerical steepest descent . . . . . . . . . . . . . . . . . . . . . . 65

5.2 Experiment using NSD for quadratic surface elements . . . . . . . . . . . 66

5.2.1 NSD decomposition for quadratic surface elements . . . . . . . . . 68

5.2.2 Algorithm details . . . . . . . . . . . . . . . . . . . . . . . . . . . 71

5.3 Results and discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . 72

6 Conclusions 75

6.1 Summary of work . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 75

6.2 Future directions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 75

A Generation of Gauss quadratures 77

A.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 77

A.2 The Golub-Welsch algorithm . . . . . . . . . . . . . . . . . . . . . . . . . 82

A.3 Sample C++ quadrature code . . . . . . . . . . . . . . . . . . . . . . . . 83

B Mesh generation with same discretization error 88

B.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 88

B.2 Discretization error when using quadratic elements . . . . . . . . . . . . 89

B.3 Discretization error when using linear elements . . . . . . . . . . . . . . . 91

B.4 MATLAB implementation . . . . . . . . . . . . . . . . . . . . . . . . . . 92

C Implementation of NSD algorithm 97

C.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 97

C.2 G1 decomposition . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 99

C.3 Steepest descent paths for G . . . . . . . . . . . . . . . . . . . . . . . . . 105

C.4 Steepest descent paths for F . . . . . . . . . . . . . . . . . . . . . . . . . 106

vii

Bibliography 108

viii

List of Tables

1.1 Values used for fig. 1.1 . . . . . . . . . . . . . . . . . . . . . . . . . . . . 17

1.2 Finite element types used in this thesis . . . . . . . . . . . . . . . . . . . 25

4.1 K-space error error maps and histograms for all the element types imple-

mented. Notice that the x-axis for the histograms in in terms of n, where

error = 10n . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 54

4.2 Test meshes for each element type and their simulated MRI images. For

the meshes with quadratic elements one of the elements is selected to show

the blue edges shown are between element nodes and do not delineate a

quadratic element . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 55

5.1 Total computation times in seconds and time ratios for different k-space

extents for an error threshold of 10−2 . . . . . . . . . . . . . . . . . . . . 72

5.2 Total computation times in seconds and time ratios for different k-space

extents for an error threshold of 10−3 . . . . . . . . . . . . . . . . . . . . 72

A.1 Fejer quadrature points and weighs . . . . . . . . . . . . . . . . . . . . . 81

ix

List of Figures

1.1 Measured (top) and simulated (bottom) magnetization and T2 maps of an

artery wall . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 18

2.1 Transformation of a plane wave in local coordinates in the standard trian-

gle to the unit square . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 32

2.2 tri3 domain s and vertices in global and local coordinate systems with a

given K space vector . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 35

2.3 The plane formed by the gradient of g with respect to ξ1 and the location

of its possible maximum absolute values in domain s. In this case the

maximum absolute value of the gradient is at ξ1 = ξ2 = 0 . . . . . . . . . 37

3.1 Flowchart of the main FE-MRI algorithm . . . . . . . . . . . . . . . . . 44

4.1 K-space error as a function of k-space location for a tet4 unit cube with a

quadrature order of 14 . . . . . . . . . . . . . . . . . . . . . . . . . . . . 57

4.2 Number of k-space points as a function of n where error = 10n for a tet4

unit cube with a quadrature order of 14 . . . . . . . . . . . . . . . . . . . 58

4.3 Mesh of segmented artery (a) and its MRI reconstruction (b) . . . . . . . 61

5.1 K-space computation time maps in milliseconds for different adaptive al-

gorithms and error thresholds. The time ratios between the algorithms is

also shown . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 74

x

B.1 Discretization of a circle using quadratic curve segments. The quantity to

be minimized is proportional to the area between the lines . . . . . . . . 89

B.2 Discretization of a circle using line segments. The quantity to be mini-

mized is proportional to the area between the lines . . . . . . . . . . . . 91

xi

Chapter 1

Introduction

1.1 Motivation

MRI simulators are primarily used to determine the origin of acquisition errors and to

design and optimize new pulse sequences. For example in [1], an MRI simulator is used

to design and optimize a MRI pulse sequence for the characterization of artherosclerotic

plaque. They are also used for education and training [2]. The motivation of this thesis

is to design an adaptive Gauss quadrature to evaluate the MRI signal equation for un-

structured meshes and to design and test an algorithm using numerical steepest descent

(NSD) [3], [4], [5], [6]. Using NSD one can compute the Fourier transform of a mesh

faster than other approaches like numeric quadrature [7]. The algorithm was tested us-

ing parameters needed for simulating an MRI acquisition. These results are then used

to implement and improve upon new functions in the FE-MRI library [7] which can be

used to simulate MRI acquisition.

1.2 Introduction to MRI theory and simulation

In this section, an introduction to magnetic resonance imaging (MRI) will be given with

emphasis on signal acquisition, its Fourier transform approximation (k-space) and image

1

Chapter 1. Introduction 2

artefacts. The introduction will have an emphasis on imaging and reconstruction methods

that are based on the Fourier transform and for 2D imaging. This introduction will also

aid the reader in understanding how in some cases an MRI acquisition of an object can be

simulated by the Fourier transform and inversion of a set of points or elements weighted

by the actual material properties of said element such as in the FE-MRI library.

The MRI acquisition process can be separated into the following steps:

1. Signal generation: generating a net magnetization in an object

2. Slice selection: selectively exciting a plane with a given thickness if a 2D acquisition

is desired

3. Spacial encoding: spatially encoding the signal in that plane

4. Signal acquisition and reconstruction: acquiring a radio frequency signal and re-

constructing the image

These steps will be described in detail in the subsequent subsections. This introduc-

tion is based on the following sources: [8], [9], [10]. An overview is given next. MRI

works by aligning nuclei in an object along their spin axes more in one direction than

another. This spin is a magnetic and quantum effect but it can be described in the same

way as a spinning top. This is done by a very strong fixed magnetic field B0. Then

another moving magnetic field that is perpendicular to the fixed field called B1 tips the

spins and they precess around the original fixed field. The amount of tipping depends on

how long the B1 field is applied. The speed of precession is dependant on the strength

of B0. The B1 field can be chosen in a way that allows for selective excitation of a slice

perpendicular to the z axis. Another smaller fixed gradient field Gx (i.e B = B0 +Gxx)

is superimposed on B0 to cause a gradient in the total fixed field. The precessing spins

emit radio waves that can be picked up by a coil. The signal’s frequency and magnitude

are related to the spatial location of the spins along the x axis and the amount of spins


in that given location. Hence an image of the spin distributions can be reconstructed by

taking the inverse Fourier transform of the signal. To obtain a 2D distribution of spins,

the locations along the y-axis can be encoded by using a gradient field Gy applied before

the measurement is taking place, so that the phase of each spin depends on the location

along the y-axis. The readout coils which do the measurement can measure the phase,

magnitude and frequency of the signal which means the 2D distribution of spins in a slice

can be determined.

1.2.1 Signal generation

The coordinate axes that will be used in this and subsequent sections are only used for

convenience and consistency with MRI literature. Some nuclei in living things, like in

hydrogen atoms (1H), can have nuclear spin angular momentum. These types of nuclei

will be referred to as spins. The most common spin used for MRI is 1H since it is a

component of water and the human body is mostly water. Spins can be described as

small bar magnets oriented randomly in a given material. When the spins are exposed

to a static magnetic field, some spins will align with the magnetic field and others align

in the opposite direction to it and in both orientations they will rotate around the z axis

with an angular frequency given by the Larmor equation:

ω = γB0 (1.1)

Where γ is the gyromagnetic ratio which for a particle is the ratio between its magnetic

dipole moment and its angular momentum. For 1H, γ/2π = 42.48 MHz/T [10]. A group

of spins in an object with the same γ value are commonly referred to as isochromats. B0

is the applied magnetic field referred to as the static field and it points in the positive

z-axis. Slightly more spins will align in the same direction as B0 because a spin is more


likely to take the lower energy state than a higher energy state. The net magnetization

is simply the spins oriented towards the positive z-axis minus the ones oriented towards

the negative z-axis and it is given by the following equation:

M0z = | ~M | =γ2~2B0Ns4KTs

(1.2)

Where ~ is Plank’s constant divided by 2π, Ns is the number of spins, K is Boltzmann

constant and Ts is absolute temperature of the spins. The derivation of equation 1.2

is similar to the one used in [11] to determine the properties of a simple paramagnetic

solid. More information is given in [10] and [11]. Equation 1.2 only applies to so called

spin-1/2 systems like 1H isochromats. The spins also can be made to rotate with other

magnetic fields with respect to any line in the x-y plane but only by discrete angles;

it is a quantum mechanical effect. However since the net magnetization is a sum of

billions of spins in the whole object, the angle of rotation for the magnetization can be

assumed to be continuous and a classical description analogous to a precessing spinning

top can be used. After applying the static field, the net magnetization will point in the

same direction. The equation that describes the time-dependent behaviour of ~M in the

presence of an applied magnetic field is called the Bloch equation and it is given by:

d ~M

dt= γ ~M × ~B − Mxî−My ĵ

T2− Mz î−M0k̂

T1(1.3)

Where ~B is the total applied magnetic field, Mx−z are the components of ~M and T1 and

T2 are the longitudinal and transverse relaxation times respectively. The range of values

for the relaxation times for 1H are T1 ∈ [100, 1500] ms and T1 ∈ [100, 1500] ms [10]. Let

us introduce a transverse magnetic field that oscillates at the Larmor frequency:


~B1(t) = 2Be1(t)cos(ωt+ ϕ)̂i (1.4)

Where Be1(t) is called an envelope function that for example can have a value of B1 from

time 0 to some time τp and 0 otherwise. τp is simply the amount of time the B1 field is

applied. It is also referred to as the pulse width. It will be shown later in this Section

that the useful factor is only the integral of Be1(t) from 0 to τp and not its shape. The

value ϕ is an initial phase angle. This longitudinal magnetic field can be decomposed

into fields rotating in opposite directions about the z-axis. One of the terms rotates in

the opposite direction as the processing spins and has a negligible effect. Therefore the

effective field that needs to be considered is:

~B1(t) = Be1(t)[cos(ωt+ ϕ)̂i− sin(ωt+ ϕ)ĵ] (1.5)

The magnetization vector 1.5 can be transformed into a vector in the Larmor-rotating

frame with the following coordinate transformations:

î′ ≡ cos(ωt)̂i− sin(ωt)ĵ

ĵ′ ≡ sin(ωt)̂i+ cos(ωt)ĵ

k̂′ ≡ k̂ (1.6)

Using the transformation 1.6, the magnetic field 1.5 becomes:

~B1,rot(t) = Be1(t)î

′ (1.7)


Which is expected since the vector ~i′+~j′ is spinning around the z-axis in the same direction

and angular speed as the original ~B1(t) vector. If the pulse width is small compared to

both relaxation times, the effect of ~B1(t) on the magnetization ~M is described by the first

term of the Bloch equation. As will be shown below, it is also advantageous to transform

that term into the Larmor-rotating frame. The transformed Bloch equation for a short

exitation using the transformation 1.6 is:

∂ ~Mrot∂t

= γ ~Mrot × ~Beff (1.8)

Where ~Beff = ~Brot + ~ω/γ is referred to as the effective magnetic field. Since the total

rotating field is B0~k′ + Be1(t)î

′, the motion of the bulk magnetization vector will be

described by:

∂ ~M

∂t= γ ~Mrot × ~Be1(t) ~i′ (1.9)

Notice that equation 1.9 will be equal to zero (i.e the bulk magnetization does not precess)

if the only field present is the original B0~k′ magnetization. The solution to equation 1.9

up to the pulse width time τp with the initial condition ~M(0) = M0z is the following:

Mx′(t) = 0 (1.10)

My′(t) = M0z sin(α) (1.11)

Mz′(t) = M0z cos(α) (1.12)


Where α is referred to as the flip angle and its given by:

α =

∫ t

0

γBe1(u)du (1.13)

If the envelope function is rectangular then α = γB1t. So in this case the bulk magneti-

zation tips away from the z′ axis about x′ or precesses (since its tipping but also rotating

about z in the laboratory frame) at an angular velocity of −γB1. Also notice that as

long as the area under B1(t) is the same, the bulk magnetization will tip to the same

final angle.

1.2.2 Slice selection

Since the envelope function Be1(t) can have any shape and still produce the desired tip

angle as long as its area is constant, it can be used to selectively excite a slice of the object

being imaged. For simplicity we will assume the slice is perpendicular to the z axis. If

we apply a gradient to the B0 field as a function of the z component of location while

the spins are being tipped, their positions will be spatially encoded by their resonant

frequencies with the following frequency selection function;

p(f) =

1, |f − fc| < ∆f

0, otherwise

(1.14)

Where:

fc = f0 +γ

2πGzz0 (1.15)


∆f =γ

2πGz∆z (1.16)

To obtain the time signal needed to produce the corresponding frequency spectrum, the

Fourier transform can be applied on equation 1.14:

B1(t) ∝∫

∞

−∞

p(f)e−2πift dt = ∆fsinc(π∆ft)e−2πifct (1.17)

From equation 1.18, we get the envelope function:

Be1(t) = Asinc(π∆ft) (1.18)

Where the constant A depends on the desired spin tip angle. This pulse is not possible

to apply in practice so it is truncated from time 0 to some time τp, with the center of the

sinc function at τp/2.

1.2.3 Spatial encoding

One can use the static field B0 to encode spatial information in the final MRI signal. A

simple way of achieving this is with a constant gradient field such as:

ω(x) = ω0 + γGxx (1.19)

So now the field B0 pointing along the z-axis changes magnitude along x. After tipping

the spins uniformly along a slice in the z-axis to 90◦ for example, this gradient can be

used to spatially encode a signal using both phase and frequency. This is needed because


the image signal in 2D acquisition is:

s(t) =

∫∫

m(x, y)e−iγ(Gxx+Gyy)tdxdy (1.20)

Which is the Fourier transform of the object’s magnetization, m(x, y), on the particular

slice being imaged:

M(kx, ky) =

∫∫

m(x, y)e−2πi(kxx+kyy)dxdy (1.21)

Where the Fourier coordinates for general G gradients are given by:

kx =γ

2π

∫ t

0

Gx(u) du (1.22)

ky =γ

2π

∫ t

0

Gy(u) du (1.23)

In the MRI literature Fourier space is referred to as k-space. Phase encoding can be

achieved as follows: first tip all spins 90◦ in the same direction using the B1(t) field

discussed in Section 1.2.1, then apply a gradient field which happens along a line in

the x-y plane (the magnetic field is still pointing in the z-axis) for a small time Tpe

which stands for phase encoding time. After Tpe seconds, the spins will have a phase

proportional to their location. For example if the G gradient that happens along the

y-axis and is constant then the phase of the signal as a function of y is:

φ(y) = −γGyyTpe (1.24)


If the signal is acquired right after Tpe for a very short time interval, we would obtain

the sum of the magnetizations for each line that contains the same phase. However, if we

apply yet another gradient field such as Gx while recording the signal every ∆t seconds,

we will also obtain the magnetization along those lines as a function of the frequency,

i.e the spatial information along the line has been frequency encoded. This procedure is

referred to as a pulse sequence. The sequence described above is referred to as spin-wrap

imaging and it is one of the most commonly used sequences because it results in equally

spaced points in k-space on a grid which can be easily transformed into an image using

the fast Fourier transform.

1.2.4 Signal acquisition and reconstruction

Equation 1.20 for the MRI signal is obtained from radio waves given off by the object

being imaged which are measured by coils in the MRI machine. The signal from the

coils is then converted to a digital signal and its phase and frequency components are

determined. If the signal is sampled in k-space at uniform intervals such as in spin-warp

imaging, the signal can be obtained applying the inverse discrete Fourier transform of

the discrete version of the signal equation. This method can take O(N2) operations.

The same transform can be performed in O(Nlog2N) operations using the fast Fourier

transform or FFT. FFT is the reconstruction method most likely used in MRI equipment.

Other transforms are available for k-space acquisition that is sampled radially such as

the Radon transform. For more information refer to [9].

1.2.5 Image resolution

When the MRI signal is being acquired, k-space is recorded at discrete time intervals.

Using equations 1.23 and 1.22 the k-space spacing is:


∆ky =γ

2π∆Gyty (1.25)

∆kx =γ

2πGx,max∆t (1.26)

Because of the Fourier sampling theorem [9], the field of view of the image should be the

following:

FOVx =1

∆Kx(1.27)

Which also applies for the y and z axes. The spatial resolution of the image in the x axis

is the field of view divided by the number of readout samples:

∆x =FOVxNread

=1

∆kxNread(1.28)

While the spatial resolution in the y axis is the field of view divided by the number of

phase encoding steps:

∆y =FOVyNpe

=1

∆kyNpe(1.29)

Notice the above relations apply only to spin-warp imaging. They do not apply to other

pulse sequences that sample k-space non-uniformly.


1.2.6 Image artefacts

Image artefacts arise in MRI acquisition either due to insufficient or inaccurate data or

both. The motivation of this thesis is to provide a tool to simulate MRI such that the

cause of image errors can be understood and also to design and test protocols such as the

resolution and how it affects the measurement accuracy of some anatomic feature similar

to [12] for example. With this in mind, the following image errors are described together

with their causes and how they can be simulated in general or in the FE-MRI library.

Gibbs or truncation artefact

The Gibbs or truncation artefact is seen as spurious ringing around sharp discontinuities

in the reconstructed image. If the image is a smooth function, the inverse transform of

the signal equation uniformly converges to the true image when the number of samples

N → ∞ for a given FOV . However if the function is discontinuous, the inverse transform

will have ringing with an overshoot and an undershoot of about 9% of the difference at

the discontinuity. Surprisingly the overshoot and undershoot percent value is constant

regardless of N . However the period of the ringing will decrease with higher N values.

If the signal was not truncated to a finite number of N samples then the ringing period

would be zero and the exact image would be reconstructed hence the name truncation

artefact. This artefact can be made to appear on any simulated MRI acquisition simply

by using a low enough number of N samples.

Aliasing

Aliasing happens when k-space is under-sampled. The artefact is a wraparound of the

image in the direction that is being under-sampled. This can be clearly understood by

the relationship between the true image function and its Fourier series reconstruction [9]:


Î =∞∑

n=−∞

I

(

x− nx∆kx

)

(1.30)

Where the term 1/∆kx = FOVx. This error can be readily be simulated by simply

using an FOV value that is smaller than the object size along the same direction. This

error can be eliminated simply by choosing an FOV value larger than the object being

imaged.

Chemical shift artefacts

Chemical shift artefacts happen when an object with different isochromats is being im-

aged. Different isochromats will have a different γ value. Since spatial position is fre-

quency encoded in the readout direction when an MRI signal is being obtained, signals

coming from two different isochromats in the same location such as water and fat will

end up in different locations along the frequency encode signal. The artefact will look

like a region of the object such as fat which is not being explicitly imaged will be shifted

compared to the rest of the image features. The frequency shift due to the different

isochromat is given by [9]:

∆ωc = γδB0 (1.31)

The frequency bandwidth of a pixel in k-space is:

∆ωx = γG∆x (1.32)

Since the pixel bandwidth is proportional to the pixel length, the spatial displacement

in units of a pixel is:


δx =∆ωc∆ωx

=δB0G∆x

(1.33)

And the spatial shift will be:

∆xc = δx∆x =δB0G

(1.34)

Therefore an effective way to reduce the chemical shift is to use a strong readout

gradient. In practise this will also reduce the signal to noise ratio of the acquired signal.

Therefore an optimal value for G has to be determined. This error can be simulated

simply by obtaining the k-space or each isochromat separately and then adding the

signal.

Motion artefacts

A type of very obvious motion artefact that happens when the whole object being scanned

is in motion is called ‘ghosting’. It can be clearly identified by copies of the object being

imaged along the direction of motion and with a lower contrast than the actual image.

This error happens when the object moves along the phase-encoding direction. For

example, assuming the object moves at a constant speed vx along the x-direction and

there is a constant phase encoding gradient Gx applied for some time interval τ starting

from time tn, the phase shift is the following:

φ = γ

∫ tn+τ

tn

Gn(x+ vxt) dt

= γGnxτ + γGnvxτ(tn + τ/2)

= 2πknx+ 2πknvx(tn + τ/2) (1.35)


So the motion induced phase shift is:

∆φm = 2πknvxtn + 2πknvxτ/2 (1.36)

Since the time τ is much smaller than the total acquisition time up to that point, the

second term of equation 1.36 can in most cases be omitted. Equation 1.36 without the

second term can be obtained directly by applying the Fourier shift theorem to the MRI

signal equation. For constant vx, Gx and a negligible τ , the signal equation is:

sm(k, t) = e−2πiknvxts(k, t) (1.37)

Equation 1.37 has the same phase shift as the one given in equation 1.36. To simulate

this error, either equation 1.37 can be applied directly to the simulated k-space map for

a constant speed vx or the object can be translated a certain distance based on vx in the

x direction and then its corresponding k-space points calculated to simulate the error. A

further discussion on other types of motion artefacts is given in [9].

Flow artefacts

Flow artefacts are caused by the motion of fluids such as blood in the living tissue being

imaged. The main feature of these errors is a loss or gain of signal intensity where the

flow is taking place like inside an artery [13]. The only simulation method that has been

used successfully for these types of error is to either measure the flow directly using other

imaging methods or to simulate the flow using computational fluid dynamics (CFD).

Then a certain number of streamlines are extracted from the flow field. Afterwards the

individual effect of spins moving along the streamlines is computed to obtain an MRI


signal. References [13] and [14] use this method of simulation. Since a set of discrete

spins are used to simulate a continuum fluid are used, the number of spins needed and

the related computation time is high.

1.3 Literature review

The different MRI simulation methods can be grouped into the following types:

1. Analytic solutions

2. Simulators using magnetization maps

3. Simulators using T1, T2 and proton density maps

4. Finite element simulators

5. Bloch equation simulators

The order on the list is based on increasing physical accuracy of the models in most

cases at the expense of computation time. Except for some Bloch equation simulators, the

other types use the Fourier transform of a pre-described signal therefore they technically

are not simulators in the sense that they are not simulating a full MRI machine. A full

simulation would require modelling both the object using the Bloch equation and the

MRI components such as the coils and the B0 magnet. However the methods described

below will be referred to as simulations since they do still crudely simulate the object

and parts of the MRI machine. Each type is described in the following sections.


1.3.1 Analytic solutions

In some situations, an analytic solution for k-space can be found. The author co-authored

a paper, [12], where MRI acquisition of an artery was performed to determine how the

total arterial wall thickness obtained from the segmentation of the MRI image depends

on the T2 values of the arterial layers (intima, media and adventitia) and TE . To test

if this is due to a partial volume effect (a low resolution) or because of the different

T2 values of each layer, an analytic k-space was constructed for each of the 4 acquired

images by adding the k-spaces of 4 concentric circles. Then the obtained images were

segmented. The values used are given in the following table:

Layer ro(mm) pTE=17ms pTE=34ms pTE=51ms pTE=68ms

arterial blood 3.7 0.0000 0.0000 0.0000 0.0000

intima 3.82 1.0242 0.8800 0.7561 0.6496

media 4.35 1.0000 0.7699 0.5927 0.4563adventitia 4.42 0.7063 0.3539 0.1773 0.0888

background N/A 0.0000 0.0000 0.0000 0.0000

Table 1.1: Values used for fig. 1.1

Where ro and p in Table 1.1 stands the outer radius for the image intensity. The image

intensity is in arbitrary units; the actual value depends on the specific MRI scanner. The

thickness changes as a function of TE closely match the experimental results. TE is called

the echo time and is simply the time where the spins are allowed to relax after a RF

filed is applied. Then a mono-exponential fit was performed (p = e−TE/T2 , where p is the

image intensity at a pixel) on the 4 simulated images to get a T2 map; it also closely

matches the T2 map obtained from the actual MRI images. For the mono-exponential

fit, if a pixel had p values lower than 1.5 × 10−5 on all the 4 images, it was not taken

into consideration and was given an arbitrary value of 0. The measured and simulated

magnetization and T2 maps are shown in Figure 1.1.

The magnetization image shown has a magnetization normalized to 1. This image

weighted with different magnetization values based on the echo time and the T2 decay are


Image x−axis (mm)

Image y

−axis

(m

m)

−6 −4 −2 0 2 4 6

−6

−4

−2

0

2

4

6

Image x−axis (mm)

Image y

−axis

(m

m)

−6 −4 −2 0 2 4 6

−6

−4

−2

0

2

4

6

T2 (m

s)

50

100

150

200

250

Figure 1.1: Measured (top) and simulated (bottom) magnetization and T2 maps of anartery wall

used in the paper. To obtain the k-space, corresponding to a simplified artery, 4 circle k-

spaces were added, this can be done since the Fourier transform is a linear transform [9].

If the artery is long compared to its radius, a 2D Fourier transform can be used instead

of the Fourier transform for a cylinder. If image and k-space coordinates are changed to

cylindrical coordinates: r =√

(∆x)2 + (∆y)2, θ = tan−1(∆x/∆y), kr =√

k2x + k2y and

φ = tan−1(kx/ky). Therefore:

Kxx+Kyy = krr(cosφcosθ + sinφsinθ) = krrcos(θ − φ) (1.38)


Then the Fourier transform of a circle with unit radius and height is given in cylin-

drical coordinates as follows:

Fcirclesimple(kr) =

∫ 1

0

r

[∫ 2π

0

e2πikrrcos(θ−φ)dθ

]

dr (1.39)

The integral can be simplified in cylindrical coordinates since the function for a cycle

is circularly symmetric. The solution of equation 1.39 is given in terms of the first order

Bessel function of the first kind J1 [9] and its the following:

Fcirclesimple(kr) =J1(2πkr)

kr(1.40)

In the case of kr = 0, a Taylor series can be taken of equation 1.40 about kr = 0, and

simplified to yield:

Fcirclesimple(0) = π (1.41)

Applying the scaling and linearity properties of the Fourier transform [9] to equations

1.40 and 1.41 yield the following equations for an arbitrary image intensity (h) and radius

(r):

Fcircle(kr) = hrJ1(2πrkr)

kr(1.42)

Fcircle(0) = hr2π (1.43)


This k-space is implemented in the FE-MRI library as a python script. An example

of its use in Unix is given below:

femri2dcylinderkspace -fov 12.0 12.0 1.0 -fovcenter 0.0 0.0 0.0

-matrix 32 32 1 -mvalue 1.0 -slicethickness 0.0 -radius 3.0

-center 0.0 0.0 0.0 --pipe femrikspacetoimage -dimensionality 2

-ofile circleImage.vti

Both in the magnetization maps derived from measurements and the analytic model

it was found that the mean wall thickness obtained by segmentation decreases with echo

time because of the small T2 time of the adventitia. Therefore a short echo time should be

included to properly image it. This layer plays an important role in plaque development

[12].

1.3.2 Simulators using magnetization maps

In [15], artefacts due to relaxation, linear flow and motion after an RF-pulse are added

to the k-space of the magnetization also called proton density map. Some limitations

of this method are not able to simulate errors due to non-linear gradients and motion

during and RF-pulse, since they are considered to be momentary.

1.3.3 Simulators using T1, T2 and proton density maps

In [2], the T1, T2 and proton density maps are Fourier transformed and changed according

to limited properties of MRI pulse sequences. Although the method is computationally

fast and can simulate wraparound and simple motion artefacts accurately, by design it

cannot simulate errors related to signal generation. This type of simulation is used in [1]

to optimize a MRI pulse sequence for the characterization of atherosclerotic plaque.


1.3.4 Finite element based simulators

Finite element based simulators also use a magnetization map and can include T1, T2

maps, however they use the finite element discretization which allows the magnetization

to vary continuously within the object. The elements used for this thesis are shown

in Table 1.2. The names given in the figure will be used in the rest of the thesis.

Both types of simulators have to transform the magnetization map into the Fourier

domain. The Fourier transformation is then done in two different ways: discretizing the

object into a set of points and then performing an FFT, or discretizing the object into

elements with continuous values inside them as stated above and then computing exactly

or approximating to a given error the Fourier transform for each element. The total

acquired signal using the finite element discretization is the following:

Stot(~k) =∑

m

[

fm(T1, T2, ρ)∑

e

se(~k)

]

(1.44)

Where m stands for materials with different T1, T2, or ρ values, and e stands for all

the elements representing the object. The advantage of this method is that if the error

incurred by the finite element discretization is low enough, the error between the actual

k-space values of the object and the simulated values can be either zero or can be made

as low as is required (depending on the allowable computation time) depending on the

element type. Also this approach can be faster than Fourier transform simulators if one

starts with a mesh of a complex object, ideally an unstructured mesh to minimize anti-

aliasing or “stair-step” effects, and takes into account the interpolation step to obtain

a structured grid. This last operation scales as O(E) where E is the number of finite

elements which also scales with the mesh volume [16]. Publications using this approach

are [17] and [18]. In [17], the linear elements tri3 and tet4 are integrated analytically.

If the magnetization inside a certain region of an object can be assumed to be constant


Gauss’ theorem can be used to transform the signal equation into a surface integral [19]:

se(~k) =

∫∫∫

element

me−2πi~k·~xd~x

= m

∫∫∫

element

i

2π~k · ~k∇ · e−2πi~k·~xd~x

=im

2π~k · ~k

∫∫

surface

~k · n̂(x)e−2πi~k·~xdS (1.45)

Where dS is an infinitesimal surface area and n̂(x) is the unit normal pointing outwards

from the closed surface. The normal is constant for a given linear element but varies

with the global coordinates for quadratic elements. As explained in [19], this method is

more efficient than volumetric integration of the signal equation in the following three

ways. No volumetric meshing step is needed where volume elements are constructed

from the surface mesh. The number of elements over which integration is performed is

reduced and scales with the surface area of the shape and not its volume. Also fewer

integration points are needed for surface vs. volume elements for the same integration

order. The analytic solution for tri3 in [17] can also be adapted to solve equation 1.45

exactly. This method also reduces the storage requirements for the phantom compared

to Fourier transform simulation methods that require a structured grid with a resolution

greater than the nominal element size of the mesh [19]. The drawback of this method is

that one has to decide in how many parts and object should be segmented so that the

difference between material properties within each part is kept to a minimum.

The advantage of FE-MRI over voxel based methods is that if the imaging resolution

approaches the voxel size then “stair-step” errors will be produced. These error are also

referred to as partial volume effects. This does not happen if an unstructured mesh

with continuous changes in material properties within each element are used. These

property makes this simulation method ideally suited to study how changing MRI imaging


parameters affect the image quality. With voxel based methods the image quality might

show errors that are due to the simulation itself.

1.3.5 Bloch equation simulators

Bloch equation simulators such as [20] and [21] allow for close to all possible computation

of errors seen in MRI such as the errors due to static field inhomogeneity. However, the

excessive time needed to compute one voxel means that not enough voxels can be used

such that there will be voxel discretization artefacts [20].

1.4 MRI acquisition within the FE-MRI library

The volumetric and surface integration steps described in Subsection 1.3.4 have been

implemented in the FE-MRI library for all the elements in Table 1.2. The main advantage

compared to simulators using a structured grid of T1, T2 and proton density values is

that the interpolation step is not needed if one has an unstructured mesh. If one only has

a surface mesh and can assume these values are constant, the Fourier transform of the

surface mesh can be computed directly from it as explained in Subsection 1.3.4. This is a

fair assumption since organs with the same tissue types will have the same isochromats.

Also the Fourier transform can be calculated exactly for linear elements such as tri3 and

tet4, or with a desired image error for quadratic elements tri6 and tet10.

1.5 Research objectives

In previous work, [18], an error metric for linear elements was derived and used for an

adaptive Gauss quadrature to compute 1.44 and 1.45. The error metric is an estimate

of the error between the calculated and exact signal for a given element and k-space

value. The error metric is used to estimate the number of points for an adaptive Gauss


quadrature that integrates the MRI signal for each element in an unstructured grid.

This was implemented by prof. Steinman and Luca Antiga in the FE-MRI library. The

contributions of this thesis are as follows:

1. A derivation of the error metric found in [18] using the element’s shape functions

and k-space values and derivations of error metrics for three more element types

including quadratic elements

2. An addition to FE-MRI that can estimate errors of Gauss quadratures in many

dimensions from the error obtained in one dimension using a result in [22] and ad-

ditions to FE-MRI that allow it to compute the MRI signal using the new elements

3. A test to see if the most promising of the methods to compute highly oscillatory

integrals (NSD) could be used instead of adaptive Gauss quadrature which is cur-

rently used in FE-MRI

The next Chapters explain in detail each of the above contributions.


Element diagram Name Nodes Shorthand

Linear triangle 3 tri3

Quadratic triangle 6 tri6

Linear tetrahedron 4 tet4

Quadratic tetrahedron 10 tet10

Table 1.2: Finite element types used in this thesis

Chapter 2

Adaptive Gauss quadrature for

FE-MRI elements

2.1 Introduction

To obtain a signal with a uniform error in k-space, an adaptive Gauss quadrature scheme

can be used. Gauss quadrature is a method of numerical integration where the function

to be integrated is evaluated at some point called a quadrature point and multiplied by

a weight for each point called a quadrature weight. The sum of all these products is an

estimate of the integral:

1∫

−1

f(x)W (x) dx =N∑

i=1

wif(xi) + E (2.1)

Where f(x) is the function to be evaluated and wi and xi are the quadrature points and

weights. W (x) is another function that has a constant form. A gauss quadrature is a

special quadrature for numerical integration that allows the exact integration of f(x) for

a given W (x) such as 1 or e−x if f(x) is a polynomial of degree 2N − 1 or less. If the

function is ‘smooth’ in the sense that it can be approximated by a polynomial (such as a

26

Chapter 2. Adaptive Gauss quadrature for FE-MRI elements 27

sinusoid), the integration error will be low. Gaussian quadrature can be readily extended

for higher dimensions. These quadratures are referred to as product rules. Although

these are easily derived from 1D rules, they require more function evaluations than is

necessary to integrate over higher dimensional domains. Optimal Gauss weights and

point locations exist for linear tetrahedral domains, but are only available for low inte-

gration orders and are not easily derived for higher orders. Non-optimal rules use more

Gauss points and weights than optimal rules but they are easier to extend to arbitrary

integration orders. The parametric representation of finite elements allows them to be

integrated in a standard local tetrahedral domain. One can do a coordinate transfor-

mation referred to as a Duffy transformation [23] that maps the quadrature points and

weights of the quadrature product rule to the same local standard tetrahedral domain.

One of the integrals we would like to compute is the MRI signal equation for volumetric

finite elements which is the following:

se(~k) =

∫∫∫

element

m(~x) e−2πi~k·~x d~x

=

∫ 1

0

∫ 1−ξ1

0

∫ 1−ξ1−ξ2

0

|J |m(~ξ ) e−2πi~k·~x d~ξ (2.2)

The term J is the element’s Jacobian which is given by:

J = det

∂x∂ξ1

∂y∂ξ1

∂z∂ξ1

∂x∂ξ2

∂y∂ξ2

∂z∂ξ2

∂x∂ξ3

∂y∂ξ3

∂z∂ξ3

(2.3)

As can be seen from 2.2, the higher the magnitude of the k-space vector, the more


oscillatory the integral is which can only be approximated by a higher order polynomial

for a given approximation error. Therefore on average more quadrature points will be

needed to compute with a fixed error when the magnitude of the k-space vector is higher.

The global coordinates in terms of local coordinates are:

xj =

n∑

i=1

Ni(~ξ ) xij (2.4)

And the magnetization in local coordinates is given by:

m(~ξ) =n∑

i=1

Ni(~ξ ) mi (2.5)

The terms Ni(~ξ ) are called the element’s shape functions and they are polynomials

and n is the number of element nodes which are shown in Table 1.2. The term xij is the

j component of node i. So the integral is a 3D polynomial multiplied by a sinusoid. The

MRI signal equation for surface elements as derived in Subsection 1.3.4 is the following:

se(~k) =im

2π~k · ~k

∫∫

surface

~k · n̂(x)e−2πi~k·~xdS

=im

2π~k · ~k

∫ 1

0

∫ 1−ξ1

0

|J |~k · n̂(x)e−2πi~k·~xd~ξ (2.6)

=im

2π~k · ~k

∫ 1

0

∫ 1−ξ1

0

~k · ~n(x)e−2πi~k·~xd~ξ (2.7)

The vector n̂(x) is the unit normal pointing outwards from the element which is part

of a closed surface. The unit normal is given by:


n̂(x) =~n(x)

|~n(x)| =

∂~x

∂ξ1× ∂~x

∂ξ2∣

∣

∣

∣

∂~x

∂ξ1× ∂~x

∂ξ2

∣

∣

∣

∣

(2.8)

The simplification from equation 2.6 to equation 2.7 can be made since the denomina-

tor in equation 2.8 is also the element’s Jacobian for surface elements. Again the integral

is a polynomial multiplied by a sinusoid.

By changing variables to change the integration domain to the standard tetrahedron,

the same product quadrature rules can be used each time. Each element signal is then

weighted with the appropriate material properties and added to get the full MRI signal

as shown in equation 1.44. To find out the number of quadrature points needed for both

integrals, an estimate of the error for a given number of quadrature points is needed.

Since the integration is carried out in the same standard domain with a fixed size, the

error is only related to the maximum local frequency.

2.2 Gauss quadrature to compute oscillatory inte-

grals

The maximum local oscillation of a function will determine the minimum distance be-

tween quadrature points such that the Nyquist criteria is followed and the equation can

be approximated properly in that area. We can then use a Gauss quadrature with enough

points to guarantee a given accuracy at that point and the rest of the function since the

oscillation everywhere else will be lower. Gauss quadrature also has the advantage of

integrating the magnetization term exactly since it is described by a polynomial in finite

elements. A way of estimating the maximum local oscillation is as follows: Consider the

following 1D function:


y = eig(x) (2.9)

The equation can be approximated at x0 by taking the Taylor series of g(x) at x0 and

substituting the result back into equation 2.9:

y = eig(x0)eig′(x0)(x−x0)ei

1

2g′′(x0)(x−x0)2 ... (2.10)

The term that is non-constant and will usually have the the highest oscillation fre-

quency is eig′(x0)(x−x0). This term produces a frequency of g′(x0). To determine the

maximum local frequency of y to perform a numerical integration from a standard do-

main such as 0 to 1 one should therefore determine the following:

max freq = max|g′(x)|, x ∈ [0, 1] (2.11)

2.3 Computing oscillatory integrals in higher dimen-

sions

In 2D, if a product rule using Gauss quadrature is used with the same number of points

in both coordinates axes to integrate a function in the unit square, the error will be

dependant on the maximum local frequency in the unit square in the directions of the

coordinate axes. Therefore this value will be given by:

max freq = max

(∣

∣

∣

∣

∂g

∂η1

∣

∣

∣

∣

,

∣

∣

∣

∣

∂g

∂η2

∣

∣

∣

∣

)

, x ∈ s (2.12)


Where η1 and η2 are the unit square coordinates and s is the boundary of the unit

square.

2.4 Error estimation for tri3 elements

The integration in local coordinates for triangular elements in the finite element method

and FE-MRI is done using a Gauss quadrature product rule for the unit square where

the point coordinates are transformed using the Duffy coordinate transform [23] to a

unit triangle. This is the same as transforming g in the triangle to the unit square

and integrating it with a Gauss quadrature product rule in the unit square. So to

find the maximum oscillation in the standard triangle one should find how the maximum

absolute values of the gradients of the function g in the unit square relate to the maximum

absolute values of the gradients of g in the unit triangle. Although there are symmetric

quadrature rules for triangles that are optimal in the Gauss-quadrature sense such as

[24], their computation requires finding a non-linear least square solution for a system of

polynomial equations which in [24] could only be done numerically up to 20 points per

dimension due to round-off error. Also the estimation of the quadrature error is very

non-trivial [25]. For a survey of different cubature (quadrature rules for 2 and higher

dimensions) formulas over triangles refer to [25].

To find these relationship lets look more closely at what happens to a plane wave

(produced by the MRI signal equation with a given k-space vector for example) in local

coordinates in the standard triangle when it is transformed to the standard square. Such

a transformation is shown in Figure 2.1.

In this case the gradient of g from the plane wave is constant in local coordinates

along both coordinate axes. This gradient is also proportional to the number of cycles in

the direction of each coordinate axes. The number of cycles is also proportional to the


ξ1

ξ 2

0 0.2 0.4 0.6 0.8 10

0.1

0.2

0.3

0.4

0.5

0.6

0.7

0.8

0.9

1

η1

η2

0 0.2 0.4 0.6 0.8 10

0.1

0.2

0.3

0.4

0.5

0.6

0.7

0.8

0.9

1

Figure 2.1: Transformation of a plane wave in local coordinates in the standard triangleto the unit square

number of red ‘crests’ or blue ‘troughs’ for a given length along a given direction. We

can see that the magnitude of the gradient along η1 remains the same as the gradient

along ξ1 in the unit square at ξ2 = 0 and decreases to 0 when ξ2 increases to 1. We can

also see that the magnitude of the gradient along η2 appears to remain the same as the

gradient along ξ2 in the unit square at ξ1 = 0. However in this case when ξ1 increases

from 0 to 1, the gradient decreases to 0 and increases again to what appears to be the

gradient along the edge from (0, 1) to (0, 1) of the standard triangle. These observations

can be made precise as follows. Based on the derivation in [23] for the standard square

with (−1, 1) bounds, the Duffy coordinate transform between the standard triangle and

square used in FE-MRI is the following:

ξ1 = η1(1− η2) (2.13)

ξ2 = η2 (2.14)


Using the chain rule for a multi-varied function we can relate the gradients of g in

the unit square to the ones in the unit triangle as follows:

∂g

∂η1=

∂g

∂ξ1

∂ξ1∂η1

+∂g

∂ξ2

∂ξ2∂η1

=∂g

∂ξ1(1− η2) (2.15)

∂g

∂η2=

∂g

∂ξ1

∂ξ1∂η2

+∂g

∂ξ2

∂ξ2∂η2

=∂g

∂ξ2− ∂g

∂ξ1η1 (2.16)

The maximum absolute value of equation 2.15 will happen when η2 is 0 and the

maximum absolute value of equation 2.16 will happen when η1 is 0 is either 0 or 1. When

η1 = 0 the gradient is along the same direction as the side from (0, 1) to (1, 0) in the

standard triangle.

In the MRI signal equation g is kxx + kyy or K · x. In finite element analysis, for

triangular linear elements in 2D, the global coordinates are related to local coordinates

in the following way:

x =n∑

i=1

Ni xi, y =n∑

i=1

Ni yi (2.17)

Where n is number of points in the element xi, yi are the global coordinates of point

i and Ni are the element’s shape functions. For the tri3 element they are given by:

N1 = ξ1, N2 = ξ2, N3 = 1− ξ1 − ξ2 (2.18)


Using the element’s shape functions the maximum oscillation in the ξ1 and ξ2 direc-

tions can be determined by:

∂g

∂ξ1= kx

∂x

∂ξ1+ ky

∂y

∂ξ1= K · e31 (2.19)

∂g

∂ξ2= kx

∂x

∂ξ2+ ky

∂y

∂ξ2= K · e32 (2.20)

Where eij is a vector from element vertices i to j in global coordinates. Using the

above two results yields:

∂g

∂ξ2− ∂g

∂ξ1= K · e32 −K · e31 = K · e12 (2.21)

Note that these values are constant in local coordinates since the partial derivatives

of the shape functions for linear finite elements are constant, therefore checking their

maximum value in domain s is not necessary. These are the same error metrics used

in [18]. In [18] the error metric is arrived at by observing that it should be invariant

with respect to element size and shape and the k-space vector. These results can be

interpreted as the dot product of the k space vector in local coordinates with each of the

standard triangle side vectors expressed in terms of ξ1 and ξ2 which give the oscillation

frequency per unit length in the local coordinate axes. The domain s in global and local

coordinates is given in Figure 2.2.

Since the k space vector in global coordinates is the following:


y

x(0,0) (1,0)

(0,1)

Figure 2.2: tri3 domain s and vertices in global and local coordinate systems with a givenK space vector

K =

kx

ky

=

∂g∂x

∂g∂y

(2.22)

The k space vector in local coordinates is given by:

Kl =

∂g∂ξ1

∂g∂ξ2

=

∂x∂ξ1

∂y∂ξ1

∂x∂ξ2

∂y∂ξ2

∂g∂x

∂g∂y

= JK (2.23)

Where J is the element’s Jacobian. Using the above result, the maximum oscillation

estimates can be written as follows:

∂g

∂ξ1= Kl · (1, 0) = Kl · ξ̂1 (2.24)

∂g

∂ξ2= Kl · (0, 1) = Kl · ξ̂2 (2.25)

∂g

∂ξ2− ∂g

∂ξ1= Kl · (−1, 1) = Kl · (ξ̂2 − ξ̂1) (2.26)


Where ξ̂1, ξ̂2 and ξ̂2 − ξ̂1 are the standard triangle side vectors. For the tri3 surface

element the same results hold, except that the element sides eij and the k-space vector

are in three dimensions. In this case the local k-space vector in the still 2D local domain

is a projection of the original k-space vector.

2.5 Error estimation for tri6 elements

For quadratic elements equation 2.12 still applies, however the values of the gradients of

g with respect to ξ1 and ξ2 in local coordinates will vary linearly with ξ1 and ξ2. The

boundary s in local coordinates for tri6 is the same as the one for the tri3 element. Since

the gradient of g for a given local coordinate g′(ξ1, ξ2) is a plane in the space formed by

ξ2 and ξ2 and the possible values of g′(ξ1, ξ2), therefore the maximum absolute value of

g′(ξ2, ξ2) will be at one of the vertices of the standard triangular boundary. Therefore

equation 2.12 will have to be computed 3 times for each gradient for a total of 3 x 3 =

9 gradients. A plane formed by the gradient of g in one of the 3 coordinates and the

possible maximum absolute values are shown in Figure 2.3.

The number of gradients that need to be computed is the number of gradients times

the number of element vertices which is 9. The computation time will be lower if the

gradients with respect to ξ1 and ξ2 are computed first and then the results are used to

compute the gradient given by equation 2.16 with ξ1 = 1.

These gradients can be computed even faster by storing and reusing some of the

intermediate results. The needed computations for the tri6 element are as follows. The


Figure 2.3: The plane formed by the gradient of g with respect to ξ1 and the location ofits possible maximum absolute values in domain s. In this case the maximum absolutevalue of the gradient is at ξ1 = ξ2 = 0

element’s shape functions are as follows:

N1 = ζ(2ζ − 1)

N2 = ξ(2ξ − 1)

N3 = η(2η − 1)

N4 = 4ξζ

N5 = 4ξη

N6 = 4ηζ (2.27)

For the gradient in the ξ1 direction:


∂g

∂ξ1= kx

∂x

∂ξ1+ ky

∂y

∂ξ1(2.28)

= kx

6∑

i=1

∂Ni∂ξ1

xi + ky

6∑

i=1

∂Ni∂ξ1

yi (2.29)

= K

[

x

y

]

∂N

∂ξ1(2.30)

Where K is a 1 by 2 matrix, [x y]T is a 2 by 6 matrix with the element’s node

coordinates and ∂N/∂ξ1 is a 6 by 1 matrix. The same result applies for the gradient in

the ξ2 direction. The matrices ∂N/∂ξ1 and ∂N/∂ξ2 can be precomputed for all elements

by evaluating them at the local domain vertices (0,0),(1,0),(0,1). Half of the entries in

these matrices are 0 therefore it is more useful to calculate the product [x y]T∂N/∂ξ1

directly.

2.6 Error estimation for tet4 elements

For linear tetrahedral elements, to estimate the quadrature error, one has to estimate

the maximum oscillation frequency in the standard cube given by equation 2.31.

max freq = max

(∣

∣

∣

∣

∂g

∂η1

∣

∣

∣

∣

,

∣

∣

∣

∣

∂g

∂η2

∣

∣

∣

∣

,

∣

∣

∣

∣

∂g

∂η3

∣

∣

∣

∣

)

, x ∈ s (2.31)

Where η1−3 are the unit cube coordinates and s is the boundary of the unit cube.

The Duffy transformation for the unit cube, again based on the derivation in [23] for the

standard cube with (−1, 1) bounds is the following:


ξ1 = η1(1− η3)[1− η2(1− η3)] (2.32)

ξ2 = η2(1− η3) (2.33)

ξ3 = η3 (2.34)

The derivation can be understood as using the 2D transformation recursively in two

of the coordinate axes to ‘collapse’ the standard cube into the standard tetrahedron.

Using equations 2.32, 2.33 and 2.34, the gradients of g with respect to the standard cube

coordinates are the following:

∂g

∂η1=

∂g

∂ξ1(1− η3)[1− η2(1− η3)] (2.35)

∂g

∂η2= − ∂g

∂ξ1η1(1− η3)2 +

∂g

∂ξ2(1− η3) (2.36)

∂g

∂η3=

∂g

∂ξ1η1[2η2(1− η3)− 1]−

∂g

∂ξ2η2 +

∂g

∂ξ3(2.37)

Since the values of ξ1−3 can only vary from 0 to 1, the maximum absolute value of

the following gradients will determine the maximum oscillation frequency of g:

∂g

∂ξ1,

∂g

∂ξ2,

∂g

∂ξ3(2.38)

∂g

∂ξ2− ∂g

∂ξ1,

∂g

∂ξ3− ∂g

∂ξ2,

∂g

∂ξ3− ∂g

∂ξ1(2.39)


∂g

∂ξ3− ∂g

∂ξ2− ∂g

∂ξ1,

∂g

∂ξ3− ∂g

∂ξ2+

∂g

∂ξ1(2.40)

The gradients shown in 2.38 and 2.39 are gradients along all the side edges of the

standard tetrahedron as shown in the 2D case for the triangle. However the gradients

shown in 2.40 which are derived from equation 2.37 are not related to any side of the

standard tetrahedron, and would not have been taken into account in the maximum

frequency calculation if extended the results for 2D elements (that the gradients along

all the element side edges determine the maximum oscillation frequency) directly to 3D

elements.

2.7 Error estimation for tet10 elements

For quadratic tetrahedral elements equation 2.31 still applies, however their values will

vary linearly on ξ1−3 like in the tri6 element. The boundary s for tet10 is the same as the

one for the tet4 element. To estimate the maximum frequency gradients 2.38, 2.39 and

2.40 should be calculated at the standard tetrahedron vertices. The gradient calculations

can be performed faster if all the components of the gradients are computed before hand

and then used for gradients 2.39 and 2.40. Also the values of some of the intermediate

steps can be stored and reused like in the case of the tri6 element.

Chapter 3

Algorithm details and

implementation

3.1 Introduction

The FE-MRI algorithm simulates either a 2D or 3D MRI signal acquisition. If a 2D

signal is required the program computes a given number of k-space planes depending

on the slice thickness and the needed resolution. If not the program computes a given

k-space volume. If using an acquisition with a fixed number of quadrature points, the

program computes the needed Gauss quadrature and uses that quadrature for all elements

and k-space values. If the adaptive quadrature is used, the program computes all the

Gauss quadratures requested, then it computes the ones needed for all the element/k-

space point combinations and uses the appropriate ones in order to keep the k-space

image error constant. If more quadrature points are needed in some elements the user

is warned. Then the user can either use the result knowing some points at the edge of

k-space might have a slightly higher image error or he/she can run the function again

with more quadrature points. This is a useful feature if the computation time even using

the adaptive algorithm with all the needed quadratures is prohibitive.

41

Chapter 3. Algorithm details and implementation 42

3.2 FE-MRI software architecture

The FE-MRI library is implemented as a set of python functions that can be used through

a command line interface by typing a script in Linux or Unix. An example of a script

that can be typed in the Unix command line after installing the FE-MRI library is given

below:

femritrisurfmeshkspace -ifile modifiedQuadSphere.vtu -dimensionality 2

-fov 30 30 30 -fovcenter 0.0 0.0 0.0 -matrix 32 32 1 -useoptimal 1

-error 0.001 -qorder 59 -sliceprofile quadratic -slicethickness 6

--pipe femrikspacezeropadding -padding 64 64 64

--pipe femrikspacetoimage --pipe vmtkimageviewer

The above script reads an unstructured mesh of a sphere-like closed surface composed

of tri6 elements. Then a 2D acquisition is performed with a field-of-view of 30, a 32 by

32 k-space matrix and a quadratic slice profile in k-space. This is done using the optimal

algorithm which can use a Gauss quadrature up to order 59 or points per dimension. Then

the k-space is zero-padded and transformed back into an image which is displayed on the

screen. Since the size of the shape is 40 by 40 by 40 units, the resulting image displays

an aliasing error. In FE-MRI information from one function can be used into the next

needed function without writing files by using the –pipe command. Some of the python

functions call C++ functions to perform the needed computations. Both languages are

used together so the user is able to both use fast and efficient C++ functions while being

able to easily customize either the script or the python functions. The C++ functions can

also be modified if needed however it is usually harder to modify that a python script.

Both the python and C++ functions use data structures and visualization functions

from the visualization toolkit or VTK1 and the vascular modelling toolkit or VMTK2

1Visualization toolkit website: https://www.vtk.org/2Vascular modelling toolkit website: https://www.vmtk.org/


libraries. The mesh used in the example script above is in VTK’s .vtu format. The

FE-MRI acquisition algorithm and its related C++ functions are shown in Figure 3.1.

The algorithm shown in Figure 3.1 is implemented in the vtkfemriTetVolNumer-

icKSpaceGenerator C++ class for meshes composed of tet4 and tet10 elements. The

algorithm can use a mesh with both elements present. The same is the case for a surface

mesh when it is used by its C++ class vtkfemriTriSurfNumericKSpaceGenerator, it can

contain both tri3 and tri6 elements. The class contains a set of data structures such as

the k-space and functions that operate on them. This class is used to declare an object

(an instance of a class) that then can call other objects declared using the classes vtkfem-

riVolTetQuadratureOrderCalculator and vtkfemriGaussArbQuadrature which compute a

list of the cycles per element as described in Chapter 2 and their corresponding num-

ber of quadrature points to obtain a desired image error and the needed quadratures

respectively.

3.3 K-space acquisition algorithm

The k-space signal acquisition in the flowchart in Figure 3.1 is done using Algorithm

1. The k-space acquisition algorithm shown in Algorithm 1 starts by calculating the

maximum cycles/element from all the k-space point and element combinations. This is

done using a design table generated for a user specified error. The design table used for

the optimal algorithm is a table that contains all the number of quadrature rules required

and its corresponding cycles per element for a given error. Then it computes and stores

all quadratures from orders 1 to the order needed for the maximum cycles/element to get

the required error if the optimal algorithm will be used. The quadratures are calculated

by the vtkfemriGaussArbQuadrature C++ class. Then the algorithm loops through all

the k-space values needed and all the elements in the mesh and uses the given Gauss

quadrature to compute the MRI signal. One thing to take into account in the algorithm


Start

Read mesh and MRIacquisition parameters

Dimensionality?

Find location andboundary of slices

Set slice magnetization profile

k-space aquisition

More slices?

Output k-space file and/orsend data to next python

function

Stop

2D

No, or 3D acquisition

3D

Yes

Figure 3.1: Flowchart of the main FE-MRI algorithm


is to make sure the magnetization and the Jacobian are being evaluated exactly. If

the optimal algorithm is used, for k-space vectors with a low magnitude the number of

quadrature points needed will be lower than the ones needed to compute the polynomial

composed of the Jacobian multiplied by the magnetization. This is because even if the

sinusoidal term is close to a straight line at low magnitudes of the k-space vector, needing

only an order of 1 for example, one still wants enough quadrature order to compute the

quadrature of the polynomial composed of the Jacobian multiplied by the magnetization.

Algorithm 1 K-space acquisition algorithm

if optimal algorithm thenmake design table relating cycles/element to quadrature order at desired image errorcompute Gauss quadratures from order 1 to specified order

end iffor each needed k-space location k doset S(k) = 0for each element e doif optimal algorithm thencalculate cycles/element from k and eset Gauss quadrature to get desired error for calculated cycles/element usingdesign table

elseset to user-defined Gauss quadrature

end ifif order of Gauss quadrature < order of magnetization and Jacobian thenset Gauss quadrature to same order as magnetization and Jacobian

end ifobtain Gauss points pobtain corresponding Gauss weights wset s(k) = 0for each p doevaluate s(k) = s(k) + w m(p) exp(k, p)

end forevaluate S(k) = S(k) + f(T1, T2, ρ) s(k)

end forend for


3.4 Cycles per element quadrature table

The algorithm to find the design table is shown in Algorithm 2.

Algorithm 2 Design table algorithm

compute max cycles/element (max cpe) for all k-space points and elementsfor 1 to max number of quadrature points specified, n doset point list, pointList(n) = nset start cpe to 0, cpe = 0set error to 0while error < error threshold doadd cpe interval to cpe, cpe = cpe+ dcpeget exact integral, Ie(cpe)obtain Gauss points pobtain corresponding Gauss weights wset Ic = 0for each p doevaluate Ic = Ic + w exp(cpe, p)

end forevaluate error = |mag(Ie)−mag(Ic)|dim(e)

end whileset cpe list, cpeList(n) = cpe

end forif max cpe > cpeList(max n) thendisplay warning

end if

The algorithm is carried out in the vtkfemriQuadratureOrderCalculator C++ class.

The algorithm first calculates the maximum cycles/element from the mesh and the MRI

acquisition parameters. Then it loops over all the quadratures from 1 to a number

specified by the user, storing each number in pointList. Within that loop it cycles

thought increasing cycle/element values, until the error is above the user defined value.

Then the cycles/element are stored in the corresponding entry in cpeList. The integral

being computed is akin to a 1D MRI signal equation:

I =

∫ 1

0

e−2πik (3.1)


Where k is the cycles/element. The exact solution of its magnitude is:

mag(Ie) =1

2πk

√

2− 2 cos(2πk) (3.2)

The error is defined as follows:

E = |mag(Ie)−mag(Ic)|dim(e) (3.3)

Where Ic is an approximation of equation 3.1 calculated using Gauss quadrature and

dim(e) is the dimensionality of the elements that the mesh is composed of. Note that

the algorithm could have looped over the cycles/element first. However this could have

caused some integrals computed with the quadrature rule to have a lower error simply

because they are not being ‘sampled’ properly by the points; the Nyquist criteria is not

followed with that loop ordering. With the current loop ordering the oscillatory integral

gets more oscillatory while the number of quadrature points is kept the same, making

sure that once the error is larger than the threshold there were also enough points to

sample the function. The only difference between the algorithm when computing it for

2D or 3D elements is that the 1D error is multiplied by 2 and 3 respectively. This is

because the error of a product rule if using the same number of quadrature points in

all dimensions can be estimated to first order as the error in one dimension times the

number of dimensions of the product rule [22]. So the algorithm finds the error in one

dimension and multiplies it by the number of dimensions to get an error estimate for the

product rule which is then used to evaluate the element signal s(k) in Algorithm 1.


3.5 Calculation of quadrature rules

The quadrature rules are computed for the design table and the k-space acquisition cal-

culations done in the C++ classes vtkfemriVolTetQuadratureOrderCalculator and vtk-

femriTriSurfNumericKSpaceGenerator. Both classes make an object which is an instance

of the class vtkfemriQuadratureOrderCalculator. The class can produce Gauss-Laguerre

and Gauss-Jacobi quadratures. The class also computes all the product rules needed by

all the classes. These quadratures have different weighting functions W (x) as shown in

equation 2.1. This allows one to make a product rule that integrates the Jacobian from

the Duffy transformation exactly. Therefore the only error comes from the transformed

coordinates in the oscillatory term. That is why the derived error metrics in Chapter

2 only take into account that term. The details of how the quadratures are computed

are given in Appendix A. The author has added a simpler design table calculation, all

the error metrics in Chapter 2 and a new class that can compute an arbitrary number of

quadrature points and weights for the relevant Gauss quadratures.

3.6 Slice profile simulation for volumetric meshes

A method to simulate an arbitrary slice profile could be to adaptively subdivide the

elements inside the slice to linearly approximate the slice. This is referred to in the

finite element literature as h-refinement. Another approach is to use a higher polynomial

order for the elements inside the slice to adaptively get a more accurate simulation.

This method is referred to as p-refinement. While those methods will give an accurate

solution, they are computationally expensive. A more practical method that does not

require modifying the mesh or the elements is to weigh the s(k) signal in Algorithm 1

after it is evaluated with a quadrature point, by using the location of the quadrature

point and the required slice profile. For example, if one requires a quadratic slice profile

along the z-direction, the weighting based on the location of the quadrature point within


the slice is the following:

m = 1− 4t2(pz − z0)2 (3.4)

Where m is the weighting, t is the slice thickness, z0 is the slice origin location and

pz is the z-component of the quadrature point. Everywhere else the weighting is 0. An

ideal slice profile is simulated simply by making the weighting 1 within the slice and

0 otherwise. Notice this method does not compromise the accuracy of the result given

by the quadrature since the weighting convoluted with the magnetization distribution

forms a new function that the quadrature is integrating. The ideal, quadratic and a

trapezoidal slice simulation was already implemented in the C++ class vtkfemriTetVol-

NumericKSpaceGenerator of the FE-MRI library.

3.7 Additions and improvements to current FE-MRI

functions

The new error metrics in Chapter 2 have been implemented as C++ functions in the FE-

MRI library. The code has been improved by removing the normalization and Jacobian

calculation step for surface elements which are the same value for a given quadrature

point and cancel each other. The algorithm to find the design table shown in Algorithm

2 was rewritten using dynamically allocated C++ arrays instead of vtk arrays to make

its execution faster and to aid in readability. Also instead of having a lookup table for

the quadrature weights up to a certain number of points, the user can specify how many

points can be used. The points are calculated from first principles once and stored at the

beginning of the calculation. The details of the quadrature point and weights calculation

can be found in Appendix A. If the program needs more points to calculate a certain


k-space point there is a warning shown that recommends using more points and running

the function again. If not the program uses the maximum number of points available.

The memory to store the points is dynamically allocated and the points are stored in a

contiguous memory block for faster access. This is done in C++ as follows:

inline double** newQuadratureMatrix(int numPs)

{

double *matrix = new double[numPs*(numPs+1)*(2*numPs+1)/6];

double **m = new double *[numPs];

for(int i = 0; i < numPs; i++)

{

m[i] = &matrix[i*(i+1)*(2*i+1)/6];

}

return m;

}

The above function returns an array of pointers in m to the memory block named

matrix, which contains all the points from 1 to numPs. The actual number of points is

the sum of the square of the the 1 to numPs series since we are storing Gauss quadrature

product rules which in this case are for 2D surface elements. The series needed and its

result is shown below:

n∑

x=1

x2 =n(n+ 1)(2n+ 1)

6(3.5)

This type of sum can be found using finite calculus [26]. Also the computation time

has been cut in close to half by computing only half of k-space using the property [9]:

se(−k) = s∗e(k) (3.6)


Where s∗e(k) is the complex conjugate of the original MRI signal. Equation 3.6 applies

for the Fourier transform of a real signal. However in actual MRI acquisition this con-

straint cannot be exploited because of object motion and magnetic field inhomogeneities

[9]. In the case of a simulated signal object motion can be added later to the simulated

signal by adding a motion dependent phase to the computed k-space for example.

Chapter 4

Results and discussion

4.1 Introduction

This chapter shows the k-space error maps and error histograms for meshes with each

of the element types discussed in this thesis. This is done to confirm the efficacy of the

algorithms in Chapter 3. Also the MRI acquisition of an anatomically realistic carotid

artery was simulated to show the potential uses of the surface FE-MRI algorithm. The

desired error threshold for all elements tested was 10−3.

4.2 Validation of optimal algorithm for tri3 elements

To validate the optimal algorithm for tri3 elements first the k-space error map and a

histogram of the errors found in Table 4.1 for the adaptive algorithm were computed

for a deformed unit cube. The deformed cube mesh and its simulated MRI image is

shown in Table 4.2. The histogram shows a sharp peak at an error of 10−4. A sharp

peak is expected since the oscillation due to the k-space value is constant within each

element. The lower value for the error might be due to the error estimate for product

rules given in [22] being a conservative estimate. The shape was chosen because it was

easy to generate in the Paraview program starting with a meshed cube. Paraview is

52

Chapter 4. Results and discussion 53

based on VTK and can produce files that can be read by the FE-MRI library. Then

the shape was made asymmetric by moving the nodes. This was done to highlight any

possible errors in the algorithm which might not be evident by imaging a symmetric

object. The imaging parameters are as follows: FOV = 2.5 x 2.5, image size = 32 x 32

pixels. The magnetization value is 1 throughout. The exact image was computed using

the exact algorithm available for surface linear elements in FE-MRI. As an example, the

exact k-space using the analytic solution for the tri3 element

efficientfourier transformation of unstructured meshes and ......derivatives and is then extended...

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