Supplemental File 1 - Basic information on MRI
Nuclei most commonly used in in vivo NMR
Especially due to their importance and relative high concentration, the nuclei shown
in Table 1 are most commonly used in in vivo MRI. The Larmor frequencies at
17.6T of the listed nuclei are provided to illustrate the enormous frequency range.
Table 1: NMR properties of nuclei most commonly used in in vivo NMR
Isotope
Spin
γ/[MHz/T]
Frequency @17.6 T
/[MHz]
natural abundance/ [%]
1H
2H
13C
19F
23Na
31P
39K
1/2
1
1/2
1/2
1 1/2
1/2
1 1/2
42.58
6.54
10.71
40.08
11.27
17.25
1.99
750
115
189
706
198
304
35
99.985
0.015
1.108
100
100
100
4.672
Standard elements of an MRI experiment
In general a MRI experiment can be divided in three parts, slice selection, phaseencoding and frequency-encoding as shown in a standard two-dimensional gradient
echo sequence (Fig. S1A).
-
slice selection: The combination of a frequency selective 90o RF pulse with a
frequency bandwidth of z and a magnetic gradient selectively excites protons
within a slice thickness given by z:
z 
 z
Gslice
The slice thickness can be derived by looking at the spatial dependency of the
Larmor frequency when an additional linear magnetic field gradient in z direction
is applied:
(   z )   ( B0  zG z )
The opposite gradient following the slice selection shown in Fig. 1A is refocusing
the signal.
1
-
phase encoding: The phase encoding gradient G y imparts a spatially dependent
phase shift into the signal. The whole experiment has to be repeated N PE times
with different Gy values to cover the whole k space.
ky 
-

G y PE
2
frequency encoding: During data acquisition the frequency encoding gradient
generates a spatially dependent precessional frequency in the acquired signal.
kx 

Gxt
2
The opposite gradient shown in Figure 1A in the frequency encoding direction
has the purpose to dephase the signal before it gets rephrased during the
acquisition. In each repetition a whole k-space line with NFE points is acquired
(see Fig. S1B).
A
Spin echo pulse sequence
900
1800
B
k – space scheme
ky
echo
RF
slice Gz
NPE
freq. Gx
NFE
acqu.
kx
phase encoding (NPE)
phase Gy
TE
frequency encoding (NFE)
TR
Figure S1. The left panel (A) shows the most commonly used T2 weighted spin
echo (SE) sequence. The schematic representation of a two dimensional k-space is
shown on the right (B). The light grey arrow indicates the acquisition of all
frequency encoding points during one repetition (repetition time: TR) with the phase
encoding gradient being at its third step in positive direction.
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Standard imaging contrast
To illustrate the different contrasts achievable with MRI a slice of a carrot tap root
was acquired with different imaging parameters using standard imaging sequences.
Example of relaxation times for protons in plants at 11.75 tesla are 1.5 s for T1, 15
ms for T2 and 5 – 10 ms for T2* respectively (Köckenberger, 2001). An image is
referred to being T1 weighted if the TR is kept relatively short. While tissues with
long T1 values will produce a hypo intense signal, tissues with shorter T1 will
appear brighter. The overall signal intensity of the image will be reduced compared
to no T1 weighting (TR = 5 x T1). Figure S2A shows the same slice of the carrot tap
root acquired with different repetition times (TR = 10s, 2s, and 0.5s). The signal
loss from A(i) to A(iii) is evident and is equal to a loss of signal to noise ratio (SNR)
from 129 in A(i) to 36 in A(iii). To increase the visibility of A(iii), the brightness of the
image was adjusted in A(iv). Compared to the non T1 weighted image A(i), the
meristematic tissue in this image (white arrow) is much more pronounced,
indicating shorter T1 values compared to the surrounding tissue.
T2/T2* weighting is achieved by extending the echo time. With a longer TE the time
for relaxation increases and tissues with shorter T2/T2* times will generate a more
hypo intense signal compared to the ones with longer T2/T2*.
The before mentioned GE sequence produces T2* weighting rather than T2
weighting in the images as the relaxation of M xy due to the inhomogeneous B 0 is not
recovered. An application of an (slice selective) 180 degree RF pulse in the middle
of TE has the ability to refocus those losses. This sequence (based on the Hahn
spin echo) is therefore called spin echo sequence and is the most commonly used
sequence for T2 weighting.
T2 weighted images from a spin echo sequence are shown in Fig. S2B. With
increasing echo time (TE = 9.5, 15, and 25 ms) the areas with shorter T2 are
getting darker (arrows) compared to the surrounding tissue. Fig. S2C displays
images from the same slice acquired with a GE sequence (TE = 2.6, 5, 10 ms). It
can be seen that the signal is reducing much faster at much shorter TE compared
to the SE images indicating T2* instead of T2 weighting. While the side root in the
SE images at a TE of 25 ms is still visible (Fig. S2B(iii)), it can hardly bee identified
at 10 ms TE in the GE image (Fig. S2C(iii)).
The image closest to a proton density weighted image is Fig S2A(i) (same as
S2B(i)) as the TR is very long and only a modest T2 weighting was applied. As TE
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in standard pulse sequences cannot be reduced to zero and a TR of 5 x T1 results
in enormous imaging times, most of the MRI images are both T1, as well as T2/T2*
weighted, depending on the desired contrast, and therefore the imaging parameters
used, one weighting might be chosen to be bigger than the other.
Figure S2. Images of the same slice of a carrot tap root. While the first two rows
are SE images the third row are GE images with different image weighting. A(i)A(iii) show the effect of T1 weighting (TE= 9.5ms; TR= 10s, 2s, 0.5s). A(iv) is the
same image as A(iii) with an increased image intensity. B(i)-B(iii) illustrate different
T2 (TR= 10s; TE=9.5, 15, and 25ms) and C(i)-C(iii) different T2* (TR=100ms;
TE=2.6,5,10ms) weightings in the image. Besides the mentioned weightings
additional signal reduction is present in every image due to diffusion effects and
susceptibility differences within the sample (see: ‘Image resolution and its limits’).
Furthermore, quantitative parameter maps showing the distribution of MRI
parameters can be generated. The most common maps display the distribution of
the equilibrium magnetization M0, the T1, the T2 and the T2* times. To generate
such a map the data needs to be acquired in a certain way. A multi spin echo
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sequence (RARE, Hennig et al., 1986) for example can be used to acquire a T2
map. Instead of using the acceleration factor of the sequence (see Supplement 2)
each echo is used to generate its own image. As the phase encoding gradient is
changed after each echo train (N echoes) and not after each echo, the resulting
imaging time increases by a factor of N. The resulting N images can be fitted on a
pixel by pixel basis to the following equation:
S (t )  S0 exp( t / T2 )
While S0 is the signal at TE=0, t represents the different echo times of the N
images. Solving the equation for T2 reveals the T2 values of the sample in each
pixel. Example of such T2 maps on the developing barley grain can be found in
Glidewell et al (2006).
As discussed in the next paragraph additional signal reduction caused e.g. by
diffusion effects and magnetic susceptibility changes within the sample need to be
considered especially when moving towards high resolutions. Instead of T2 values
often apperent T2 values are reported. A comprehensive overview how the field
strength B0, the spatial resolution and the echo time influences the observed T2
values can be found in Edzes et al (1998).
Image resolution and its limits
In general, the voxel size of a two dimensional MRI image is defined by the slice
thickness and the ‘in plane’ resolution x and y:
x 
FOV x
1
1
2



Nx
N x k x k x ,max G x Tacq
y 
FOV y
Ny

1
1
2


N y k yx k y ,max
N y G y ,max  PE
While in this example Nx is the number of frequency and N y is the number of phase
encoding points, FOV x and FOVy are the field of views in the corresponding
directions. The maximum k x,max is reached with t = T acq (Tacq = acquisition time). It
needs to be noted that MRI is not limited to two dimensional imaging, a real three
dimensional dataset can be achieved by applying an additional phase encoding
(NPE2) in the slice selection direction.
If the resolution is not high enough to resolve the desired structure, or if the voxel
lies on an edge of a compartment within the specimen, the voxel intensity is a
mixture from signals of different tissues. This ‘partial volume’ effect will result in a
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blurred image and real structures might not get resolved. An example is shown in
Fig. S3. Although the images are not acquired at the exact same location it is
obvious that structures resolved in the higher resolution image, are blurred in the
low resolution one.
Figure S3. High (left panel) and low (right panel) resolution images of a carrot tap
root. Due to the partial volume effect the structures within the low resolution image
appear blurred.
Due to the T2 relaxation during the acquisition of the data (frequency encoding
direction) the point spread function for each pixel will have a line width which also
influences the resolution. If the point spread functions of two adjacent pixels does
not fulfil the Rayleigh criterion the two pixels cannot be separated. As the T2 times
in plants are usually relatively short and therefore yield large line widths (FWHM =
1/T2) the bandwidth (BW) of the experiment needs to be chosen to make sure that
the T2 broadening is less than one pixel.
As plants have many air pockets susceptibility effects in T2* weighted imaging can
lead not only to a reduced resolution but also to image distortions. A good review
about susceptibility effects on resolution is given by Callaghan et al. (1994).
Another effect that needs to be taken into account is the diffusion of the spins out of
the voxel in the frequency encoding direction during the acquisition of the signal
Tacq. Although the effect is almost negligible at bigger voxel sizes, it needs to be
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considered and can be substantial in smaller voxels. The average displacement
<x> of, in our example, water molecules undergoing Brownian motion with a selfdiffusion constant DH20 = 2.23 x 10-9 m2s-1 at 20 oC can be written as:
 x  2  D  Tacq
To achieve a resolution x of 10 m with a FOVx of 4 mm e.g., N x = 400 points
need to be acquired in the frequency encoding direction. Assuming a typical
bandwidth (BW) of 50 kHz will result in an acquisition time of T acq = Nx/BW = 8 ms.
With a resulting mean diffusion length of 6 m the average displacement <x> is
already 60% of the voxel size resulting in substantial attenuation of the signal.
According to the equation for resolution (x) calculations the necessary gradient
strength Gx would be 1.8 T/m which exceeds already the capability of most
commercially available gradient sets and shows another limit for higher resolution.
To reduce the necessary gradient strength the T acq needs to be extended resulting
in an even bigger signal loss due to diffusion.
These remarks show that by increasing the resolution, besides the fact that less
spins are within one voxel resulting in a lower SNR, new problems arise that need
to be taken into consideration.
SNR improvements
Except moving to a lower resolution other measures can be applied to improve the
SNR. Moving to a stronger magnet will in general increase the SNR by a substantial
amount. While at lower field strength the theoretical increase is proportional to B 07/4
at high field strength the SNR gain tends to become closer to linear to B 0. As plants
have a lot of susceptibility changes within their tissue mostly due to air enclosures
using stronger magnets is not always advantages as additional artefacts can arise
(T2*).
Using an adapted RF resonator is another measure to increase the SNR. Besides
using the before mentioned cryoprobes many other possibilities are available to
construct an optimized RF resonator. Adapting the size of the RF resonator to ‘just
fit’ the sample increases the SNR (high filling factor). Furthermore, smaller
resonators have a higher sensitivity than larger ones and therefore achieve a higher
SNR. While volume resonators like a solenoid or a birdcage resonator can be used
for small samples (e.g. seeds, leaves, pods) a surface resonator like a simple single
loop resonator is more appropriate for studies of small structures on larger plants
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(e.g. the bark of a tree). Adding a preamplifier right next to the surface resonator
increases the SNR furthermore. The need of a switch between transmission and
reception of the signal of this setup would complicate the construction by a fair
amount.
The most straight forward way to increase the SNR is to use averaging and the
appropriate (fast) imaging sequence.
Chemical shift and principles of chemical shift imaging
The electrons surrounding a nucleus produce a magnetic field which is in the
opposite direction to B 0. This field, which is proportional to the static magnetic field
B0, shields the nucleus from B 0 and the resonance frequency of the nucleus
changes slightly depending on the strength of the additional field. Nuclei in different
chemical environments (different positions in a molecule) have distinct but different
resonance frequencies. The effective magnetic field at the nucleus is described by:
B  B0 (1   )
where  is the shielding constant. In this way, different molecules can be identified
in a NMR spectrum which is the result of a FID after a Fourier transformation. As
the frequency shifts are usually very small  is measured in parts per million (ppm).
For in vivo proton NMR spectroscopy all the peaks that occur in a spectrum are in
the range of about 12 ppm. By applying a water suppression technique (e.g.
CHESS (Haase et al, 1985), or VAPOR (Griffey et al, 1990) the dominant water
resonance can be suppressed and the resonances from much lower concentrated
sugars or amino acids for example can be detected.
In a standard three dimensional chemical shift imaging (CSI, Brown et al, 1982)
experiment all three directions are phase encoded and a FID is acquired without
any gradients present. The result is a four dimensional dataset in which each pixel
of the three spatial dimensions contains a NMR spectrum. Integrating a peak allows
to construct a three dimensional distribution of the observed resonance (metabolite
map). The biggest trade-off of a CSI experiment is the long acquisition time since
every point in k- space has to be acquired in a different transient and the relative
low resolution due to the low metabolite concentrations.
References
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Brown T.R., Kincaid B.M., Ugurbil K. (1982) NMR chemical shift imaging in three
dimensions Proc. Natl. Acad. Sci. USA 79, 3523-3526.
Callaghan, P.T. (1994) Principles of Nuclear Magnetic Resonance Microscopy.
Clarendon Press, Oxford, England.
Edzes, H.T., van Dusschoten, D., and Van As, H. (1998) Quantitative T2 imaging
of plant tissues by means of multi-echo MRI microscopy. Magn. Reson. Imaging
16, 185-196.
Griffey, R.H., and Flamig, D.P. (1990) VAPOR for solvent – suppressed, short
echo, volume – localized proton spectroscopy. J. Magn. Res. 88, 161-166.
Haase, A., Frahm, J., Hanicke, W., and Matthaei, D. (1985) 1H NMR chemical
shift selective (CHESS) imaging. Phys. Med. Biol. 30, 341-344.
Köckenberger, W. (2001) Nuclear magnetic resonance micro-imaging in the
investigation of plant cell metabolism. J. Exp. Bot. 52, 641-652.
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