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Electromagnetic Fields in Biological Systems
• Its amplitude is dependent on the number of involved hydrogen nuclei and hence
on the water content of tissue.
• Its time course is mainly influenced by two different mechanisms. The first is the
release of the absorbed RF EMF energy quantum and return of nuclei to their basic
state. Consequently the longitudinal magnetization is recovering at the expense of
the transversal vector component. The time constant of this mechanism is dependent on the tissue structure and is called spin–lattice time constant T 1 .
• The second mechanism responsible for the signal decay is dephasing of the spinning nuclei due to local B-field variations. These are caused by the influence of
neighboring nuclear magnetic moments (spins). Fluctuating local B-fields cause
fluctuating Larmor resonance frequencies and consequently varying phase angles
of nuclear spins. The vectorial summation of increasingly dephased magnetic components results in the ongoing reduction of the macroscopic transverse magnetization. Because nuclei continue spinning, they still keep their absorbed energy and do
not return to longitudinal alignment. Therefore, the longitudinal magnetic vector
component remains unaffected. The time constant of this mechanism is dependent
on tissue density and rigidity and is called spin–spin time constant T 2 .
Both time constants are valuable for tissue discrimination and hence are used for imaging. Therefore, in addition to the initial RF pulse to tip the macroscopic magnetization
to a flip angle, measurement of both time constants requires specific sequences of RF
pulses differing in their duration and time interval (e.g., inversion recovery sequence to
measure T 1 or spin echo sequence for T 2 ). To acquire sufficient independent information
for image generation, this complex structure of RF pulse sequences is rapidly repeated.
Unfortunately, the receiving coil is measuring a signal originating from an intracorporal volume that is not capable of localizing the exact signal origin. To allow this, two
additional measures are needed: (1) the application of a longitudinal static magnetic
field gradient B(z). This assures that only nuclei within a small cross-sectional slice are
able to absorb the RF energy, which fulfill the condition that their Larmor frequency
ω L = γ ⋅ B(z) equals the RF frequency ω RF . (2) Spatial coding is performed by applying
transverse static magnetic field gradients B(x) and B(y), which code the site of spinning nuclei by different Larmor frequencies and hence widen the original monofrequent
measured signal into a broadband signal. Consequently, frequency analysis of the measured broadband signal allows identifying the integral signal components originating
from all nuclei along a line in the cross section, which exhibits the same Larmor frequency. From a series of such integral projections of different orientations, a numerical
reconstruction algorithm is able to generate a high-resolution cross-sectional anatomic
image of the body. With selectable weighting of the three basic parameters (hydrogen
density and time constants T 1 and T 2 ), MRI offers many possibilities to display and differentiate tissues with selectable contrast.
However, the gyromagnetic ratio γ does not only depend on the atomic nucleus as
such but also on its type of chemical binding to other nuclei. This results in a chemical
(frequency) shift Δω L , which increases proportional to B, and allows molecular imaging.
To provide this information to medical diagnosis and improve the signal-to-noise ratio,
high static magnetic fields are required.
Electromagnetic Fields in Biological Systems
• Its amplitude is dependent on the number of involved hydrogen nuclei and hence
on the water content of tissue.
• Its time course is mainly influenced by two different mechanisms. The first is the
release of the absorbed RF EMF energy quantum and return of nuclei to their basic
state. Consequently the longitudinal magnetization is recovering at the expense of
the transversal vector component. The time constant of this mechanism is dependent on the tissue structure and is called spin–lattice time constant T 1 .
• The second mechanism responsible for the signal decay is dephasing of the spinning nuclei due to local B-field variations. These are caused by the influence of
neighboring nuclear magnetic moments (spins). Fluctuating local B-fields cause
fluctuating Larmor resonance frequencies and consequently varying phase angles
of nuclear spins. The vectorial summation of increasingly dephased magnetic components results in the ongoing reduction of the macroscopic transverse magnetization. Because nuclei continue spinning, they still keep their absorbed energy and do
not return to longitudinal alignment. Therefore, the longitudinal magnetic vector
component remains unaffected. The time constant of this mechanism is dependent
on tissue density and rigidity and is called spin–spin time constant T 2 .
Both time constants are valuable for tissue discrimination and hence are used for imaging. Therefore, in addition to the initial RF pulse to tip the macroscopic magnetization
to a flip angle, measurement of both time constants requires specific sequences of RF
pulses differing in their duration and time interval (e.g., inversion recovery sequence to
measure T 1 or spin echo sequence for T 2 ). To acquire sufficient independent information
for image generation, this complex structure of RF pulse sequences is rapidly repeated.
Unfortunately, the receiving coil is measuring a signal originating from an intracorporal volume that is not capable of localizing the exact signal origin. To allow this, two
additional measures are needed: (1) the application of a longitudinal static magnetic
field gradient B(z). This assures that only nuclei within a small cross-sectional slice are
able to absorb the RF energy, which fulfill the condition that their Larmor frequency
ω L = γ ⋅ B(z) equals the RF frequency ω RF . (2) Spatial coding is performed by applying
transverse static magnetic field gradients B(x) and B(y), which code the site of spinning nuclei by different Larmor frequencies and hence widen the original monofrequent
measured signal into a broadband signal. Consequently, frequency analysis of the measured broadband signal allows identifying the integral signal components originating
from all nuclei along a line in the cross section, which exhibits the same Larmor frequency. From a series of such integral projections of different orientations, a numerical
reconstruction algorithm is able to generate a high-resolution cross-sectional anatomic
image of the body. With selectable weighting of the three basic parameters (hydrogen
density and time constants T 1 and T 2 ), MRI offers many possibilities to display and differentiate tissues with selectable contrast.
However, the gyromagnetic ratio γ does not only depend on the atomic nucleus as
such but also on its type of chemical binding to other nuclei. This results in a chemical
(frequency) shift Δω L , which increases proportional to B, and allows molecular imaging.
To provide this information to medical diagnosis and improve the signal-to-noise ratio,
high static magnetic fields are required.