Las leches fermentadas deberán ser envasadas con materiales bromatológicamente aptos de conformidad con el presente Código y adecuados para las condiciones de almacenamiento

Magnetic Resonance Imaging (MRI) is based on the natural magnetisation that is induced in the human body when it is placed in the scanner. Specifically it is the signal obtained from the magnetic moment of hydrogen nuclei that forms the basis of MRI. Conventional MRI produces spatial maps of mobile hydrogen protons that are contained mainly in water molecules, providing anatomic details with exquisite resolution (on the order of 1 mm or better) (Gore, 2003).

All MR images used in this thesis were acquired using a 3 Tesla MRI scanner (for further details see Section 4.3). An MRI sequence contains radiofrequency (RF) pulses and gradient pulses which have carefully controlled durations and timings. The gradient fields are produced by three sets of gradient coils, one for each direction (x, y, z), through which large electrical pulses are applied repeatedly in a carefully controlled pulse sequence. Further information on the acquisition of MR images can be obtained elsewhere (e.g. Buxton 2002; Hashemi et al., 2004; Horowitz, 1995; Jezzard et al., 2001; Schild, 1990; Westbrook and Roth, 2005).

MRI pulse sequences

Three characteristics of the tissue being measured which influence the signal intensity of MR images are the T1 relaxation time, T2 relaxation time and proton density (PD). There are many different types of pulse sequence, but they all have timing values called TR (repetition time) and TE (echo time) which can be modified. The TR is the time between RF pulses and, for a given T1, determines the amount of longitudinal relaxation. The TE is the time between application of an RF pulse and measurement of the MR signal and, for a given T2 determines the amount of transversal relaxation.

Contrast in a T1-weighted image results from differences in longitudinal relaxation times between tissues and structures. A pulse sequence with a short TR (e.g. 300-800 milliseconds) and a short TE (e.g. ~20 milliseconds) will accentuate the effects of

- 55 -

longitudinal relaxation and reduce the loss of magnetization that occurs from T2 dephasing. For T2-weighted images a pulse sequence with a long TR (e.g. ≥1 second) and a longer TE (e.g. 100-500 milliseconds) are used. This will ensure no T1-weighting is present in the signal of interest and will exploit differences in T2 relaxation times of the tissues.

Spin echo pulse sequence

Two factors influence transversal relaxation over time: loss of phase, and inhomogenieties in the magnetic field. The loss of signal can be reduced by applying a 180º refocusing RF pulse a short time TE/2 after the 90º RF pulse. This in effect, causes the precessing protons to turn around resulting in phase coherence and a stronger transversal magnetization. Many 180º RF pulses can be applied to ‘neutralise’ effects that influence the protons in a constant manner.

Gradient Echo Pulse Sequence

The gradient echo (GRE) pulse sequence is used to reduce scan time. Instead of using a 180º refocusing pulse the GRE pulse sequence uses a magnetic field gradient to refocus the FID signal at the end of each TR, by reversing the polarity of the frequency- encoding gradient. The TR is generally the most time consuming parameter in a pulse sequence. The GRE sequence reduces this time by using a smaller flip angle of less than 90º to convert only a fraction of the longitudinal magnetisation into the transverse plane, meaning that a portion of the longitudinal magnetization will remain for the subsequent RF pulse to excite (McRobbie et al., 2003). By applying RF pulses at short TRs, the time it takes for longitudinal magnetization to recover is decreased and an ideal T1- weighted contrast can be achieved in a relatively short amount of time. However, the omission of the refocusing 180º RF pulse means that the dephasing of spins resulting from magnetic field inhomogenieties are not rephased and thus GRE sequences are more susceptible to artefacts. Quite often the standard GRE sequence is modified to obtain T1–weighted MR images. All T1-weighted MR images analysed in this thesis were obtained using a GRE pulse sequence.

- 56 - 3.2 FUNCTIONAL MR IMAGING The BOLD Signal

fMRI detects the blood-oxygenated-level-dependent (BOLD) changes in the MRI signal which result from an increase in neuronal activity in a region of cortex following a change in brain state, which may be produced by a stimulus or task. The BOLD technique is based on the fact that neural activity and haemodynamics (regulation of blood flow and oxygenation) are linked in the brain (Heeger and Ress, 2002; Ogawa et al., 1992). BOLD fMRI reveals which parts of the brain are active in certain tasks with a spatial resolution of 2-5 millimetres.

An increase in neural activity stimulates an increase in the local blood flow in order to meet the larger demand for oxygen and other substrates. The BOLD fMRI technique measures changes in the inhomogeneity of the magnetic field, which are the result of changes in the level of oxygen present in the blood (blood oxygen) (Aguire et al., 2002; Detre and Wang, 2002; Heeger et al., 2002; Ogawa et al., 1990, 1992). While blood that contains oxyhaemoglobin is not very different in terms of susceptibility from other tissues or water, deoxyhaemoglobin is significantly paramagnetic (like the agents used for MRI contrast materials such as gadolinium) and thus deoxygenated blood differs substantially in its magnetic properties from surrounding tissues. Therefore, a high level of deoxyhaemoglobin in the blood will result in a greater field inhomogeneity and therefore a decrease in the fMRI signal (Ogawa et al., 1990).

The haemodynamic response function (HRF)

The function of the BOLD fMRI signal against time in response to a temporary increase in neuronal activity is known as the haemodynamic response function (HRF) (Heeger et al., 2002). After an increase in neuronal activity there is an increase in the relative level of deoxyhaemoglobin in the blood as active neurons use oxygen, resulting in a decrease of the signal (Heeger et al., 2002; Vanzetta and Grinvald, 1999). The decrease however, is tiny and is not always found (Detre and Wang, 2002; Ugurbil et al., 2003). Following this initial decrease, there is a large increase in the BOLD fMRI signal which reaches its maximum after approximately 6 seconds, due to a massive oversupply of oxygen rich blood (Fox et al., 1988; Heeger et al., 2002). The result of this oversupply of oxygen is a large decrease in the relative level of deoxyhaemoglobin, which in turn causes the

- 57 -

increase in the BOLD fMRI signal. Finally, the level of deoxyhaemoglobin slowly returns to normal and the BOLD fMRI signal decays until it has reached its original baseline after an initial undershoot after approximately 24 seconds (Heeger et al., 2002). Further information on the signal obtained in fMRI can be found elsewhere (Gore, 2003).

fMRI signal of interest

Block design (Aguirre and D’Esposito, 2000; Donaldson and Buckner, 2001) is the most commonly used experimental design in neuroimaging, and is the design used for all fMRI tasks in this thesis. Two or more conditions are alternated in blocks. The so- called subtraction paradigm involves making the conditions in each block differ in only the cognitive process of interest (Aguirre and D’Esposito, 2000; Donaldson and Buckner, 2001). The fMRI signal that differentiates the conditions should represent the cognitive process of interest. The main advantage of block design is that the increase in fMRI signal in response to a stimulus is additive, meaning that the amplitude of the HRF increases when multiple stimuli are presented in rapid succession. When each block is alternated with a rest condition in which the HRF has enough time to return to baseline and a maximum amount of variability is introduced in the signal. Therefore, block designs offer considerable statistical power.