The technique first used to measure cerebral blood flow employed an inert tracer (nitrous oxide) and was described by Kety and Schmidt (1945). Briefly, N2O was inhaled until a
saturation point is reached at which point inhalation ceases and the concentration of the tracer declines. The direct Fick equation was applied via sampling concentrations of the tracer obtained from arterial and jugular samples. Subsequent tracers included Xe133 (Obrist et al. 1967; Thomas et al. 1979) and Kr 85 (Ingvar & Lassen 1961; Lassen & Klee 1965) with the use of extracranial gamma detectors. Whilst this technique did provide an accurate
73 account of CBF, it was invasive, had poor temporal resolution with minutes to acquire a given flow that represented global increases in CBF, and required a steady-state condition (Willie et al. 2011). TCD offers excellent temporal resolution and can be used to track dynamic changes in CBF during a variety of perturbations.
3.1.1 Transcranial Doppler Ultrasound: Ultrasonography Principles
The use of sonography in measuring flow velocity in intracranial cerebral arteries is severely limited by the skull, in that bone attenuates the ultrasonic wave. However, at lower frequencies (1-2 MHz) this attenuation from bone, as well as soft tissue, is markedly less (Aaslid et al. 1982). Further the skull itself varies in thickness, some areas such as the temporal window, just above the zygomatic arch, provide further opportunity for ultrasonic waves to penetrate as the skull is thinner in this area. This window allows the insonation of the large basal cerebral arteries including the middle cerebral, the proximal anterior cerebral, and posterior cerebral arteries (Aaslid et al. 1982; DeWitt & Wechsler 1988). A technique using a lower emitted ultrasound frequency and these anatomical ‘windows’, coined TCD, was first described by Aaslid et al. (1982), and allows real time measurement of cerebral blood flow velocity.
TCD functions via traditional Doppler methods but differs in the fact that both the sound source and the observer (the transducer) are fixed in the same position. The transducer emits sound waves that are reflected by the moving erythrocytes within the insonated vessel, which are detected by the transducer. The Doppler shift resulting from the reflected waves is proportional to the velocity of blood (DeWitt & Wechsler 1988; Stroobant & Vingerhoets 2000; Willie et al. 2011). The following equation describes the Doppler shift (Moppett & Mahajan 2004):
74 Doppler frequency shift = ʹ ൈ ܸ ൈ ܨݐ ൈ ܿݏߠȀܥ
where: V is the velocity of the reflector, which in this case is the erythrocytes within the insonated vessel; Ft is the transmission frequency of the Doppler transducer; C is the speed of sound in soft tissue and cosߠ is a correction factor based on the insonation angle ሺߠ). In pulsed Doppler the Ft is equivalent to 2MHz and C 1540 m·s-1, both of which remain constant. The remaining factors, the angle of insonation and flow velocity of blood, therefore have the largest influence on the resultant Doppler frequency shift (Moppett & Mahajan 2004). If the insonation angle varies from 0 to 30o the resulting cosine will vary between 1 and 0.86, respectively, and the maximum error associated within this acute angle insonation range will be 15% (Aaslid et al. 1982).
As the TCD transducer emits a pulsed signal, the time interval between the emission and receiving of the signal will determine the depth of the Doppler frequency shift. Thus the depth of an insonated vessel can be manipulated by varying the time interval between the emitted and received signal (Moppett & Mahajan 2004). The obtained signal represents a distribution of velocities (Willie et al. 2011). This is due to the varying velocities of erythrocytes within the different lamina of a large artery with the fastest toward the middle of the artery and the slowest next to the vessel wall (Levick 2010). The TCD unit applies a spectral analysis to the mixed frequency shifts to enable three dimensional data to be displayed in two dimensions, with time on the x axis and the frequency shift (velocity) to be displayed on the y axis, and the signal intensity represented by the colour of the trace (Moppett & Mahajan 2004).
75 3.1.2 Validity
With the drastic improvements of temporal resolution in TCD, dynamic beat-to-beat fluctuations in CBF can now be investigated, however, as the measurement obtained is a velocity rather than an absolute flow measure, several problems can arise from this when trying to infer absolute flow. Using this type of technology, the artery diameter cannot be established and therefore is assumed to stay constant (Willie et al. 2011). As reviewed in
Chapter Two (section 2.1), Poiseuille’s law states resistance is inversely proportional to the
radius of a vessel to the fourth power. Furthermore, as velocity can be defined as the flow in a vessel divided by its area (in cm2) any change in vessel diameter will have profound effects on both flow and velocity of the blood within a given artery.
For the larger basal conduit arteries of the brain the vessel diameter has been found to be constant over varying physiological conditions. Direct observations of the large cerebral arteries of the brain (carotid, middle cerebral and vertebral arteries) revealed that during moderate pharmacological alterations in blood pressure and end-tidal PCO2 the average
change was less than 4% (Giller et al. 1993). The smaller arteries such as the anterior cerebral and more distal division of the middle cerebral arteries showed large diameter changes to the aforementioned stimuli (Giller et al. 1993). This was also found to be true in conscious humans using magnetic resonance imaging during alterations in arterial CO2
(Valdueza et al. 1997) and during sympathetic nervous system activation, without a concomitant drop in blood pressure, induced via LBNP (Serrador et al. 2000). These results were also supported by a good correlation between the respective percentage changes when comparing TCD and Xenon (Xe133) clearance technique during hypercapnia (Bishop et al. 1986). Further, using digital angiography it was shown that the cerebral arteries ≥0.57
76 mm in diameter (significantly smaller than the MCA) showed no alteration in calibre to standardised changes in arterial PCO2 (Djurberg et al. 1998). It should be noted in the study
of Djurberg et al. that these patients had arteriovenous malformations (AVM), despite diameter measurements being made on the ispilateral hemisphere to the AVM, the AVM was outside the vascular territory and therefore artery structure and function was assumed to be normal. Signal power measurements from transcranial Doppler also demonstrate no change in MCA diameter during carotid artery compression (Aaslid et al. 1991) and during blood pressure reduction via thigh-cuff deflation (Aaslid et al. 1989). Therefore, it appears that over a wide range of physiological stimuli both with and without anaesthesia, the MCA diameter does not appear to change.
TCD has been validated against the intravenous Xe133 technique at rest and during hypercapnia (Bishop et al. 1986). However, only the change in MCAv correlated suitably with changes in hemispheric blood flow, with absolute velocity correlating poorly with absolute flow. Further, changes in flow within the internal carotid artery were tracked similarly by the change in velocity within the MCA (Lindegaard et al. 1987; Newell et al. 1994). Similarly, during changes in MCAv as a result of pharmacological manipulations of MAP, MCAv tracked relative changes in global CBF (Fick, Xe133 clearance) adequately down to and including the lower limits of cerebral autoregulation (Larsen et al. 1994). Thus, whilst TCD cannot provide a measure of absolute flow the relative changes can be accurately measured and correlate well with absolute changes in CBF measured via more direct methods. Accordingly, results from this thesis are presented as the absolute and relative change from baseline measures.
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