Fig. 9.7 Pulsatile inßow through the arteries and non-pulsatile outßow during the cardiac cycle. During systole ocular volume increases associated with an increase in intraocular pressure

Systole Diastole

As with most other techniques for noninvasive blood ßow assessment in humans, the quality of the readings is strongly dependent on clear ocular media. Furthermore, although the software can adjust for small eye movements, good Þxation abilities for the subject under study are needed. Thus, measurements in patients with poor central vision will result in greater variability. As with laser Doppler ßowmetry, the outcome variables are given in arbitrary units. Thus, it is difÞcult to directly compare the data obtained from different sites of measurements or from different subject eyes. Moreover, the technique provides velocity measurements only, although studies have shown that there is a high degree of correlation between measurements of NB and those obtained with hydrogen clearance [37] or the microsphere technique [38]. For measurements in the peripheral human retina, the relative contributions of the choroid and the retina cannot be separated. The use of a confocal system may potentially overcome this problem, but such a system has not yet been realized.

9.1.2Pulsatile Ocular Blood Flow

Another method for the assessment of ocular blood ßow is to measure the changes in ocular volume and pressure during the cardiac cycle. The widespread use of the Goldmann applanation tonometer allows the ophthalmologist to

easily observe the pulsatility of the IOP in daily practice and its variation depending on different physiological and pathological states. It was recognized early that differences in these pressure variations may occur in several ocular pathologies. The Þrst approach to quantify these changes in pulsatility of the IOP was introduced by Perkins [23]. The ocular pulse amplitude was measured using a Goldmann applanation tonometer. In particular, the tonometer tip was modiÞed and connected to a pressure transducer to allow for a continuous transmission of the ocular pulse, which in turn generated a pulse trace. Using these data, pulsatility was compared between the investigated eye and the fellow eye in order to identify carotid insufÞciency [23].

This technique was further developed by Langham and colleagues to assess the pulsatile choroidal blood ßow based on pneumotonometry. The basic principle of this technique is based on the observation that if a bolus of ßuid (as it happens during systole) enters the eye, the eye will expand (Fig. 9.7). Consequently, given the assumption that the eye is elastic and the ßuid is incompressible, the eye will increase by the volume of the bolus associated with an increase in IOP. When the bolus of ßuid leaves the eye again, this will in turn lead to a decrease in eye volume. Thus, if the relation between IOP change and volume change is known (as it is described by the Friedenwald equation), then pressure measurements can be transformed to blood ßow measurements [34].

Based on these considerations, the pneumotonometric OBF system has been introduced for the measurement of the pulsatile component of choroidal blood ßow. This instrument records the waveform of the ocular pulse amplitude over time. This commercially available blood ßow system (Paradigm Medical Industries, Inc., Salt Lake City, UT, Fig. 9.8) assesses changes in intraocular pressure, which are caused by the rhythmic Þlling of the intraocular vessels, with a pneumatic applanation tonometer as described above. Hereby, the maximum IOP change during the cardiac cycle is called pulse amplitude. Based on the pulse amplitude, the wave form, and the heart rate, the pulsatile component of choroidal blood ßow can be calculated.

These calculations are based on a theoretical model eye, which is dependent on several assumptions [34]. First, the measurement of the IOP and its variation over time need to be valid. Second, the relationship between volume and pressure in the eye is taken to be known so that changes in pressure can be converted into changes of volume. This means that the ocular rigidity, which is used to derive ocular volume changes from changes in IOP, is assumed to be equal for all subjects. Third, the hydrodynamic model is based upon the assumption that venous outßow from the eye is nonpulsatile.

An advantage of the pneumotonometric system is that pulsatile ocular blood ßow and IOP can be measured at the same time. It has been shown that the amplitude of the ocular pulse is inßuenced to a major extent by the axial length, and hence, the refraction of the eye [14]. This strong dependence of the pulse amplitude on the eye length can be explained by the differences in intraocular volume between smaller and larger eyes and is not necessarily related to reduced choroidal blood ßow or altered ocular rigidity in myopic subjects [3]. A major limitation of all pulsatile ocular blood ßow techniques is that only the pulsatile component of the choroidal circulation is assessed and no information on the steady component of ocular blood ßow is obtained. Estimates of the pulsatile component of ßow in comparison to total blood ßow in the choroid vary between approximately 80% and 50% [18, 19]. Another major limitation derives from the conversion of the IOP change over the cardiac cycle to volume changes over the cardiac cycle. As mentioned above, the ocular rigidity is assumed to be equal in all measured subjects. However, experimental evidence indicates that ocular rigidity may considerably vary among subjects [6]. In diseases like age-related macular degeneration [8] and glaucoma [12], evidence has accumulated that sclera stiffness is increased.

Fundus pulsation amplitude

Diastole

Systole

Fig. 9.9 Schematic drawing showing the pathways of light for the measurement of ocular fundus pulsations. The eye is illuminated by a parallel laser beam. The light is reßected at both the front side of the cornea (spherical wave) and the retina (plane wave). The waves form interference fringes

9.1.2.1 Laser Interferometry

Information regarding pulsatile choroidal blood ßow can also be obtained by measuring distance changes between cornea and retina during the cardiac cycle [32, 33]. Given that the cornealretinal distance changes with the periodic Þlling of the intraocular vessels, the distance change between cornea and retina contains information of the pulsatile component of choroidal blood ßow. The ocular volume increases when the arterial inßow exceeds venous outßow during systole. This increase can be mainly attributed to choroidal swelling during systole and leads to a reduction of the distance between cornea and retina. This phenomenon is called ocular fundus pulsation, and the maximum distance change during the cardiac cycle is named fundus pulsation amplitude (FPA).

Technically, the eye is illuminated by the parallel beam of a single mode laser diode with a wavelength of 783 nm. This laser light with a high temporal and spatial coherence is then reßected by both the front side of the cornea and the ocular fundus. The wave reßected at the front side of the cornea is spherical, whereas the wave reßected from the retina is plane (Fig. 9.9). Hence, the two waves produce nonlocalized circular interference fringes, which is then subject of further analysis. The path difference between the two waves is twice the optical length of the eye. The interferences produced from the two reemitted waves are detected via a CCD

camera, placed in the center of the interference fringes. Each readout is then plotted along the time axis. Based on these so-called synthetic interferograms, the distance changes between cornea and retina can be evaluated by simply counting the fringes moving inward and outward during the cardiac cycle.

The beam focus on the retina has a diameter of approximately 50Ð100 mm. The reßection from the posterior segment most likely occurs from the retinal pigment epithelium or BruchÕs membrane [7]. Only the interferogram resulting from the strongest reßection of the posterior pole is visible. Whereas other interferogram systems may also arise from other retinal interfaces, they are not visible because of the low intensities. The main reßection from the cornea occurs at the front site of the tear Þlm [11].

The system comprises a fundus camera, allowing for the real-time inspection of the ocular fundus (Fig. 9.10). Thus, measurements can be performed in predeÞned areas of the ocular fundus. However, given that the two reßected laser beams have to be superimposed, measurements are restricted to an area approximately 25¡ around the macula. In practice, measurements of FPA have been performed in the macular region and in the region of the optic nerve head. FPA in the macula is in the order of 4 mm in healthy young subjects, whereas in the optic nerve head region, FPA reaches up to 10 mm [32]. The higher FPA in the optic nerve head has been explained by the increased elastic properties in the optic nerve head region of the eye [30]. In the peripheral retina, FPA is slightly lower than in the macula [32].

Several studies indicate that FPA measurements provide high reproducibility and sensitivity in healthy subjects as well as in patients [25, 31, 33]. However, given that currently no gold standard exists regarding the measurement of choroidal blood ßow, estimation of validity of the method is more difÞcult. Studies comparing laser interferometric measurements with pneumotonic measurement of pulsatile ocular blood ßow reveal a strong correlation between FPA and POBF, indicating that FPA is a valid index of pulsatile choroidal blood ßow in humans [30].