- Email: [email protected]
Laser Doppler instrument to investigate retinal neural activity-induced changes in optic nerve head blood flow
ARTICLE IN PRESS
Optics and Lasers in Engineering 43 (2005) 591–602
Laser Doppler instrument to investigate retinal neural activity-induced changes in optic nerve head blood flow Eric Logeana,*, Martial H. Geiserb,a, Charles E. Rivaa a
Institut de Recherche en Ophtalmologie, CH-1950, Sion, Switzerland b Haute Ecole Valaisanne, CH-1950, Sion, Switzerland
Received 9 December 2003; received in revised form 31 March 2004; accepted 5 April 2004 Available online 20 July 2004
Abstract A fundus camera-based instrument to investigate increased retinal neuronal activity effect on optic nerve blood flow is described. It incorporates (a) near infrared fundus illumination and observation, (b) near infrared laser Doppler flowmeter (LDF), (c) pupil position monitoring and (d) delivery of visual stimuli. Two types of stimulation are currently used: diffuse heterochromatic and chromatic luminance flicker generated by light emitting diodes, and contrast reversal pattern displayed on a video monitor. Recordings of the changes in the LDF parameters (blood velocity, volume and flow) from the optic nerve in response to these stimuli illustrate the potential of the technique. r 2004 Elsevier Ltd. All rights reserved. Keywords: Retinal neural activity; Ocular circulation; Laser Doppler flowmetry; Optic nerve head; Flicker stimulation
1. Introduction In 1890, Roy and Sherrington [1], both neurophysiologists, postulated that in the brain neuronal activity, cellular metabolism and blood flow are tightly coupled. This coupling is the basis of all functional imaging techniques, whereby local changes in brain activity can be visualized by monitoring the changes in blood flow, glucose *Corresponding author. Tel.: +41-27-205-7900; fax: +41-27-205-7901. E-mail address: [email protected] (E. Logean). 0143-8166/$ - see front matter r 2004 Elsevier Ltd. All rights reserved. doi:10.1016/j.optlaseng.2004.04.011
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utilization, or oxygen consumption associated with the activity of specific neuronal circuits [2,3]. Over the past several years, laser Doppler flowmetry (LDF) has been used in animal models to characterize the hemodynamic changes that occur in the brain in response to functional activation [4]. The technique has shown its capability to document in animal models the spatio-temporal properties of the activity-induced flow changes [2]. LDF studies in the eye have demonstrated that optic nerve head (ONH) blood flow in the cat also responds to increased neuronal activity induced by visual stimulation [5,6]. Such activity-induced flow-responses were obtained recently in human subjects [7,8]. In this paper, we describe the instrument used to measure the response of flow to increased neuronal activity. However, prior to that we provide a brief review of the anatomy of the eye fundus and of the principle of LDF as applied to the measurement of blood flow.
2. The fundus of the eye A section cut of the human eye shows the different fundus layers, i.e. the sclera, the choroid and the retina (Fig. 1). The outermost layer, the sclera, wraps around the eye globe as a protecting shell. Between the retina, the innermost fundus layer, and the
choroid
cornea retina
id oro s ch tor ep rec orm xif oto ph ple lear ter uc m ou r rn o e f i inn lex s p ell er nc inn lio ng ga
sclera
pupil optic disc
central retinal artery and vein
lens
optic nerve iris
ciliary artery and vein
lamina cribrosa
Fig. 1. Diagram of a human eye. Light enters the eye through the pupil, a circular aperture formed by the iris. The cornea and the lens form an image on the fundus. The latter is composed of three layers: the sclera, choroid and retina. The inner retinal structures are detailed on the left. (Adapted with permission from Kolb [10])
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∆f (Hz) Fig. 2. (a) View of the ONH and retinal vessels. The black circle represents the probing beam at a typical recording site. (b) Typical Doppler shift power spectrum P(Df) obtained from the ONH. The analysis software removes data that are outside the analysis range, here below 30 Hz and above 2500 Hz (vertical lines).
sclera lies the choroid, a highly vascularized tissue, whose main function is to supply oxygen to the photoreceptors located in the outer part of the retina. The retinal photoreceptors convert light into electrical signals that are preprocessed by the different retinal layers and the ganglion cells. The visual information is then transmitted to the brain by the ganglion cell axons, first in the inner (neural) part of the retina and then by the optic nerve. The vascular system of the eye is derived from the ophthalmic artery. Branching from this artery, the ciliary arteries supply the choroid and the central retinal artery gives rise to main retinal arterioles which feed the inner (neural) layers of the retina. Superficial layers of the optic nerve, visible in Fig. 2(a), are fed by the central retinal artery whereas the pre-laminar and laminar layers are supplied by posterior ciliary arteries [9].
3. Laser Doppler flowmetry (LDF) LDF is based on the Doppler effect which describes the change in frequency of a sound or a light wave emitted from a moving source when the source is traveling towards or away from the observer. With the advent of the laser it became possible to detect very sensitively the Doppler shift of optical waves and measure flow velocities in a non-invasive way over a broad range of velocities. Consider a laser beam incident on a particle, such as a red blood cell (RBC), along ~ i : Light scattered towards a photodetector a direction defined by the wave vector K ~ along the direction K s is shifted in frequency by an amount Df, the Doppler shift ~i K ~ s and V ~ ; the velocity vector of frequency, equal to the scalar product between K
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the particle Df ¼
1 ~ ~s Þ V ~ ðK i K 2p
ð1Þ
~ i jDjK ~ s j ¼ 2pn=l; where n is the refractive index of the flowing medium and l the jK wavelength of the incident laser light in vacuo. 3.1. Doppler shift power spectrum for RBCs moving in a tissue When a laser beam impinges on a tissue, most photons reaching the photodetector have followed random paths through the tissue due to scattering by the static tissue structures and some of them have also interacted with RBCs moving with different velocities and directions. This distribution of velocities and directions results in a Doppler shift power spectrum [P(Df)], which has a shape close to that of an exponentially decreasing function. An example of P(Df) is shown in Fig. 2(b). P(Df) is recorded by heterodyne light mixing spectroscopy [11], a technique that makes use of the interferences between the un-shifted light scattered by the nonmoving tissue and the light (Doppler shifted) that has been scattered by the moving RBCs to generate a photocurrent whose spectrum contains only the Doppler shift beat frequencies Df’s. The application of LDF to the measurement of RBC velocity in the ONH tissue was first described in 1982 [12]. 3.2. The hemodynamic quantities derived from the spectrum Based on Bonner and Nossal’s theory of LDF [13], three parameters are derived from P(Df): the velocity (Vel) which is the first moment frequency of the P(Df) and is expressed in Hz, Vel is proportional to the mean velocity of the RBCs moving in the sampling volume; the volume (Vol) which is the area under P(Df) and is proportional to the number of moving RBCs in this volume; the flow F=k Vel Vol , where k is a constant of proportionality. These parameters are mathematically defined as follows: R max Df PðDf Þ dDf R max Vel ¼ min ; ð2Þ min P ðDf Þ dDf Vol ¼
F¼
1 A2DC
1 A2DC
Z
Z
max
PðDf Þ dDf ;
ð3Þ
Df PðDf Þ dDf ;
ð4Þ
min max
min
where ADC is the amplitude of the direct current which is proportional to the intensity of the detected light. Doppler shifts below 30 Hz (min) are filtered out in order to minimize the effect of tissue motion artifacts. When measuring from the ONH in humans, the Doppler shift frequencies are typically below 2500 Hz (max).
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Vol and F are expressed in relative units since only relative blood volume and blood flow can be measured with the LDF technique. Effectively, the penetration depth of the light, which determines the probed volume, depends upon light scattering and absorption properties of the tissue. As these optical properties generally vary spatially due to variations of blood volume, hemoglobin concentration and saturation and tissue structure, the penetration depth also varies spatially. Furthermore, local alterations in tissue structure or vascular arrangement occurring in various pathologies, will also affect the flow readings.
4. The optical setup The laser Doppler optical delivery and detection systems were mounted on the Kowa-Pro 50 fundus camera (Kowa-Pro, Kowa Co. Ltd., Japan), which was modified, as shown schematically in Fig. 3. The light source (lamps and associated optics) of the camera was replaced by 12 light emitting diodes (LED) radiating at 770720 nm (LED770-03AU, Roitner Lasertechnik, Austria). This near infrared illumination IR-I is mounted in an image plane of the eye pupil. The light from these diodes is reflected into the illumination path of the fundus camera by a hot mirror HM. Following this path, it is then reflected by the annular mirror at P1, focused in the plane of the eye pupil P by the ophthalmic lens OL and illuminates the eye fundus in Maxwellian view over a field of 50 . The radiance of this illumination at the cornea is 5 mW cm2 sr1, a level below the maximum permissible exposure [14]. The light reflected by the fundus structures that exits the pupil of the eye forms an image of the fundus at the plane R2 after passing through OL, the circular opening of the annular mirror P1 and the optical system between P1 and R2. 4.1. Laser delivery to the optic nerve A near-infrared beam lasing at 810 nm is focused at the optic nerve tissue, in areas deprived of visible vessels, as indicated in Fig. 2(a) using part of the optics of the fundus illumination of the Kowa camera. The laser diode LD (Pion III, Micro Laser System, USA) is mounted off-center on a rotating wheel (Fig. 3). This system enables the beam to enter the eye at the desired location. The off-axis beam is then deflected by an optical element, BP, consisting of two prisms mounted in an image plane of the eye pupil. Rotation of the element and of the two prisms relatively to each other allows the aiming of the beam at the desired fundus location. Lens LF focuses the beam onto a retinal image plane R3 and the 50/50 beam splitter, BS, directs this beam along the optical path of the fundus illumination pathway of the fundus camera. The power of the probing beam at the cornea is approximately 100 mW, a level below the maximum permissible exposure [14]. Choosing a camera fundus illumination/observation and laser Doppler probing beam in the near infrared range of the spectrum prevents these lights to perturb the level of light adaptation of the subject’s retina.
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Fig. 3. Schematic representation of the fundus camera, laser Doppler system and stimulation light sources (See text for details) with R: retinal plane; P: pupil of the tested eye; R1, R2, R3: retinal images; P1: pupil images; DM: dichroic mirror; SL: spectacle lens; OL: ophthalmic lens; APD: avalanche photodiode; CCD: charge coupled device; BS: beam splitter; HM: hot mirror; IR-I: infrared illumination; LF: focussing lens; BP: bi-prism; LD: laser diode. Inset at bottom left shows the pupil monitoring system consisting of: PI: pupil illumination; M: full mirror; L1: lens; IR-F: infrared filter; P2: pupil image.
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4.2. Detection of the light scattered by the RBCs In the retinal image plane R2 (Fig. 3), an optical fiber (200-mm diameter at the fundus) mounted on a x–y positioning stage collects the light from the site illuminated by the laser. An avalanche photodiode APD (C30902S, EG&G, USA) converts this light in electric current, which is amplified, digitized and fed to the digital signal processor of a NeXT computer (Red-wood city, USA) for analysis by dedicated software [15]. The three LDF-parameters are obtained at a rate of 21 Hz. The collecting fiber and eye fundus are simultaneously observed by a CCD camera (C2400-02, Hamamatsu Photonics K.K., Japan) mounted in the plane of a relayed image of R2. 4.3. Pupil illumination and observation Preliminary measurements have demonstrated that the variability of the flickerinduced blood flow response in successive recordings could be reduced by maintaining constant the position of the camera relative to the eye pupil. Therefore, the optical system drawn in Fig. 3 (inset) was designed for monitoring the eye pupil position. The iris of the subject is illuminated with PI, a LED emitting at 870 nm. The optical system of observation consists of the ophthalmic lens OL (used at its edge), a mirror M, which redirects the light out of the fundus camera instrument and a lens L1 of 20 mm of focal length, which forms an image of the iris of the subject on a CCD finger camera (Conrad, Germany). To avoid detecting unwanted light such as the probing beam reflexion at the cornea, a high pass filter IR-F (IR820, Schott, Germany) with cutoff at 820 nm was placed in front of the CCD camera. 4.4. Diffuse stimulation Diffuse stimulation is generated by six red and six green LEDs (SL905 OCU-16 and GCU-15, Sloan Precision Optoelectronics, USA) located in an image of the eye pupil (Fig. 3). These LEDs are driven by a generator based on the pulse wave modulation technique [16] and produce user defined waveform stimuli. Generally, luminance stimuli (i.e. green/black flicker) or chromatic stimuli (i.e. constant luminance red/green flicker) are used. These LEDs uniformly illuminate a field of 50 at the fundus with a maximum mean illuminance of 40 lx for each color. 4.5. Contrast reversal pattern stimulation Contrast reversal pattern (checkerboard or grating) were displayed on the monitor of a pattern generating system (VGS series 5, Cambridge Research System, England). These stimuli were reflected by a dichroic mirror DM in Fig. 3 (E30634, Edmund Industrial Optics, USA). This mirror, which reflects the visible light and transmits the near infrared light is located between the ophthalmic lens OL and the eye. After reflection at DM, the stimuli were focused at the retina using a
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spectacle lens SL with a dioptric power specific for the tested eye. Such pattern stimulus sustains a field of approximately 30 with a mean illuminance of 4.5 lx.
5. Experimental part When recording activity-induced blood flow responses, subjects were asked to fixate a pinpoint target located in the plane R1 (Fig. 3). The target was the tip of a 50-mm core diameter optical fiber transmitting light from a green LED. The luminance of the fixation target was chosen to match that of the mean retinal luminance of the stimulus. After locating the optic disk on the monitor screen and bringing the probing beam in the rim region using the bi-prisms BP, the operator aimed the beam at the precisely chosen location by moving the fixation target. Furthermore, during the measurements, the operator monitored the position of the pupil of the tested eye relative to the fundus camera and attempted to keep it constant by readjusting the camera, when necessary. A typical recording protocol consisted of a pre-stimulus period of approximately 30 s followed by a 40–50 s period of stimulation of identical mean luminance and a 30 s recovery period. Prior to the measurements, the tested eye was dilated with two drops of Tropicamide 0.5% (Novartis Ophthalmics AG, Switzerland). 5.1. Blood flow response to diffuse stimulation Fig. 4a shows a typical set of recordings of the blood flow parameters obtained in a normal volunteer in response to a 15-Hz green luminance flicker. The fast
(a)
(b)
Fig. 4. (a) Typical recording obtained from the temporal side of the neuro-retinal rim of the ONH of a normal volunteer. The stimulation was a 50 field diffuse luminance flicker (green/black) at 15 Hz. (b) Blood flow response RF to stimuli of different color ratio r at 3.2 Hz. Error bars represent standard deviations of the mean of 6 recordings.
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variations are due to the changes of the flow parameters during the heart cycle, as confirmed by the recording of the heart pulse using a plethysmograph connected to one of the ear of the subject. In response to the flicker, all three parameters increase to reach a plateau shortly after the start of the stimulation and then decrease back to the baseline following the cessation of the stimulus. For this particular recordings, the percent changes between the last 30 s of stimulation and the pre-stimulation were 19%, 3% and 23% for Vel, Vol and F, respectively. The time course and amplitude of these changes obtained in 15 normal volunteer have been recently described by Riva et al. [17]. Such recordings may be repeated several times varying the characteristics of the stimulus, such as the mean luminance, modulation depth, temporal frequency, and color ratio. The amplitude of the blood flow response (RF) expressed as percent changes from the pre-stimulus F-level was measured in response to 3.2-Hz stimuli of different red/ green luminance ratio (r=red/(red+green)). r=0 corresponds to a green/black luminance flicker, r=1 to a red/black luminance flicker, and r=0.5 to an equiluminant chromatic flicker. Such low temporal frequency chromatic flicker is known to activate predominantly the parvocellular pathway, which is responsibles for retinal color processing [18]. Fig. 4b shows a relatively large RF for a chromatic flicker (r=0.4) compared to luminance flicker (r=0 or 1). 5.2. Blood flow response to pattern stimulation Fig. 5a shows four successive responses to checkerboard pattern stimulation from a subject with excellent target fixation. The pattern was constituted of 36-arcmin black and white checks and was reversed at a rate of 8 Hz. The mean changes in the flow parameters were 9712% (stdev) for Vel, 3276% for Vol and 3876% for F. The dependence of RF on the sharpness of the stimulus image on the retina was measured in one subjects by changing the spectacle lens (SL in the inset of Fig. 3) from 3 to +4.5 diopters with 0 diopter corresponding to a pattern sharply focused on the retina. The pattern consisted in a 225-arcmin black and white vertical grating reversed at a rate of 8 Hz. Fig. 5b shows that RF are obtained for sharply focused pattern stimulus.
6. Discussion We have described a fundus camera-based laser Doppler instrument to investigate the changes in blood flow in the optic nerve of human subjects in response to visual stimulation. The fast response time of this device makes it possible to monitor the time-course of flow response to increased neural activity and fast changes in blood flow, such as those occurring during the heart cycle. To our knowledge, only the LDF-technique has been applied to the monitoring of the optic disk blood flow response to visual stimulation. In the retinal microvasculature, the amplitude of the blood flow response to diffuse luminance flicker has been investigated using the Heidelberg retina flowmeter (HRF).
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(a)
(b) Fig. 5. (a) Continuous recording of Vel, Vol and F obtained in a normal volunteer in response to four stimulation periods (checkerboard, 30 , 8 Hz, 36 arcmin). (b) Effect of pattern defocus on the blood flow responses RF (vertical line grating, 30 , 8 Hz, 225 arcmin).
Michelson et al. [19] and Harrison et al. [20] have found increases in peripapillary blood flow in response to diffuse flicker. Attempts to obtain response to pattern stimuli using the HRF remained, however, unsuccessful [21]. Indirectly, the flow response to diffuse stimuli has been investigated using the blue field entoptic simulation technique [22]. Flicker-induced increases in retinal vessel diameter was also reported [23–25]. Furthermore, the amplitude of the blood velocity changes in the central retinal artery and vein has been measured using ultrasound sonography [19]. In the optic nerve head, as in the brain, increased neuronal activity induces detectable and reproducible increases in local blood flow. This capability opens new avenues in the identification of the factors which regulate blood flow in this neural tissue of the ocular fundus. This instrument should be useful in the study of the role
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of the circulation in the pathogenesis of various ocular diseases of vascular origin, such as glaucoma and diabetic retinopathy.
Acknowledgements . The authors thanks Ce! dric Schopfer for the design and implementation of the diffuse stimulus generator and technical expertise. This work was supported by the Swiss National Science Foundation Grant 32-53785.
References [1] Roy CS, Sherrington CS. On the regulation of the blood supply to the brain. J Physiol 1890;11: 85–108. [2] Villringen A, Dirnagl U. Coupling of brain activity and cerebral blood flow: basis of functional neuroimaging. Cerebrovasc Brain Metab Rev 1995;7:240–76. [3] Grinwald A, Lieke E, Frostig RD, Gilbert CD, Wiesel TN. Functional architecture of cortex revealed by optical imaging of intrinsic signals. Nature 1986;324:361–4. [4] Dirnagl U, Kaplan B, Jacewicz M, Pulsinelli W. Continuous measurement of cerebral cortical blood flow by laser-Doppler flowmetry in a rat stroke model. J Cereb Blood Flow Metab 1989;9:589–96. [5] Riva CE, Harino S, Shonat RD, Petrig BL. Flicker evoked increase in optic nerve head blood flow in anesthetized cats. Neurosci Lett 1991;128(2):291–6. [6] Riva CE, Harino S, Petrig BL, Shonat RD. Laser-Doppler flowmetry in the optic nerve. Exp Eye Res 1992;55:499–506. [7] Riva CE, Falsini B, Logean E. Flicker-evoked changes in human optic nerve blood flow: luminance versus chromatic modulation. Invest Ophthalmol Vis Sci 2001;42(3):756–62. [8] Falsini B, Riva CE, Logean E. Flicker-evoked changes in human optic nerve blood flow: relationship with retinal neural activity. Invest Ophthalmol Vis Sci 2002;43(7):2309–16. [9] Hayreh SS. The 1994 Von Sallman lecture: the optic nerve head circulation in health and disease. Exp Eye Res 1995;61:259–72. [10] Kolb H. How the retina works. Am Scientist 2003;91:28–35. [11] Cummins HZ, Swinney HL. Light beating spectroscopy. In: Wolf E, editor. Progress in optics, vol. VIII. New York: American Elsevier Publishing Co; 1970. p. 133–200. [12] Riva CE, Grunwald JE, Sinclair SH. Laser Doppler measurement of relative blood velocity in the human optic nerve head. Invest Ophthalmol Vis Sci 1982;22(2):241–8. . [13] Bonner RF, Nossal R. Principles of laser-Doppler flowmetry. In: Shepherd AP, Oberg PA, editors. Laser-Doppler blood flowmetry. Boston: Kluwer Academic Publishers; 1990. p. 17–45. [14] American National Standard Institute. American national standard for safe use of lasers: ANSI Z136.1-2000. [15] Petrig BL, Riva CE. Optic nerve head laser Doppler flowmetry: principles and computer analysis. In: Kaiser HJ, Flammer J, Hendrickson P, editors. Ocular blood flow. Basel: Karger; 1996. p. 120–7. [16] Toi VV, Burckhardt W. Visual stimulator monitored by microcontroller. In: Noninvasive assessment of the visual system, Technical digest series. Washington: Optical Society of America; 1992. p. 198–201. [17] Riva CE, Logean E, Falsini B. Temporal dynamics and magnitude of the blood flow response at the optic disk in normal subjects during functional retinal flicker-stimulation. Neurosci Lett 2004;356(2): 75–8. [18] Lee BB, Pokorny J, Smith VC, Martin PR, Valberg A. Luminance and chromatic modulation sensitivity of macaque ganglion cells and human observers. J Opt Soc Am A 1990;7(12):2223–36.
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[19] Michelson G, Patzelt A, Harazny J. Flickering light increases retinal blood flow. Retina 2002;22(3): 336–43. [20] Harrison JM, Elliot WR, Kiel J, Sponsel WE. Effect of 10 Hz flicker on retinal hemodynamics measured by the Heidelberg retina flowmeter. In: Noninvasive assessment of the visual system, Technical digest series. Washington: Optical Society of America; 1997. p. 61–5. [21] Schmeisser ET, Harrison JM, Sutter EE, Kiel J, Elliot RW, Sponsel WE. Modification of the Heidelberg retinal flowmeter to record pattern and flicker induced blood flow changes. Doc Ophthalmol 2003;106:257–63. [22] Scheiner AJ, Riva CE, Kazahaya K, Petrig BL. Effect of flicker on macular blood flow assessed by the blue field simulation technique. Invest Ophthalmol Vis Sci 1994;35(9):3436–41. [23] Formaz F, Riva CE, Geiser M. Diffuse luminance flicker increases retinal vessel diameter in humans. Curr Eye Res 1997;16:1252–7. [24] Polak K, Schmetterer L, Riva CE. Influence of flicker frequency on flicker-induced changes of retinal vessel diameter. Invest Ophthalmol Vis Sci 2002;43(8):2721–6. [25] Nagel E, Vilser W. Flicker observation light induces diameter response in retinal arterioles: a clinical methodological study. Br J Ophthalmol 2004;88:54–6.
Optics and Lasers in Engineering 43 (2005) 591–602
Laser Doppler instrument to investigate retinal neural activity-induced changes in optic nerve head blood flow Eric Logeana,*, Martial H. Geiserb,a, Charles E. Rivaa a
Institut de Recherche en Ophtalmologie, CH-1950, Sion, Switzerland b Haute Ecole Valaisanne, CH-1950, Sion, Switzerland
Received 9 December 2003; received in revised form 31 March 2004; accepted 5 April 2004 Available online 20 July 2004
Abstract A fundus camera-based instrument to investigate increased retinal neuronal activity effect on optic nerve blood flow is described. It incorporates (a) near infrared fundus illumination and observation, (b) near infrared laser Doppler flowmeter (LDF), (c) pupil position monitoring and (d) delivery of visual stimuli. Two types of stimulation are currently used: diffuse heterochromatic and chromatic luminance flicker generated by light emitting diodes, and contrast reversal pattern displayed on a video monitor. Recordings of the changes in the LDF parameters (blood velocity, volume and flow) from the optic nerve in response to these stimuli illustrate the potential of the technique. r 2004 Elsevier Ltd. All rights reserved. Keywords: Retinal neural activity; Ocular circulation; Laser Doppler flowmetry; Optic nerve head; Flicker stimulation
1. Introduction In 1890, Roy and Sherrington [1], both neurophysiologists, postulated that in the brain neuronal activity, cellular metabolism and blood flow are tightly coupled. This coupling is the basis of all functional imaging techniques, whereby local changes in brain activity can be visualized by monitoring the changes in blood flow, glucose *Corresponding author. Tel.: +41-27-205-7900; fax: +41-27-205-7901. E-mail address: [email protected] (E. Logean). 0143-8166/$ - see front matter r 2004 Elsevier Ltd. All rights reserved. doi:10.1016/j.optlaseng.2004.04.011
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utilization, or oxygen consumption associated with the activity of specific neuronal circuits [2,3]. Over the past several years, laser Doppler flowmetry (LDF) has been used in animal models to characterize the hemodynamic changes that occur in the brain in response to functional activation [4]. The technique has shown its capability to document in animal models the spatio-temporal properties of the activity-induced flow changes [2]. LDF studies in the eye have demonstrated that optic nerve head (ONH) blood flow in the cat also responds to increased neuronal activity induced by visual stimulation [5,6]. Such activity-induced flow-responses were obtained recently in human subjects [7,8]. In this paper, we describe the instrument used to measure the response of flow to increased neuronal activity. However, prior to that we provide a brief review of the anatomy of the eye fundus and of the principle of LDF as applied to the measurement of blood flow.
2. The fundus of the eye A section cut of the human eye shows the different fundus layers, i.e. the sclera, the choroid and the retina (Fig. 1). The outermost layer, the sclera, wraps around the eye globe as a protecting shell. Between the retina, the innermost fundus layer, and the
choroid
cornea retina
id oro s ch tor ep rec orm xif oto ph ple lear ter uc m ou r rn o e f i inn lex s p ell er nc inn lio ng ga
sclera
pupil optic disc
central retinal artery and vein
lens
optic nerve iris
ciliary artery and vein
lamina cribrosa
Fig. 1. Diagram of a human eye. Light enters the eye through the pupil, a circular aperture formed by the iris. The cornea and the lens form an image on the fundus. The latter is composed of three layers: the sclera, choroid and retina. The inner retinal structures are detailed on the left. (Adapted with permission from Kolb [10])
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P (a.u.)
0.31
0 (a)
(b)
0
1000
2000
3000
4000
5000
∆f (Hz) Fig. 2. (a) View of the ONH and retinal vessels. The black circle represents the probing beam at a typical recording site. (b) Typical Doppler shift power spectrum P(Df) obtained from the ONH. The analysis software removes data that are outside the analysis range, here below 30 Hz and above 2500 Hz (vertical lines).
sclera lies the choroid, a highly vascularized tissue, whose main function is to supply oxygen to the photoreceptors located in the outer part of the retina. The retinal photoreceptors convert light into electrical signals that are preprocessed by the different retinal layers and the ganglion cells. The visual information is then transmitted to the brain by the ganglion cell axons, first in the inner (neural) part of the retina and then by the optic nerve. The vascular system of the eye is derived from the ophthalmic artery. Branching from this artery, the ciliary arteries supply the choroid and the central retinal artery gives rise to main retinal arterioles which feed the inner (neural) layers of the retina. Superficial layers of the optic nerve, visible in Fig. 2(a), are fed by the central retinal artery whereas the pre-laminar and laminar layers are supplied by posterior ciliary arteries [9].
3. Laser Doppler flowmetry (LDF) LDF is based on the Doppler effect which describes the change in frequency of a sound or a light wave emitted from a moving source when the source is traveling towards or away from the observer. With the advent of the laser it became possible to detect very sensitively the Doppler shift of optical waves and measure flow velocities in a non-invasive way over a broad range of velocities. Consider a laser beam incident on a particle, such as a red blood cell (RBC), along ~ i : Light scattered towards a photodetector a direction defined by the wave vector K ~ along the direction K s is shifted in frequency by an amount Df, the Doppler shift ~i K ~ s and V ~ ; the velocity vector of frequency, equal to the scalar product between K
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the particle Df ¼
1 ~ ~s Þ V ~ ðK i K 2p
ð1Þ
~ i jDjK ~ s j ¼ 2pn=l; where n is the refractive index of the flowing medium and l the jK wavelength of the incident laser light in vacuo. 3.1. Doppler shift power spectrum for RBCs moving in a tissue When a laser beam impinges on a tissue, most photons reaching the photodetector have followed random paths through the tissue due to scattering by the static tissue structures and some of them have also interacted with RBCs moving with different velocities and directions. This distribution of velocities and directions results in a Doppler shift power spectrum [P(Df)], which has a shape close to that of an exponentially decreasing function. An example of P(Df) is shown in Fig. 2(b). P(Df) is recorded by heterodyne light mixing spectroscopy [11], a technique that makes use of the interferences between the un-shifted light scattered by the nonmoving tissue and the light (Doppler shifted) that has been scattered by the moving RBCs to generate a photocurrent whose spectrum contains only the Doppler shift beat frequencies Df’s. The application of LDF to the measurement of RBC velocity in the ONH tissue was first described in 1982 [12]. 3.2. The hemodynamic quantities derived from the spectrum Based on Bonner and Nossal’s theory of LDF [13], three parameters are derived from P(Df): the velocity (Vel) which is the first moment frequency of the P(Df) and is expressed in Hz, Vel is proportional to the mean velocity of the RBCs moving in the sampling volume; the volume (Vol) which is the area under P(Df) and is proportional to the number of moving RBCs in this volume; the flow F=k Vel Vol , where k is a constant of proportionality. These parameters are mathematically defined as follows: R max Df PðDf Þ dDf R max Vel ¼ min ; ð2Þ min P ðDf Þ dDf Vol ¼
F¼
1 A2DC
1 A2DC
Z
Z
max
PðDf Þ dDf ;
ð3Þ
Df PðDf Þ dDf ;
ð4Þ
min max
min
where ADC is the amplitude of the direct current which is proportional to the intensity of the detected light. Doppler shifts below 30 Hz (min) are filtered out in order to minimize the effect of tissue motion artifacts. When measuring from the ONH in humans, the Doppler shift frequencies are typically below 2500 Hz (max).
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Vol and F are expressed in relative units since only relative blood volume and blood flow can be measured with the LDF technique. Effectively, the penetration depth of the light, which determines the probed volume, depends upon light scattering and absorption properties of the tissue. As these optical properties generally vary spatially due to variations of blood volume, hemoglobin concentration and saturation and tissue structure, the penetration depth also varies spatially. Furthermore, local alterations in tissue structure or vascular arrangement occurring in various pathologies, will also affect the flow readings.
4. The optical setup The laser Doppler optical delivery and detection systems were mounted on the Kowa-Pro 50 fundus camera (Kowa-Pro, Kowa Co. Ltd., Japan), which was modified, as shown schematically in Fig. 3. The light source (lamps and associated optics) of the camera was replaced by 12 light emitting diodes (LED) radiating at 770720 nm (LED770-03AU, Roitner Lasertechnik, Austria). This near infrared illumination IR-I is mounted in an image plane of the eye pupil. The light from these diodes is reflected into the illumination path of the fundus camera by a hot mirror HM. Following this path, it is then reflected by the annular mirror at P1, focused in the plane of the eye pupil P by the ophthalmic lens OL and illuminates the eye fundus in Maxwellian view over a field of 50 . The radiance of this illumination at the cornea is 5 mW cm2 sr1, a level below the maximum permissible exposure [14]. The light reflected by the fundus structures that exits the pupil of the eye forms an image of the fundus at the plane R2 after passing through OL, the circular opening of the annular mirror P1 and the optical system between P1 and R2. 4.1. Laser delivery to the optic nerve A near-infrared beam lasing at 810 nm is focused at the optic nerve tissue, in areas deprived of visible vessels, as indicated in Fig. 2(a) using part of the optics of the fundus illumination of the Kowa camera. The laser diode LD (Pion III, Micro Laser System, USA) is mounted off-center on a rotating wheel (Fig. 3). This system enables the beam to enter the eye at the desired location. The off-axis beam is then deflected by an optical element, BP, consisting of two prisms mounted in an image plane of the eye pupil. Rotation of the element and of the two prisms relatively to each other allows the aiming of the beam at the desired fundus location. Lens LF focuses the beam onto a retinal image plane R3 and the 50/50 beam splitter, BS, directs this beam along the optical path of the fundus illumination pathway of the fundus camera. The power of the probing beam at the cornea is approximately 100 mW, a level below the maximum permissible exposure [14]. Choosing a camera fundus illumination/observation and laser Doppler probing beam in the near infrared range of the spectrum prevents these lights to perturb the level of light adaptation of the subject’s retina.
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Fig. 3. Schematic representation of the fundus camera, laser Doppler system and stimulation light sources (See text for details) with R: retinal plane; P: pupil of the tested eye; R1, R2, R3: retinal images; P1: pupil images; DM: dichroic mirror; SL: spectacle lens; OL: ophthalmic lens; APD: avalanche photodiode; CCD: charge coupled device; BS: beam splitter; HM: hot mirror; IR-I: infrared illumination; LF: focussing lens; BP: bi-prism; LD: laser diode. Inset at bottom left shows the pupil monitoring system consisting of: PI: pupil illumination; M: full mirror; L1: lens; IR-F: infrared filter; P2: pupil image.
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4.2. Detection of the light scattered by the RBCs In the retinal image plane R2 (Fig. 3), an optical fiber (200-mm diameter at the fundus) mounted on a x–y positioning stage collects the light from the site illuminated by the laser. An avalanche photodiode APD (C30902S, EG&G, USA) converts this light in electric current, which is amplified, digitized and fed to the digital signal processor of a NeXT computer (Red-wood city, USA) for analysis by dedicated software [15]. The three LDF-parameters are obtained at a rate of 21 Hz. The collecting fiber and eye fundus are simultaneously observed by a CCD camera (C2400-02, Hamamatsu Photonics K.K., Japan) mounted in the plane of a relayed image of R2. 4.3. Pupil illumination and observation Preliminary measurements have demonstrated that the variability of the flickerinduced blood flow response in successive recordings could be reduced by maintaining constant the position of the camera relative to the eye pupil. Therefore, the optical system drawn in Fig. 3 (inset) was designed for monitoring the eye pupil position. The iris of the subject is illuminated with PI, a LED emitting at 870 nm. The optical system of observation consists of the ophthalmic lens OL (used at its edge), a mirror M, which redirects the light out of the fundus camera instrument and a lens L1 of 20 mm of focal length, which forms an image of the iris of the subject on a CCD finger camera (Conrad, Germany). To avoid detecting unwanted light such as the probing beam reflexion at the cornea, a high pass filter IR-F (IR820, Schott, Germany) with cutoff at 820 nm was placed in front of the CCD camera. 4.4. Diffuse stimulation Diffuse stimulation is generated by six red and six green LEDs (SL905 OCU-16 and GCU-15, Sloan Precision Optoelectronics, USA) located in an image of the eye pupil (Fig. 3). These LEDs are driven by a generator based on the pulse wave modulation technique [16] and produce user defined waveform stimuli. Generally, luminance stimuli (i.e. green/black flicker) or chromatic stimuli (i.e. constant luminance red/green flicker) are used. These LEDs uniformly illuminate a field of 50 at the fundus with a maximum mean illuminance of 40 lx for each color. 4.5. Contrast reversal pattern stimulation Contrast reversal pattern (checkerboard or grating) were displayed on the monitor of a pattern generating system (VGS series 5, Cambridge Research System, England). These stimuli were reflected by a dichroic mirror DM in Fig. 3 (E30634, Edmund Industrial Optics, USA). This mirror, which reflects the visible light and transmits the near infrared light is located between the ophthalmic lens OL and the eye. After reflection at DM, the stimuli were focused at the retina using a
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spectacle lens SL with a dioptric power specific for the tested eye. Such pattern stimulus sustains a field of approximately 30 with a mean illuminance of 4.5 lx.
5. Experimental part When recording activity-induced blood flow responses, subjects were asked to fixate a pinpoint target located in the plane R1 (Fig. 3). The target was the tip of a 50-mm core diameter optical fiber transmitting light from a green LED. The luminance of the fixation target was chosen to match that of the mean retinal luminance of the stimulus. After locating the optic disk on the monitor screen and bringing the probing beam in the rim region using the bi-prisms BP, the operator aimed the beam at the precisely chosen location by moving the fixation target. Furthermore, during the measurements, the operator monitored the position of the pupil of the tested eye relative to the fundus camera and attempted to keep it constant by readjusting the camera, when necessary. A typical recording protocol consisted of a pre-stimulus period of approximately 30 s followed by a 40–50 s period of stimulation of identical mean luminance and a 30 s recovery period. Prior to the measurements, the tested eye was dilated with two drops of Tropicamide 0.5% (Novartis Ophthalmics AG, Switzerland). 5.1. Blood flow response to diffuse stimulation Fig. 4a shows a typical set of recordings of the blood flow parameters obtained in a normal volunteer in response to a 15-Hz green luminance flicker. The fast
(a)
(b)
Fig. 4. (a) Typical recording obtained from the temporal side of the neuro-retinal rim of the ONH of a normal volunteer. The stimulation was a 50 field diffuse luminance flicker (green/black) at 15 Hz. (b) Blood flow response RF to stimuli of different color ratio r at 3.2 Hz. Error bars represent standard deviations of the mean of 6 recordings.
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variations are due to the changes of the flow parameters during the heart cycle, as confirmed by the recording of the heart pulse using a plethysmograph connected to one of the ear of the subject. In response to the flicker, all three parameters increase to reach a plateau shortly after the start of the stimulation and then decrease back to the baseline following the cessation of the stimulus. For this particular recordings, the percent changes between the last 30 s of stimulation and the pre-stimulation were 19%, 3% and 23% for Vel, Vol and F, respectively. The time course and amplitude of these changes obtained in 15 normal volunteer have been recently described by Riva et al. [17]. Such recordings may be repeated several times varying the characteristics of the stimulus, such as the mean luminance, modulation depth, temporal frequency, and color ratio. The amplitude of the blood flow response (RF) expressed as percent changes from the pre-stimulus F-level was measured in response to 3.2-Hz stimuli of different red/ green luminance ratio (r=red/(red+green)). r=0 corresponds to a green/black luminance flicker, r=1 to a red/black luminance flicker, and r=0.5 to an equiluminant chromatic flicker. Such low temporal frequency chromatic flicker is known to activate predominantly the parvocellular pathway, which is responsibles for retinal color processing [18]. Fig. 4b shows a relatively large RF for a chromatic flicker (r=0.4) compared to luminance flicker (r=0 or 1). 5.2. Blood flow response to pattern stimulation Fig. 5a shows four successive responses to checkerboard pattern stimulation from a subject with excellent target fixation. The pattern was constituted of 36-arcmin black and white checks and was reversed at a rate of 8 Hz. The mean changes in the flow parameters were 9712% (stdev) for Vel, 3276% for Vol and 3876% for F. The dependence of RF on the sharpness of the stimulus image on the retina was measured in one subjects by changing the spectacle lens (SL in the inset of Fig. 3) from 3 to +4.5 diopters with 0 diopter corresponding to a pattern sharply focused on the retina. The pattern consisted in a 225-arcmin black and white vertical grating reversed at a rate of 8 Hz. Fig. 5b shows that RF are obtained for sharply focused pattern stimulus.
6. Discussion We have described a fundus camera-based laser Doppler instrument to investigate the changes in blood flow in the optic nerve of human subjects in response to visual stimulation. The fast response time of this device makes it possible to monitor the time-course of flow response to increased neural activity and fast changes in blood flow, such as those occurring during the heart cycle. To our knowledge, only the LDF-technique has been applied to the monitoring of the optic disk blood flow response to visual stimulation. In the retinal microvasculature, the amplitude of the blood flow response to diffuse luminance flicker has been investigated using the Heidelberg retina flowmeter (HRF).
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(a)
(b) Fig. 5. (a) Continuous recording of Vel, Vol and F obtained in a normal volunteer in response to four stimulation periods (checkerboard, 30 , 8 Hz, 36 arcmin). (b) Effect of pattern defocus on the blood flow responses RF (vertical line grating, 30 , 8 Hz, 225 arcmin).
Michelson et al. [19] and Harrison et al. [20] have found increases in peripapillary blood flow in response to diffuse flicker. Attempts to obtain response to pattern stimuli using the HRF remained, however, unsuccessful [21]. Indirectly, the flow response to diffuse stimuli has been investigated using the blue field entoptic simulation technique [22]. Flicker-induced increases in retinal vessel diameter was also reported [23–25]. Furthermore, the amplitude of the blood velocity changes in the central retinal artery and vein has been measured using ultrasound sonography [19]. In the optic nerve head, as in the brain, increased neuronal activity induces detectable and reproducible increases in local blood flow. This capability opens new avenues in the identification of the factors which regulate blood flow in this neural tissue of the ocular fundus. This instrument should be useful in the study of the role
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of the circulation in the pathogenesis of various ocular diseases of vascular origin, such as glaucoma and diabetic retinopathy.
Acknowledgements . The authors thanks Ce! dric Schopfer for the design and implementation of the diffuse stimulus generator and technical expertise. This work was supported by the Swiss National Science Foundation Grant 32-53785.
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