Principles of optical coherence tomography

Introduction

In modern ophthalmology, there is an increasing demand for high-resolution imaging of the ocular tissues. In particular, the diagnosis of retinal disorders has been dramatically improved by the introduction of modern imaging techniques such as scanning laser ophthalmoscopy, fluorescein angiography, and nerve fiber polarimetry. While these techniques are able to image retinal structures at a high transverse resolution, depth resolution is limited by the numerical aperture of the eye and by aberrations of its optical components to about 300 µm. This prevents the use of these technologies for measuring and imaging the depth of individual fundus layers (an exception is nerve fiber polarimetry, which can determine the thickness of the retinal nerve fiber layer by measuring its birefringence).

Optical coherence tomography (OCT) has overcome this limitation. Based on the special coherence properties of superluminescent diodes (SLDs) and femtosecond pulse lasers, OCT enables the recording of cross-sectional images of transparent and semi-transparent tissues at a high resolution.1,2 The main application field of OCT is retinal imag- ing,3-6 while another ophthalmic application field is imaging of the anterior eye segment.7

OCT imaging is similar to ultrasound B-mode imaging, except that near infrared light is used instead of sound to probe the tissue. OCT images are made up of several optical A scans (in analogy to ultrasound A scans). The basic ranging technique of OCT that provides the optical A scans is partial (or low) coherence interferometry (PCI or LCI), a technique developed after the mid-1980s for measurement of intraocular distances.8-12 PCI enables high-depth resolution that is completely indepen-

dent of the numerical aperture and the aberrations of the eye. Instead, the axial resolution solely depends on the coherence properties of the probing light beam. In standard applications, an SLD with a bandwidth of ~25 nm yields an axial resolution of ~10 µm. This provides OCT images whose resolution is at least one order of magnitude better than that of ultrasound B-mode imaging. By using state- of-the-art femtosecond pulse lasers as the light source, a depth resolution of ~3 µm can be achieved.13 Another advantage of OCT, compared to ultrasound imaging, is that it operates as a non-contact technique and is therefore much more comfortable for the patient and avoids any risk of corneal infection.

This chapter describes the basic principles of OCT, the generation of optical A scans, the synthesis of A scans to an OCT image, and discusses axial and transversal resolution, image contrast, and the advantages and limitations of the technique. Alternative methods and extensions, such as dual beam OCT, transversal (or en-face) OCT, and Doppler OCT, are discussed briefly, and finally, an outlook regarding future developments is presented.

Basics of optical coherence tomography

Partial coherence interferometry

Optical A scans are generated by PCI, the basic ranging technology of OCT. The goal of PCI is, very precisely, to locate the depth position of the back-reflecting or backscattering layers within a sample. Contrary to classical interferometry, which uses laser light of high spatial and temporal coher-

Fig. 1. Basic principle of partial coherence interferometry.

ence, PCI needs a light source that is only partly coherent. It has to have high spatial coherence, but very low temporal coherence (or short coherence length lc).14 Roughly speaking, high space coherence means that all light components across the beam cross-section are in phase with each other, i.e., any parts of the beam that have travelled the same distance from the source can interfere with each other. Short coherence length means that only those beam components that have travelled approximately the same distance from the light source can interfere with each other (components whose travel distances from the light source differ by more than lc cannot interfere).

Figure 1 illustrates the basic principle of PCI. A low coherence light source (usually an SLD) emits a short coherence light beam towards a Michelson interferometer. At the beam splitter, the beam is divided into two components: the reference beam directed towards the movable reference mirror, where it is retroreflected, and the sample beam directed towards the sample. This beam is backreflected or backscattered, usually at several interfaces within the sample whose positions are to be measured. Both beam components are recombined at the beam splitter and superimposed on the photodetector. In order to determine the positions of the sample’s interfaces, the reference mirror is moved, while the light intensity is measured as a function of the position of the mirror. The signal is recorded on a PC and displayed on the computer monitor after the measurement. Any time the path difference from the beam splitter to the reference mirror equals one of the path differences from the beam splitter to one of the sample interfaces, interferometric light modulation occurs: an oscillating signal intensity is recorded. The frequency of the oscillations equals the Doppler shift frequency of the reference light caused by the moving reference mirror, and acts as the carrier frequency of the coherence signal. At the position of maximum os-

cillation strength, the reference mirror position corresponds to the position of the sample interface (see Fig. 1). It should be mentioned that distances measured by this method are optical ones, i.e., geometric distances multiplied by the (group) refractive index of the sample.9 In order to obtain the true geometric distances, the optical distances have to be divided by that index.

Usually, only the envelope of the interferometric signal is recorded. This saves computer memory and improves the readability of the signal. The position of a sample interface is now indicated by a signal peak in the A scan. The width of the peak is equal to the ‘round trip’ coherence length (lc/2, because the beam travels back and forth in the interferometer), and determines the depth resolution of the technique.

Figure 2 illustrates the application of PCI for measuring intraocular distances. The bulk Michelson interferometer is now replaced by a fiber optic Michelson interferometer, as is used in the majority of present OCT set ups.1,4-7 The advantages of the fiber optic set up are its compact, light-weight design, its robustness, and ease of alignment. The beam splitter is replaced by a fiber optic 50:50 X-coupler, and the reference mirror by a retroreflector. An SLD and a photodetector are directly coupled to the interferometer via single mode fibers; therefore, these components need no further alignment. The eye is placed in the sample path and, as indicated in Figure 2, a signal peak is recorded every time the retroreflector position coincides with an intraocular interface. In this way, intraocular distances can be determined with micrometer precision.9-12,15

Synthesis of optical coherence tomograms

In order to obtain cross-sectional OCT images, several optical A scans are recorded at adjacent sample

Fig. 2. Partial coherence interferometry for intraocular ranging. Principle of optical A-scan recording.

positions. For this purpose, a pair of scanning mirrors deflects the probing beam to different sample positions and an A scan is recorded at each position (see Fig. 3). The signal intensities are converted to gray levels or color values, and the individual scans are mounted to form a two-dimensional gray level or false color image showing a cross-section of the sample, the OCT tomogram.

In conventional OCT of the retina, realized by the first commercially available system, the ZeissHumphrey OCT scanner, 100 A scans are typically recorded within ~1 second. The transversal orientation of the tomographic section can be arbitrarily chosen by tilting the scanning mirrors between the individual A scans in a suitable, predetermined way. Thereby, horizontal, vertical, arbitrarily oriented, and even circular sections (e.g., around the optic nerve head) can be recorded.

Image contrast

Conventional OCT is based on the intensity backreflected or backscattered at object structures. Backreflection occurs at interfaces within the sample where the refractive index changes abruptly. However, such an interface is only observable in OCT images if the light is retroreflected, i.e., if the interface is oriented (nearly) perpendicular to the probing light beam. In this case, the interface shows up as a bright line. Backscattering is caused by particles having a refractive index different from that of the surrounding matrix. Backscattering occurs in all directions, however, the intensity will vary with backscattering angle, depending on particle size, shape, and refractive index mismatch to

Fig. 3. Synthesis of an OCT tomogram of the retina by recording and mounting several A scans.

the matrix.16 Therefore, backscattering structures are visible in OCT images independently of the orientation of the backscattering structure. Backscattering surfaces show up as bright boundaries, volume backscatterers (particles dispersed in a surrounding matrix) are imaged as ‘speckled’ structures; due to attenuation by scattering, the backscattered intensity decreases exponentially with depth. Absorption is another effect that attenuates light remitted by the sample. Strongly-absorbing material can shadow tissue structures beneath, and render them invisible.

It should be mentioned that OCT is a ‘high pass’ technique.2 This means that only areas where the refractive index changes over a very short distance (of the order of a wavelength or less) give rise to detectable backscattered (or back–reflected) intensity. Areas with only slowly spatially varying refractive indexes remain invisible in standard OCT (enhanced OCT contrasting techniques, e.g., polar- ization-sensitive OCT or phase-contrast OCT, which are currently under development, can improve this situation in certain cases).

As already mentioned, scattering (and also absorption) attenuates the probing light exponentially with depth. Furthermore, the back-reflection or backscattering coefficient can vary by a large amount, depending on the optical properties of the tissue structures. This causes a large variation in backscattered intensity, usually too large to be

displayed on a linear color or gray level scale. Therefore, a logarithmic intensity scale is usually adopted in OCT.

Figure 4 shows an OCT tomogram recorded horizontally through the optic nerve head of a human eye in vivo. It illustrates the image contrast typically obtained from retinal structures: a strong backscattered intensity is obtained from the retinal nerve fiber layer and from the posterior surface of the retina, probably caused by the retinal pigment epithelium (RPE) and the choriocapillaris. Intermediate layers show weaker backscattering with a typical ‘speckled’ appearance. Due to light attenuation, only very weak light intensity can be observed from structures below the intense RPE/ choriocapillaris band. Structures with a rather constant refractive index, such as the vitreous or edema (not shown here), do not backscatter light and are displayed in black. Blood absorbs the wavelength used for retinal OCT imaging (~800–850 nm). Therefore, blood vessels can shadow structures beneath them.

System performance

Light source

The most important component of an OCT system with respect to its performance is the light source. Via absorption and scattering coefficients, the central wavelength determines how deep the light penetrates into the tissue. The bandwidth defines the depth resolution, and the power determines the sensitivity of the system.

OCT light sources have to fulfill very special criteria: good spatial coherence is required, while the temporal coherence has to be very low (short coherence length). Three types of light sources that fulfill these criteria are presently in use for OCT

applications: femtosecond pulse lasers, broadband fiber optic sources, and superluminescent diodes. Femtosecond pulse lasers achieve the shortest coherence lengths and therefore the best axial resolution.13,17 However, they are very expensive and difficult to operate. Therefore, their use for OCT applications is presently restricted to a very few research laboratories.

Broadband fiber optic sources are derived from the fiber optic amplifiers used in the telecommunications industry. Because of their origin, they are presently restricted to the wavelengths most commonly used in telecommunications: wavelength bands centered around 1300 and 1550 nm. The advantage of these sources is that they provide rather a high output power (up to ~30 mW) from a single mode fiber, which ensures excellent spatial coherence. This type of source is preferably used for OCT of scattering tissue, such as skin, teeth, mucosa, etc., because the longer wavelengths of 1.3–1.55 µm are less scattered and allow greater penetration depth (~1–2 mm) into these types of tissue. However, these wavelengths cannot be used for retinal OCT because they are absorbed by the water contained in the ocular media. An ophthalmic application for 1300 nm light is OCT of the anterior eye segment,7 where the better transmission of this wavelength is used to image the sclera and iris.

SLDs are most commonly used in OCT. They are comparatively cheap, lightweight, easy to operate, and available in a comparatively large selection of wavelengths. Available center wavelengths presently range from 670-1550 nm. The wavelength range used for retinal OCT is ~800–850 nm. In this range, output powers from ~0.5-20 mW are available from single mode fibers. The ocular media are usually perfectly transparent in this wavelength range, and the retina tolerates more light intensity than at the shorter wavelength of 670 nm.

With a probing beam power of 750 µW at λ0 = 0.83 µm, the maximum permissible exposure time of a point at the retina is eight seconds18. Within this time, several OCT images can be recorded. Furthermore, since the beam scans the retina (i.e., the light energy is distributed over several points of the retina), OCT imaging is well below laser safety limits.

Resolution in OCT

One of the main advantages of OCT is the complete decoupling of axial (depth) resolution from transverse resolution. Contrary to scanning laser ophthalmoscopy, axial resolution is influenced by neither the numerical aperture of the eye nor aberrations of its optical components. The axial resolution of OCT is determined by the coherence length lc of the probing light. For a light source with a Gaussian emission spectrum with emission bandwidth ∆λ (full width at half maximum, FWHM) and center wavelength λ0, the axial resolution in air, for a dispersion balanced interferometer, is given by the following equation:19

A single reflecting interface in the sample arm gives rise to a coherence signal with an FWHM width δz. δz equals the so-called round trip coherence length which is equal to lc/2. If we assume a center wavelength of λ0 h 830 nm for retinal OCT (compare Light source above), the main parameter determining the axial resolution is ∆λ. The larger the bandwidth, the better the resolution. For a typical SLD bandwidth of 25 nm, we obtain δz h 12 µm in air. For OCT imaging within a sample with a refractive index of n > 1, the resolution is improved:

(Strictly speaking, we would have to use the group refractive index ng instead of the usual phase refractive index n9. However, for the ocular media we are interested in, the difference between n and ng is only in the order of ~1%.) If we assume an index of n h 1.3520 for ocular media, we obtain

δzMedium h 9 µm.

It should be mentioned that the resolution figures provided by equations (1) and (2) are only valid if the interferometer is dispersion-compensated. In other words, the amount of (group) dispersion introduced by dispersive media (lenses, prisms, glass fibers, etc.) in the sample arm has to be compensated by an equal amount of dispersion by equivalent media in the reference arm. Otherwise the resolution is degraded. This is not usually a problem for the technical components in the interfero-

meter arms. However, in case of retinal OCT, the dispersion caused by the ocular media also has to be compensated.21 The amount of this dispersion differs between individual eyes, mainly due to different axial eye lengths. In order to obtain OCT images with standard resolution, compensation assuming a mean axial eye length of ~24 mm is sufficient. However, for high resolution OCT employing broadband femtosecond laser sources, careful compensation of the individual eye’s dispersion must be used.13

The transversal resolution in OCT is limited by criteria similar to those in confocal microscopy.22 The smallest distance δx that can be resolved in a transversal direction can be defined as the FWHM diameter of the beam scanning the sample. Normally, a Gaussian beam in the fundamental transversal mode is used in OCT. High transverse resolution demands focusing of the beam to a small beam waist radius w0 (the radius at which the beam intensity falls to 1/e2 of its central value23). The FWHM diameter of this beam, and hence the transversal resolution, can be calculated by:

where f is the focal length of the lens (in the case of retinal imaging, the focal length of the optical components of the eye), and d the FWHM beam diameter of the collimated beam. With a beam diameter of about 1–2 mm, a transverse resolution of the order of ~10 µm can be expected. This is equal to the transverse resolution obtained with scanning laser ophthalmoscopes. Beam diameters larger than ~3–4 mm would not further improve the transverse resolution, because of the aberrations caused by the ocular optical components. Furthermore, a larger beam diameter would reduce the depth of focus considerably, requiring dynamic focusing,24 which would make the system more complex and expensive.

It should be mentioned that the transverse resolution defined by equation (3) requires that the transverse sampling rate is high enough, i.e., optical A scans are recorded at transverse sampling intervals ≤ δx. However, this is usually not the case in retinal OCT. A typical OCT image of the retina consists of 100 A scans equally spaced over a transverse scanning width of ~3 mm. In this case, the transverse resolution is equal to the transverse sampling distance, i.e., 30 µm in our example.

System sensitivity

An important issue in OCT imaging is detection sensitivity. Sensitivity can be defined as the smallest signal just discernible from noise. OCT signals are not generated by the tissue, but are caused by backscattered light. The amount of backscattered

light depends on the incident light intensity and the sample reflectivity Rs. Since we are not interested in the absolute amount of backscattered light but in the tissue parameter Rs, a suitable definition for the sensitivity S in the context of OCT is the ratio between the reflectivity of a perfectly reflecting mirror (R = 1) and the smallest sample reflection coefficient, Rs,min, yielding a signal strength equal to noise:

Rs,min

The sensitivity is usually specified in dB:

ru 1 iu

S(dB) = 10logu u. (5)

qRs,mint

As a rule of thumb, a sensitivity of at least 50 dB is necessary to observe a signal from the retina. However, with S = 50 dB, the strongest backscatterer at the retina, the RPE/choriocapillaris complex, is just discernible from noise, and no further details can be observed. For good quality OCT images of the retina, a sensitivity of about 90–100 dB is required.

The sensitivity of an OCT system is determined by noise. Several noise components contribute to the overall noise. The main noise components are shot noise (quantum noise), electronic amplifier noise, and excess noise caused by the light source.25,26 If SLDs with low to medium output power are used, as in retinal OCT, and if the electronic components are properly chosen, shot noise is the factor limiting the sensitivity (shot noise limitation is the optimum case; this type of noise is caused by the inherent quantum nature of light, and cannot be eliminated). The main factors determining shot noise are light power and detection bandwidth. In the shot-noise-dominated regime, sensitivity increases linearly in proportion to the power incident on the sample. Therefore, light power should be as high as the safety limit permits. Sensitivity is inversely proportional to detection bandwidth. For high sensitivity, a low bandwidth is preferable, however, this implies a slow scanning speed and long recording times. Therefore, there is a trade-off between recording time and sensitivity. The recording time of 100 A scans within one second, as presently employed in the ZeissHumphrey OCT scanner, seems to be a good compromise.

Motion artifacts

Ocular motions such as microsaccades or axial eye motions during the recording can cause image artifacts and distortions. Microsaccades are small, rapid rotational eye movements with an amplitude

of 1-20 minutes of arc. If a large microsaccade occurs during the recording period, the tomographic scan should be repeated. Axial eye motions are more difficult to avoid. Even very small axial movements of about a few tens of microns can severely distort the image. Such movements can easily occur, especially in elderly patients. In order to correct for distortions caused by axial eye motions, digital image processing techniques can be used.

Figure 5 presents an example.27 An OCT image was recorded across the fovea of a human retina. Figure 5a shows the raw image. Axial motions are clearly visible: there is an overall backward drift of the eye, overlaid with a small tremor. Figure 5b shows the image after digital post-processing. As can be seen, motion artifacts are considerably reduced by this technique. However, it should be mentioned that the result of the image processing technique depends on the assumption of what the image should look like. One of the common image processing techniques assumes that the RPE should resemble a smooth, straight line. Each deviation is assumed to be caused by axial eye motions. Consequently, the image is corrected to display a flat RPE. While the assumption is meaningful in most cases, there are retinal disorders in which the RPE is not flat. The image processing algorithm would also flatten these images and the disorder might be missed. Therefore, care has to be taken if postprocessed images are used.

Another technique that avoids this type of motion artifact problem is the so-called dual-beam OCT technique. This will briefly be discussed in the following section.

Alternative and extended OCT techniques

Dual-beam OCT

The dual-beam OCT technique completely eliminates artifacts caused by axial eye motions. It is based on dual-beam PCI, which was initially introduced to measure axial eye length and other intraocular distances with micrometer precision.8,9,15 A wide range of applications of dual-beam PCI in physiological studies and in therapeutic applications, especially in the context of cataract surgery, have been reported.12,28,29

The main idea of dual-beam PCI and OCT is to use the cornea as a reference surface. Both the reference and the sample beam are directed towards the eye (they form a coaxial ‘dual beam’, see Fig. 6a). Interferometric modulation is recorded at the photodetector, if the path difference cornea–retina is matched by a corresponding path difference on an external Michelson interferometer. Since no absolute positions are measured, but just path differences, any influence of axial eye motions is completely eliminated. In other words: the sample

(retina) moves synchronously with the reference (cornea).

Figure 6b is an example of a horizontal section recorded across the fovea of a human eye in vivo.30 Because of the slow scanning speed used with the experimental set up, the transverse resolution is limited by microsaccades to ~100 µm. No digital image post-processing was necessary for producing the tomogram seen in this Figure. The disadvantage of the technique is the increased system complexity and the rather complicated alignment procedure needed to receive the beams reflected at the cornea and retina simultaneously.

En-face OCT

Conventional OCT images are made up of A scans, i.e., the information on the reflectivity distribution

within the sample is obtained by subsequent axial scans recorded at different transversal sample positions. In order to obtain information on a whole 3D sample volume, A scans have to be recorded along a 2D grid distributed over the sample. Other scanning patterns can also be used to obtain 3D information. The information can be obtained by transversally ‘slicing’ the sample volume:31 a complete transversal image corresponding to an object depth defined by the reference arm length is recorded by raster scanning the sample beam across the object. Thereby, the distribution of backscattered light within a slice of thickness equal to the coherence length is recorded. After recording one transversal slice, the reference arm length is altered and the next transversal image slice is obtained. This process is repeated until a whole 3D data set has been recorded. While the A-scan tech-

nique provides a carrier frequency based on the Doppler shift of the reference light induced by the scanning reference mirror, transversal, or en-face, OCT usually needs additional optical components, such as phase modulators or frequency shifters, for generation of the carrier frequency.

Podoleanu et al. have adapted the transversal OCT technique for retinal imaging.32 They used the path length modulation induced by transversally raster scanning the retina with the sample

beam by a galvo mirror pair for generating the carrier frequency. The disadvantage of this method is that, due to the nonlinear path length dependence on the scanning angle, a frequency spread of the photoelectric signal occurs, leading to a reduced signal-to-noise ratio. The advantage is that no additional phase modulators are needed. The images thus recorded look rather fragmented because, due to the short lc, the coherence condition within one transverse image is only met for small

areas of the retina. Therefore, several subsequent transverse OCT images have to be recorded and mounted to a 3D data set with the help of a simultaneously recorded SLO image33 (an SLO detection system is integrated with the instrument). From the resulting 3D data set, sections of arbitrary orientation through the retina can be derived by software.34 This can be a considerable advantage because section geometries do not have to be defined in advance; instead, they can be chosen after data acquisition, based on the recorded information. Reported recording times are about 0.5 seconds for one transversal image; a stack of 112 en-face images can be recorded within 56 seconds.

Figure 7 shows an image derived from a 3D data set of the optic nerve head of a human eye recorded in vivo.34 Two mutually perpendicular sections across the nerve head, oriented perpendicular to the retinal surface (along the x-z and y-z planes; z: axial direction), are shown. Imaged structures comprise the RPE, retinal nerve fiber layer, and lamina cribrosa.

Doppler OCT

effect is used in laser Doppler velocimetry to measure flow velocity in a wide range of applications. The short coherence light sources used in OCT enable this effect to be used for depth resolved flow velocity measurements and flow imaging.

Different schemes for Doppler OCT have been reported.35-37 The method first reported by Izatt et al.37 is based on the conventional A-scan OCT technique described above under Basics of optical coherence tomography, and is therefore best suited for a brief explanation regarding the basic principles of Doppler OCT in the context of this book.

As mentioned above (Partial coherence interferometry), an oscillating coherence signal is generated every time the reference mirror position coincides with the position of a backscatterer in the sample arm. The frequency f0 of this signal equals the Doppler frequency shift of the reference beam caused by the reference mirror moving with speed vr:

If light is backscattered by a moving particle, its frequency is shifted by the Doppler frequency. This

If the scatterer in the sample arm moves with a speed vs, the frequency of the sample beam is also

Doppler shifted, by a frequency:

where α is the angle between the incident sample beam and the direction of the scatterer movement, and n is the refractive index of the medium in which the scatterer moves. The resulting frequency detected by the photodetector is:

the sign in equation (8) depends on the movement direction of the scattering particle.

To measure and image flow velocity by OCT requires the recording of the full interferometric signal. Just to record the signal envelope is not sufficient. After recording the A scans making up an OCT image, each A scan is analyzed by a shorttime fast Fourier transform, which provides the local distribution of frequencies within the A scan. Deviations from the center frequency f0 are caused by moving scatterers. These deviations are extracted and converted into color values, indicating the local particle speed. The data sets are then mounted in the usual way in order to provide a map of velocity distribution within the sample. The signal intensities contained in the A scans can be separately displayed as conventional OCT images of the sample structure.

Figure 8 shows a Doppler OCT image recorded in this way.38 A cross-section recorded in vivo in a human retina, at a location superior to the optic nerve head, is shown. A structural OCT image

displaying reflected intensities (gray scale image) is superimposed with flow information reconstructed from Doppler shifts. The sample contains two vessels, branches of the central retinal artery and vein superior, whose cross-sections can be seen in the tomogram. Blood cells of opposite flow directions cause positive and negative frequency shifts in the light backscattered from within the two vessels, indicated in red and blue. The saturation of the color is a measure of the blood cell speed.

It should be mentioned that there is a trade-off between spatial resolution and velocity resolution. A good velocity resolution requires a larger Fourier window, i.e., a larger segment of an A scan is needed to obtain the necessary amount of data. However, a longer A-scan segment implies reduced spatial resolution in an axial direction. More sophisticated Doppler OCT schemes that overcome this problem have recently been reported.39,40

Conclusions and Outlook

Several improvements and extensions of OCT are presently under development. The aim of these developments is essentially improved resolution, speed, and image contrast.

Improved axial resolution requires light sources of a higher bandwidth. Recently, the first retinal OCT images recorded with a state-of-the-art Ti: Al2O3 femtosecond pulse laser were presented.13 Corresponding images can be seen in Figures 4 and 5 of the chapter by J.S. Schuman et al. (see p. 147). Using a bandwidth of 155 nm, a depth re-solution of ~3 µm was obtained. Comparison with a standard resolution image clearly demonstrates the dramatic

improvement in resolution. Implementation of this technique into clinically-useful OCT systems depends on the availability of a cheap and robust light source.

With presently available standard OCT systems, one OCT image consisting of 100 A scans is recorded in one second. Acquiring 3D information is presently rather time consuming. Different approaches for solving this problem are under investigation. Rapid scanning optical delay lines have been reported that achieve more than 1000 A scans per second.41,42 Parallel OCT schemes recording up to 58 × 58 A scans simultaneously with speciallydesigned detector arrays have been suggested43. Finally, a variant of en-face OCT that can record a 3D data set consisting of 32 transverse OCT images within one second has been announced.44 All these techniques depend on light sources providing sufficient output power to achieve high sensitivity.

Image contrast in conventional OCT is based on the intensity of backscattered light. Several structures yield poor or no contrast on an intensity basis. Other properties of light, as wavelength and polarization, can be exploited to improve contrast or to generate other types of contrast. OCT at wavelengths with different absorption coefficients in oxygenated and de-oxygenated blood has been suggested for measuring and imaging oxygen saturation in vessels.45 Polarization-sensitive OCT has been reported as a means of measuring the thickness of the birefringent nerve fiber layer for glaucoma diagnostics.46

These developments seem promising technologies for future, enhanced OCT applications. However, much additional research and developmental work has to be done before they can be converted into robust, clinically-applicable tools.

Acknowledgments

The author acknowledges the permission of several authors and editors to reproduce figures from their works. Furthermore, the author wishes to thank Professor A.F. Fercher, head of the Institute of Medical Physics, University of Vienna, and several other co-workers, for fruitful discussions. Part of the work reported in this chapter was financed by the Austrian Science Foundation.

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Abstract

Retinal photocoagulation has been used for diabetic and other ischemic retinopathies for decades. While most ophthalmologists are convinced of its clinical effect, many do not understand how if works, i.e., the mechanism of effect is unclear. The purpose of this chapter is to present a theory on how the physical energy of visible laser light, such as that of the argon laser, is transferred into a biological effect and a therapeutic effect on retinopathy.

Introduction

The physiological mechanism of photocoagulation can be seen in following steps:

•The physical light energy is absorbed in the melanin of the retinal pigment epithelium. The adjacent photoreceptors are destroyed and are replaced by a glial scar, and the oxygen consumption of the outer retina is reduced.

•Oxygen that normally diffuses from the choriocapillaris into the retina can now diffuse through the laser scars in the photoreceptor layer without being consumed in the mitochondria of the photoreceptors.

•This oxygen flux reaches the inner retina in order to relieve inner retinal hypoxia and raise oxygen tension. As a result, the retinal arterioles constrict and the blood flow decreases.

•Vasoconstriction increases arteriolar resistance, decreases hydrostatic pressure in the capillaries and venules, and reduces edema formation according to Starling’s law.

•Hypoxia relief reduces the production of growth factors such as vascular endothelial growth factor (VEGF), and neovascularization is reduced or halted.

•Retinal laser photocoagulation improves inner retinal oxygenation, which affects retinopathy through the relief of hypoxia and consequent changes in growth factor production and hemodynamics.

Light coagulation and laser treatment of the retina were introduced to ophthalmology around the middle of the last century. They are widely used for the treatment of diabetic retinopathy and other ischemic retinopathies. In this chapter we will focus on argon laser treatment and diabetic retinopathy, even though the theory applies to all visible light photocoagulation for ischemic retinopathies.

The usefulness of retinal photocoagulation was established in clinical trials, such as the Diabetic Retinopathy Study1,2 and the Early Treatment Diabetic Retinopathy Study.3-8 The mechanism of treatment was not investigated in these trials and eludes many ophthalmologists to this day. Ophthalmic physiologists have strived to understand the physiologic mechanism using histological, physiological, and clinical research techniques. This paper proposes a general theory on the mechanism of action of retinal laser treatment, based on these studies.

Histology

Visible light that is sent into the eye is predominantly absorbed in the melanin of the retinal pigment epithelium, and the heat results in thermal destruction of retinal pigment epithelium cells, the adjacent photoreceptors and, to some degree, the choriocapil- laris.9-14 Figure 1 shows a pigmented rabbit retina with a two-week-old argon laser burn. The outer retina and retinal pigment epithelium are photoco-

Fig. 2. Schematic drawing indicating an oxygen flux coming from the choroid and passing through a laser scar into the ischemic inner retina. In the laser scar, photoreceptors are replaced by glia, and the oxygen consumption is decreased. The oxygen flux from the choroid would normally be consumed by the photoreceptors, but since the oxygen consumption of the glia is less, the oxygen flux reaches the inner retina, where ischemic hypoxia may be relieved. (Redrawn with permission from Stefánsson E, Graefe’s Arch Clin Exp Ophthalmol 228:120-123, 1990.)

agulated and the photoreceptors are absent, whereas the inner retina is relatively intact.

The histology predicts the physiological effect. In the normal situation, oxygen and nutrients diffuse from the choriocapillaris into the retina and these are consumed by the photoreceptors, which have a very high density of mitochondria and a high oxygen consumption. In laser scars, the photoreceptors

are replaced by glia, which have few mitochondria and a low oxygen consumption. This means that the laser scars function as windows where oxygen consumption is low and it can diffuse from the choroid through the photoreceptor layer into the inner retina (Fig. 2). This should lead to increased oxygen tension in the inner retina.

Oxygen physiology

Oxygen tension in the inner retina following laser treatment has been studied in a number of experimental animals and diabetic patients. The improved oxygenation following retinal laser treatment was first shown in rhesus monkeys, when Stefánsson et al.15 showed that retinal oxygen tension was much higher in laser-treated areas than in untreated areas of the same retina. These findings have been confirmed by a number of researchers,16-18 including Pournaras et al., who showed that laser treatment reverses retinal hypoxia induced by branch retinal vein occlusion.19 Diddie and Ernest found an initial increase in retinal oxygen tension following retinal laser treatment, but this effect was short lived.20

Figure 3 shows data from rabbits, part of whose retina was photocoagulated and the other part not. The oxygen tension of the retina was much higher over the photocoagulated area compared to the untreated area. This effect was seen even as early as one day following treatment.21 Oxygen tension measurements from the preretinal vitreous of human diabetics undergoing vitrectomy have shown the same result, where the oxygen tension over lasertreated areas of the retina was significantly higher than the oxygen tension over untreated areas of the same retina.22

Fig. 3. Graph showing preretinal oxygen tension (torr = mmHg) in rabbits where part of the rabbit retina was photocoagulated and another part was not. The oxygen tension in the retina was significantly higher in the photocoagulated area compared to the untreated area. This effect was seen even as early as one day following treatment. (Reproduced from Novak et al.21 by courtesy of the publisher.)

Vascular physiology

The retinal circulation autoregulates its own diameter and blood flow. It dilates in hypoxia and constricts if the oxygenation is increased. The improved oxygenation of the inner retina should lead to constriction of the retinal arterioles and venules, and to decreased blood flow following laser treatment. This was originally proposed by Wolbarsht et al.,23,24 and has been tested and demonstrated in numerous other studies.25-32 Hessemer and Schmidt suggested that total ocular blood flow is also decreased after scatter laser treatment of the retina.33

Wilson et al. used photographs and data from the Diabetic Retinopathy Study (DRS) 2 to evaluate the diameters of retinal vessels in patients treated with panretinal photocoagulation for proliferative diabetic retinopathy.34 On DRS photographs, arteriolar and venular constriction was seen following both laser and xenon arc photocoagulation. Furthermore, there was a statistically significant correlation between the vasoconstriction and the disappearance of disc neovascularization, according to the DRS graders (Fig. 4). Remky et al. also observed constriction of the retinal venules following laser treatment.35

Retinal vasoconstriction is also seen following macular grid photocoagulation.36 The temporal arterioles and venules and their macular branches all constrict significantly following photocoagulation for diabetic macular edema, and the same is also true of laser treatment in branch retinal vein occlusion.37

In diabetic macular edema, Kristinsson et al. documented progressive vasodilatation prior to the onset of edema, and constriction following laser treatment as the edema disappears (Fig. 5).38

Fig. 4. Graph showing the correlation between decreasing arteriolar diameter following panretinal photocoagulation in the Diabetic Retinopathy Study and the regression of disc neovascularization. There was a statistically significant correlation between the vasoconstriction and the disappearance of disc neovascularization, according to the DRS graders. (Reproduced with permission from Wilson et al., Am J Ophthalmol 106:657-664, 1988.)

Fig. 5. Graph showing the diameter of a temporal retinal arteriole in a diabetic patient. In 1989, the female patient was diagnosed with diabetic macular oedema (DMO) and received macular grid laser treatment. The graph demonstrates progressive vasodilatation prior to the onset of edema and constriction following laser treatment as the edema disappears. (Reproduced from Kristinsson et al.36,38 by courtesy of the publisher.)

Proliferative retinopathy

Improved oxygenation of the inner retina and the relief of hypoxia lower the production of VEGF in the retina, and hinder neovascularization. Aiello et al.39 and Augustin et al.40 showed that VEGF levels in the vitreous of diabetic patients are elevated in patients with proliferative retinopathy, and are significantly lower following retinal photocoagulation. Lip et al.41 showed that VEGF levels were elevated in the blood of diabetics with proliferative

retinopathy, and fell following panretinal photocoagulation. Previously, Boulton et al.42 and Smith et al.43 had shown that VEGF and VEGF receptors in the diabetic retina are present in proliferative retinopathy, and are reduced in eyes that have had laser treatment. Pournaras et al.44 demonstrated that hyperoxia also reduces VEGF production in ischemic retina, suggesting that laser treatment and hyperoxia have the same effect on VEGF production. Miller et al.45 showed, in an experimental animal, that VEGF levels are elevated when retinal ischemia is induced by means of a branch vein occlusion, demonstrating the correlation between ischemia (hypoxia) and VEGF levels.

In addition to VEGF reduction following retinal photocoagulation, reduced vasodilatation and endothelial stretching may also reduce the effect of growth factors on vascular endothelium. This was first suggested in 1983,46 based on clinical observations, and recent laboratory data have pointed to the role of vascular tissue stretching in vasoproliferation.47-50 Suzuma et al. found that capillary stretching increases thymidine uptake and VEGF production in capillaries.51 Retinal hypoxia stimulates neovascularization in two ways, firstly by hypoxia-induced VEGF production, and secondly through the autoregulatory

dilatation of capillaries, which stimulates growth directly. Reduction in retinal hypoxia through laser treatment will reduce VEGF formation and capillary stretching, and reduce neovascularization through both mechanisms (Fig. 6).

Macular edema

By definition, edema is an abnormal accumulation of water in a tissue, and the development and regression of edema is based on the movement of water between vascular and tissue compartments. Starling’s law describes the steady-state water exchange between the vascular compartment and the extracellular tissue compartment. Hydrostatic pressure in the vessel drives water into the tissue, and this is opposed by oncotic (osmotic) pressure differences between blood and tissue. In the normal state, these forces are in balance and there is no net movement of water between tissue and vascular compartments. However, if the hydrostatic pressure in the capillaries and venules is increased, this drives water into the tissue and creates edema, whereas decreased hydrostatic blood pressure would decrease edema, assuming that the oncotic pressures are constant.

The retinal arterioles serve as resistance vessels and control the hydrostatic pressure downstream. Dilated arterioles have less resistance, and consequently the blood flow and hydrostatic pressure are increased downstream in the capillaries and venules, where high hydrostatic pressure dilates these thinwalled vessels, according to LaPlace’s law. The diameter of the arterioles and venules in the retina is an indicator of the hydrostatic blood pressure in the retinal microcirculation.

We have observed that, prior to the development of diabetic macular edema, the retinal arterioles and venules gradually dilate.38 Bek also observed the role of vasodilatation in diabetic macular edema.52 Following macular laser treatment, the arterioles and venules constrict as the retinal edema regresses (Fig. 5).36 The same pattern is seen in branch retinal vein occlusion.37

The regression of retinal edema following macular laser treatment is easily understood in light of Starling’s law. The laser treatment reduces oxygen consumption in the outer retina, and extends the oxygen flux from the choroid into the inner retina. This causes the arterioles to constrict (autoregulation), their resistance is increased to the fourth power of the radius, and the hydrostatic pressure downstream is decreased. Starling’s law predicts that this will reduce the water flux from the vessel into the tissue, and the oncotic pressure will now manage to drive the water back into the vessels and reduce the edema. The decreased hydrostatic pressure will also reduce the dilatation of the venules, as has been observed (Fig. 6).

Since vascular hydrostatic pressure plays a central role in Starling’s law, this explains why arterial hypertension is so important in diabetic edema, and why reducing the blood pressure may be helpful against diabetic macular edema.53

Conclusions

This chapter describes a general theory on the mechanism of action of retinal photocoagulation in diabetic retinopathy and related retinopathies (Fig. 6). The theory focuses on the effect of the treatment on retinal oxygenation, and the consequence of the oxygenation effect. The same mechanism can be used to understand the effect of laser treatment on different aspects of diabetic retinopathy, as well as the effect of vitrectomy on diabetic retinopathy.54

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diabetic retinopathy. Am J Ophthalmol 113:36-38, 1992

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24.Wolbarsht ML, Landers MB III, Stefánsson E: Vasodilation and the etiology of diabetic retinopathy: a new model. Ophthalmic Surg 12(2):2104-2107, 1981

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26.Richard G, Kreissig I: Effect of laser therapy in diabetic retinopathy on the hemodynamics of retinal vessels. Klin Mbl Augenheilk 186(2):107-109, 1985

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31.Mendivil A: Ocular blood flow velocities in patients with proliferative diabetic retinopathy after panretinal photocoagulation. Surv Ophthalmol Suppl 41:S89-95, 1997

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Y:Alterations of retinal microcirculation in response to scatter photocoagulation. Invest Ophthalmol Vis Sci 39(5): 769-776, 1998

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B:Retinal vessels before and after photocoagulation in diabetic retinopathy. Determining the diameter using digitized color fundus slides. Klin Mbl Augenheilk 209(2/3):7983, 1996

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Introduction

Conventional laser techniques for corneal surgery are based on the application of high energy ultraviolet (UV) nanosecond (ns) laser pulses for the precise photoablation of stromal tissue.1 In general, an argon fluorite (ArF) excimer laser at 193 nm, at a low pulse repetition rate of some hundred Hz, is used as laser source. The energy of a single photon of 6.4 eV is sufficient to break molecular bonds (photodecomposition). For example, a dissociation energy of 3.6 eV is required for the C-C bond as well as the C-O bond, and 4.8 eV for the O-H bond. An energy of 6.4 eV is necessary to break C=C bonds. As a result of 193-nm laser exposure, fragments of low molecular weight are ejected. This non-thermal process is called ablative photodecomposition.

Due to the high absorption coefficients, the light penetration depth for 193-nm radiation is limited to < 5 µm.2 The radiation is mainly absorbed by cytoplasm of the outermost cell layer. Therefore, ablation can only occur at the surface. Non-invasive intratissue ablation is impossible with excimer lasers. Ablation of in-depth tissue, such as the stroma underneath the epithelial layer, first requires the removal of the superficial layer. This can be done either (i) optically with the same laser source, (ii) mechanically with a microtome, or (iii) by mechanical manipulation after chemical disconnection of the epithelial layer from the stroma.

Photorefractive keratectomy (PRK)3 removes the epithelial layer optically by subsequent layer-by-layer UV photoablation, starting from the outermost layer. After the epithelium has been removed, photoablation of the stroma can be performed. PRK has the disadvantage of a long healing period and relatively severe postoperative pain.

The laser-assisted in situ keratomileusis (LASIK) method implies UV laser ablation of corneal stromal structures after partial removal of the epithelium by means of a microtome.4,5 A corneal flap is produced. After the laser procedure, the flap is relocated, and it covers the stroma again. Although the epithelium and Bowman’s layer can be preserved, LASIK faces problems due to microtome-related complications and the weak binding forces between the flap and the stroma. The healing process is much faster than with PRK.

When LASEK (laser epithelial keratomileusis) is being used,6 excimer laser treatment follows the removal of the epithelium, using a combined knife technique and 20% alcohol. The postoperative pain level is between those of PRK and LASIK.

Although millions of excimer laser treatments on human cornea have been performed within the last couple of years, the application of high-energy UV photons and their possible long-term harmful effects are still under discussion. A haze of unknown origin is frequently observed within a few years of corneal surgery with the ArF excimer laser, after PRK and LASEK for the treatment of higher corrections.2

A new direction in laser corneal surgery implies the use of ultrashort wavelength lasers in the visible (VIS) and near infrared (NIR) spectral range. Due to the high light penetration depth in this wavelength region, non-invasive intratissue surgery becomes possible. Laser beams which are focussed into the tissue are required in order to confine the surgical effect to the intratissue region of interest (Fig. 1).

Clinical studies have been performed with nanosecond, picosecond (ps), and femtosecond (fs) laser pulses. When using the pulsed Nd:YLF laser at 1053 nm with a 30-60 ps pulse width, better results

Fig. 1. Principle of intrastromal surgery using VIS/NIR laser beams focused to an intratissue region of interest.

were obtained than when ns laser pulses were used. However, no fully contiguous intrastromal effect could be obtained.7,8

Femtosecond lasers work much more precisely than ps and ns lasers. Using femtosecond laser pulses, precise intrastromal cuts in animal cadaver eyes and human eyes can be obtained.9-16 The American company InterLase Inc. provides a femtosecond laser system (INTRALASE™ FS, Intralase, Irvine) operating with a 3-µm spot at 1053 nm for the optical production of flaps, based on photodisruptive effects.17

However, destructive collateral effects have been reported, based on the formation of large gas bubbles and their strong disruptive mechanical effects. Bubble

diameters of more than 12 µm have been hypothesized.15 Also, strong destructive mechanical shock waves can occur. A third side-effect is uncontrolled beam propagation towards the inner eye by a nonlinear effect known as self-focussing. All these destructive collateral effects have been observed with amplified femtosecond laser systems providing high laser peak energies in the range of microjoules (µJ) or even millijoules (mJ). The intensity of these destructive effects changes in parallel with the pulse energy. If it were possible to use less strong pulses at lower pulse energy, by maintaining the ability to ablate corneal tissue, this collateral damage should be reduced or even avoided.

In this chapter, we report on the use of nanojoule (nJ) and sub-nJ NIR femtosecond laser pulses of compact non-amplified laser sources for precise nanoprocessing of biological structures without significant collateral damage.

Using NIR femtosecond laser pulses at low subnJ pulse energy, we were able to achieve cut sizes of sub-200 nm in biological structures (Fig. 2).18 A minimum full-width half-maximum (FWHM) cut size of 85 nm was determined after femtosecond laser treatment of human chromosome 1. This amazing low cut size is even below the diffraction limited laser spot size. Multiphoton effects which occur in the central part of the spot only make such cuts and holes possible below the size of the illumination spot. Therefore, NIR femtosecond lasers enable similar or better nanoprocessing of cells and tissues than UV lasers, although the diffraction-limited spot size at 800 nm is a factor of four larger than the value for a high quality 193-nm laser beam (spot size changes with wavelength).

Using nJ and sub-nJ NIR femtosecond pulses, it was possible to optically knock out single organelles and the dissection of chromosomes within living animal cells, without collateral damage to the surround-

ing organelles or cellular membranes.19,20 We were also able to drill transient holes in cell membranes, and to realize optical gene transfer and efficient targeted transfection with foreign DNA.21

After modification of our femtosecond laser system, this nanotechnology was used for the first time to realize precise femtosecond laser surgery of corneal structures with an intrastromal cut size in the sub-micron range.22 But not just ultraprecise surgical corneal procedures can be performed with NIR femtosecond laser pulses of low pulse energy, we also used the same system to realize three-dimen- sional (3D) optical tomography of the cornea, with subcellular resolution based on multiphoton-excited autofluorescence and other non-linear effects.

In contrast to excimer lasers where ablation occurs layer-by-layer and the ablative effect is confined to the outermost structure, intratissue femtosecond laser procedures require sophisticated monitoring systems in order to make sure of the right location of the focal volume at the target structure of interest.

Such a system can be based on optical coherence tomography (OCT), however, the typical spatial resolution is in the range of 10-100 µm. Better subcellular resolution can be obtained with confocal scanning microscopes based on reflected backscattered light (for a review, see, Koester,23 this volume). In order to obtain sufficient depth resolution, out- of-focus photons have to be suppressed. This can be done by the introduction of a spatial filter (pinhole). By changing the focal plane, optical sectioning (optical tomography) can be performed. A major advantage of this confocal imaging technique is the possibility of using low-cost NIR laser diodes as light sources. Disadvantages include the use of a precise adjusted pinhole, low photon collection efficiency, artifacts due to the detection of some scattered out- of-focus photons, and limitation to signals due to changes in the refractive index. Today’s laser surgical systems have no integrated high-resolution 3D monitoring devices.

In this chapter, we also report on a pinhole-free method of optical sectioning based on multiphoton excitation of endogenous biomolecules. The use of NIR femtosecond laser pulses allows the detection of fluorescence photons emitted from endogenous fluorophores, such as the coenzyme NAD(P)H which emits in the blue/green spectral range. Collagen can also be selectively imaged. This multiphoton imaging method with a spatial resolution of 1 µm or better, provides different information than confocal reflection microscopy. In particular, multiphoton imaging allows highly sensitive functional imaging. NAD(P)H, which is mainly located in mitochondria, acts as a sensitive bioindicator because only the reduced form of this coenzyme is fluorescent (neither NAD nor NADP are fluorescent).24

In general, fluorescence detection is a very sensitive method compared to the detection of transmitted or reflected photons. In principle, single fluorescence photons can be detected.

We used our 3D diagnostic tool for multiphoton imaging with subcellular resolution to determine the target of interest for nano-/microsurgery as well as to control the surgical procedure. In particular, NAD(P)H and collagen have been imaged. After localization of the intratissue target by multiphoton imaging, laser ablation was performed with the same system. Immediately after laser treatment, optical sectioning was carried out in order to ascertain the surgical effect.

Femtosecond laser systems

Ultrafast laser systems emitting femtosecond pulses are considered to be the highly-precise surgical and diagnostic laser tools of the future. Potential fields of application for femtosecond laser surgery include treatment of the cornea and other ocular structures, the teeth, inner ear, and brain. The first clinical studies with femtosecond laser pulses have been conducted for the optical production of corneal flaps in LASIK.16 In a variety of experiments in material sciences, it was shown that femtosecond lasers may generate a more precise cut or hole in a variety of materials than nanosecond or picosecond laser pulses.

So far, the shortest laser pulse produced is of the order of attoseconds (1 as = 10-18 s).25 Attosecond laser pulses have been realized in the X-ray spectral range. The shortest pulses obtained in the NIR spectral range are about 3 fs (3 × 10-15 s). In this chapter, we report on applications of ultrashort pulses in the range of about 100 fs. During 100 fs, light travels a distance of 30 µm in air, within the cornea (n = 1.4) 21 µm. Ultrashort laser pulses are not monochromatic. For example, the spectral bandwidth of a 50-fs pulse is about 19 nm. This broad spectrum has to be considered in the choice of coated optics. Due to optical dispersion effects, the pulse width of the femtosecond laser pulse will be increased during transmission through glass. For example, the transmission of a 100-fs laser pulse through a typical microscope results in a pulse width at the target of about 170 fs, whereas an incident 10-fs pulse will have a pulse width at the target of more than 1 ps. The additional use of pulse compression units can realize a pulse width at the target which is nearly as good as at the laser output.

Femtosecond lasers as sources of pulsed infrared radiation belong to high-risk class IV.

The most popular femtosecond laser is a modelocked solid state laser with a titanium (Ti)-doped sapphire as the laser medium. The Ti3+ ion is responsible for the action of the laser in the NIR spectral range. These lasers are capable of producing pulses of 100 fs or less at a high repetition rate of about 80 MHz (80 million pulses per second). Recent developments include turn-key, compact, sealed tuneable Ti:sapphire laser systems which include the pump laser system (diode pumped frequency doubled neodymium:yttrium vanadate Nd:

Fig. 3. Pulse train of a femtosecond laser at 80 MHz with 100fs pulse width and 1-W mean power. L: length of laser cavity, c: velocity of light.

YVO4 cw laser at 532 nm) and the Ti:sapphire laser oscillator in one box. Ti:sapphire lasers can be tuned over a wavelength range of between about 690 and 1060 nm. For example, the compact system

MaiTai (Spectra Physics Inc., USA) has a typical tuning range of 780-920 nm, the Chameleon system (Coherent Inc., USA) of 720-930 nm.

In a non-amplified, 1-W, mode-locked Ti:sapphire laser system, the output pulse width is in the range of 60-100 fs and the repetition rate is about 80 MHz (Fig. 3). The pulse energy is 1 W/80 MHz = 12.5 nJ.

Further developments in femtosecond laser technology include the use of other laser materials such as Cr:LiSaF, neodymium-doped yttrium lithium fluoride (Nd:YLF, emission at 1047 nm), and the construction of frequency-doubled fiber lasers at 780 nm.

So far, material processing including cornea surgery has been conducted with high-pulse-energy amplified femtosecond pulses in the µJ to mJ range, and µm-sized illumination spots. In order to obtain such high energy pulses, the output beam of the Ti:sapphire laser oscillator has to be amplified. A typical amplified femtosecond laser system consists of the pump-laser unit, laser oscillator, a pulse stretcher to transform the pulses into picosecond pulses that do not harm the amplifier material, the amplifier with additional pump source unit, and a pulse compressor. The repetition frequency of amplified

systems is in the order of kHz (Fig. 4). Amplified laser systems are complex, expensive, and their use requires a laser expert.

In this chapter, we focus on non-amplified 80 MHz fs laser systems emitting picojoule (pJ) and nanojoule (nJ) laser pulses.

Multiphoton effects

Nanoand microsurgery, as well as fluorescence diagnostics of cells and tissues with NIR femtosecond laser systems, are based on multiphoton effects.

Multiphoton effects, in particular two-photon absorption, were predicted in 1931 by a young woman, Maria Göppert-Mayer, in her doctoral dissertation (Fig. 5).26 Her supervisor was Max Born. GöppertMayer hypothesized that one molecule can simultaneously absorb two photons (two-photon absorption) within a short temporal window (100 as). This means that two low-energy NIR photons can be absorbed by the molecule, and can lead to an excited state that would normally require high-energy UV or blue photons. From this excited state, the molecule can emit fluorescence in the visible range. Therefore, NIR photons can induce visible fluorescence in the blue, green, yellow, and red spectral ranges (Fig. 6). Because the process of fluorescence depends on absorbed photon pairs, there is a nonlinear, squared dependence between fluorescence intensity and incident light intensity or laser power, respectively. Göppert-Mayer’s, the later Nobel prize laureate, two-

Fig. 5. Photograph of Maria Göppert-Mayer, who predicted multiphoton processes in her PhD thesis. (Source: University Goettingen)

photon hypothesis was in direct contrast to the dogma that the fluorophore has to emit at a longer wavelength than the fluorescence excitation light. It took 30 years to confirm her ideas, due to the low probability of such a quantum event. Photon fluxes of about 1024 photons per second and per square centimeter are required, which could not be provided until the laser was invented in 1960.

In 1961, Kaiser and Garret reported first two-pho- ton excited fluorescence using CaF2:Eu2+ crystals.27 Another two-photon effect, the second-harmonic generation (SHG) was found by Franken et al.28 SHG occurs in certain non-centrosymmetric molecules and produces light at exactly half the incident laser light (Fig. 6). In contrast to fluorescence light, SHG occurs immediately (within femtoseconds compared to nanoseconds), it has a low spectral bandwidth and the same direction as the incident light.

In 1970, Rentzepis et al. produced images of threephoton excited fluorescence in organic dyes.29 Based on these studies, Sheppard and Kompfner suggested the utilization of multiphoton-induced fluorescence and harmonic generation for a new type of microscopy, nonlinear microscopy.30 In 1978, they wrote: “In the scanning optical microscope, nonlinear interactions are expected to occur between the object and highly focused beam of light, which we hope will open new ways of studying matter in microscopic detail hitherto not available.” Their idea was to use highly focused laser beams which provide the required high-photon flux density in the focal volume. By means of a scanning unit (flying-spot-tech- nology), nonlinear excitation within the focal volume can be used to probe the target in three dimensions.

In 1990, Denk et al. from Cornell University produced the first multiphoton microscope for applications in life sciences, based on femtosecond laser pulses from a dye laser at high repetition frequency.31 Their studies on two-photon excited fluorescence revolutionized fluorescence microscopic imaging.

Twoand three-photon excitation of fluorescence in NIR require MW/cm2 and GW/cm2 laser intensities which are produced within a sub-femtoliter focal volume of a high numerical aperture objective. Using 80 MHz sub-200 fs laser pulses, mean powers at the target of 10 mW or less of the tightly focussed laser beam are sufficient. The probability of twophoton excitation decreases with the forth power of the distance from the focus. There is no out-of-focus nonlinear effect, such as generation of VIS fluorescence or photodamage. This is in contrast to conventional one-photon imaging in which reflected photons and fluorescence arise within the entire illumination cone and in which pinholes are required for optical sectioning in order to suppress out-of- focus photons. A major advantage of multiphoton imaging is the possibility of pinhole-free optical sectioning by (i) moving the tiny multiphoton excitation volume across the target of interest, and (ii) the efficient collection of fluorescence and SHG photons with respect to the intratissue position of the excitation volume.

The efficiency of two-photon excitation and the fluorescence yield emulates the following relationship:31

n ≈ P2α / (τf2) . π2NA4 / (hcλ)2

with n = the number of absorbed photon pairs, P = mean power, α = molecular two-photon absorption coefficient, τ = pulse width, f = repetition frequency, NA = numerical aperture, h: Plack quantum of action, c: velocity of light, and λ = wavelength.

Because the fluorescence yield depends on a P2/τ relationship in two-photon microscopy, the efficiency increases for high peak power and low pulse width.

Despite the high light intensity, safe imaging can be performed below certain intensity thresholds, as demonstrated on laser-exposed single hamster ovarian cells and their next generations.32 Squirrel et al. imaged whole hamster embryos for 24 hours with a femtosecond laser scanning microscope without any damage to the embryos, in contrast to shortwavelength visible light.33

However, when the light intensity is increased to TW/cm2 light intensities, immediate destructive effects occur. The intensity is high enough to induce optical breakdown and plasma formation. At 800nm laser wavelength, four-photon absorption is sufficient to induce ionization and the formation of quasi-free electrons, leading to optical breakdown and plasma formation in water, and organic molecules (Fig. 6). Within the plasma, material is removed due to high temperatures, and plasma-filled bubbles are formed. However, there is another optomechanical effect causing photodisruption. Photodisruption is based on rapid expansion of the laser-induced plasma with Gigapascal pressures and the development and further collapse of cavitation bubbles, accompanied by the formation of destructive shock waves. So far, plasma-mediated ablation of the cornea by highenergy ultrashort laser pulses is mainly due to photodisruptive effects. Photodisruption also leads to destructive effects outside the illumination spot and causes undesired side-effects.

In order to obtain a desired highly-localized destructive effect without significant collateral damage, small bubble diameters and low optomechanical effects in the surroundings are required. The threshold for optical breakdown in water is decreased by a factor of more than 100 when comparing 100-fs with 3-ns pulses accompanied by less transformation into destructive mechanical energy (factor 6).13 Kurtz et al. reported that femtosecond pulses require about one-tenth the energy of picosecond and nanosecond pulses to produce corneal disruption.14 But also the photodisruptive effects of µJ femtosecond laser pulses count as destructive effects outside the illumination volume. In addition to optomechanical effects such as the formation of large cavitation bubbles, gas bubbles, and shock waves, self-focusing effects occur. Self-focusing effects based on laser-induced modulation of the refractive index of the illuminated

tissue occur at high pulse energy and weak focusing, and lead to uncontrolled cutting effects in the direction of the incident laser light. Self-focusing and photodisruptive effects depend linearly on pulse energy.

It would helpful to induce multiphoton effects, including optical breakdown and plasma-mediated material ablation, with femtosecond laser pulses at pulse energies as low as possible, to realize highlylocalized destructive effects within the illumination spot only.

Experimental set-up and biological samples

Preliminary studies on nanoprocessing and multiphoton imaging of the cornea with pJ and nJ non-am- plified femtosecond laser pulses have been conducted with an inverted femtosecond laser scanning microscope (FLSM, JenLab GmbH, Jena, Germany; Fig. 7). This microscope is based on the Zeiss LSM 410 system. The 80-MHz, 1-W beam of a tunable MaiTai laser was introduced into the FLSM via the LasCon interface, which consists of a beam expander, fast shutter, beam attenuator, power control, synchronization unit, and mirrors for beam adjustment. The FLSM contains baseport detectors, such as photomultipliers (PMT) and CCD cameras. Autofluorescence and SHG images were obtained by processing the PMT signals depending on the position of the galvomirrors (Fig. 5). In order to avoid the detection of backscattered laser light, two short-pass filters SP 730 (Chroma Techn., USA) were placed in front of the PMT. In part, the Ti:sapphire Vitesse laser (Coherent Inc., USA) at 800 nm was also used.

Transillumination of light from a halogen lamp through the whole ex-vivo porcine eye was monitored on-line with a baseport CCD camera. The laser beam was focussed on its sub-micron diffractionlimited spot size by a 40× objective of 1.3 numerical aperture (NA). The pulse width at the sample was determined to be 170 fs pulses by autocorrelation techniques. For two-photon autofluorescence

Fig. 7. The femtosecond laser scanning microscope from the JenLab GmbH, consisting of the MaiTai compact laser system, an interface between the laser and microscope, and the modified LSM 410 microscope.

Fig. 8. Fresh porcine eyes were chemically marked using AgNO3 (upper part). The eyes were placed in special eye chambers (JenLab GmbH).

imaging, a mean power of 5-10 mW at the target surface was used. For corneal processing, a beam with a mean power of 80 mW was used. This power corresponds to 1 nJ pulse energy.

We studied the effect of the ablation of corneal structures on porcine eyes that were placed in special tissue chambers with a 170-µm glass window (MiniCeM-biopsy, JenLab GmbH, Jena). In order to localize the imaged and laser-processed structures after laser treatment, parts of the cornea were marked with silver nitrate (Fig. 8).

Multiphoton imaging

Optical sectioning of corneal structures was performed by beam scanning using fast galvomirrors and a piezo-driven z-stage. A typical scan in one focal plane consisted of 512 × 512 pixels and took one or eight seconds. The illuminated field covered 320 × 320 µm = 0.1 mm2. The intratissue focal plane was varied in z-steps of 1 or 5 µm.

Figure 9 demonstrates a typical intracorneal image showing two-photon excited autofluorescence. It is obvious that multiphoton imaging provides subcellular spatial resolution. The fluorescence arises mainly from mitochondria. Using a variety of broadband filters in the detection path, maximum emission was found to be in the blue/green spectral range. Very likely, the origin of this autofluorescence is mainly determined by the reduced coenzyme NAD (P)H. The nuclei are non-fluorescent. Different cells can be clearly differentiated between, as can nuclear

Fig. 9. High-resolution image of two-photon excited autofluorescence of the cornea at 20-µm tissue depth. The autofluorescence is based on NAD(P)H located in the mitochondria of epithelial cells.

area and cytoplasm. Images from different tissue depths are depicted in Figure 10. In addition, the dependence on excitation wavelength is shown in the same figure. The 12 images in the figure represent only some out of a large stack of images. As shown by these representative optical sections, 3D multiphoton-induced autofluorescence imaging enables the various corneal tissue layers (epithelium, Bowman’s layer, stroma) and individual cells and collagen structures to be distinguished. By tuning the laser wavelength to 750 nm, autofluorescent mitochondria can clearly be seen. The mitochondrial fluorescence became less intense when the laser was tuned to longer excitation wavelengths. This behavior correlates with the expected two-photon fluorescence excitation spectrum of NAD(P)H. Because the onephoton absorption band is around 340 nm, the optimum two-photon excitation wavelength should be 750 nm or shorter.

Significant changes of the images at 750, 790, and 830 nm occur in the depth range of the junction between the epithelial layer and stroma, and in deeper tissue. The image at 750 nm represents autofluorescence signals, the 830-nm image mainly SHG radiation, while the image taken at 790 nm exhibits a mixture of fluorescence and SHG signals. The SHG

signal arises from the non-centromeric molecule collagen. Significant SHG signals cannot be obtained from structures within the epithelial layer. SHG imaging can be used to determine clearly the onset of stromal tissue with an axial resolution of about 1 µm. An SHG signal was found if excitation wavelengths of 790 nm or higher were used. In principle, SHG radiation also occurs at laser wavelengths shorter than 790 nm. However, no SHG signal in the UV can be imaged with this experimental set-up due to the reduced UV transmission of short-pass filters in front of the detector. For that reason, the 750-nm signal represents autofluorescence without contributions from SHG radiation, while the 830-nm signal is mainly influenced by backscattered SHG light from collagen structures at 415 nm.

Nanoand microprocessing

Autofluorescence and SHG imaging were used to find a particular intratissue target region of interest, and to ‘park’ the laser beam there. In order to perform a precise single intrastromal cut, the mean laser power was increased to 80 mW (1 nJ pulse energy). The intense 80-mW beam was scanned along one 512 pixel line at a typical beam dwell time on a pixel of 4 µs. This corresponds to about 320 pulses and a real exposure time of 320 × 170 fs = 54 ps per pixel. The photomultiplier became saturated (gray level: 255) during the laser exposure, probably due to plasma luminescence. Therefore, recording the PMT signal during laser treatment provides on-line information about multiphoton-induced optical breakdown and plasma formation. Interestingly, seconds and minutes after the laser treatment, the effect of intense laser exposure can also be imaged with high spatial resolution at a lower laser power. An intratissue, highly-fluorescent structure of sub-micron lateral size was formed along the cut. Figure 11 shows multiphoton images of stroma after the laser procedure along five lines. The region of intense laser exposure can clearly be seen by the luminescent lines with a typical dimension of 0.8 ± 0.4 µm.

In order to study the femtosecond laser effects more carefully, laser-treated eyes underwent histological examination. Upon analysis of hematoxilin/ eosin (HE)-stained cryosections, ultrathin sections revealed precise intraocular sub-micron cuts without collateral damage (Fig. 12). In part, the cuts even went through single nuclei of individual cells without any further visible damage. Laser scanning microscopy of these sections revealed a cut size of about half a micron. This corresponds to the resolution limit in VIS/NIR light microscopes.

In order to obtain more precise information on the cut size and quality of the NIR femtosecond laser procedure, scanning force microscopy and electron microscopy (EM) were conducted. Figures 13 and 14 demonstrate EM images obtained after a laser procedure in which a cube of material was ablated. To obtain such a cube, the bottom was first prepared by scanning a squared region deep into the stroma.

Fig. 11. Nanoprocessing along line scans. The effect of the surgical procedure can be monitored by multiphoton imaging. After laser treatment, an intratissue highly luminescent structure of sub-micron lateral size is formed along the cut.

Fig. 12. Histological examination of HE-stained cryosections after laser exposure by 488-nm laser scanning microscopy reveals precise sub-micron line cuts. Even a single nucleus at 90-µm tissue depth can be clearly cut through without visible collateral damage (image in the left corner).

Thereafter, the four walls were prepared up to the surface by x,z scans and y,z scans. As can be seen in the figures, material can be removed by relatively precise laser cuts. At higher magnification (Fig. 14), the image demonstrates in detail that collagen fibers can be cut through without any significant, large damage zones. There is a small visible layer along the cut of about 100 nm in size, which could be formed as the result of photothermal or photome-

Fig. 13. Electron micrographs of laser-treated stromal and epithelial tissue.

Fig. 15. During line scans with a power near the threshold for optical breakdown, some small bubbles with diameters of less than 5 µm can be monitored.

Fig. 14. REM images demonstrate clear cuts through collagen fibers.

chanical effects. The origin of these tiny changes in tissue structure still have to be investigated. Nevertheless, the image demonstrates that no significant signs of collateral destructive effects were present.

In addition to the removal of material, the successful removal of small intratissue structures, such

as single intraocular cells, was realized by single- point-illumination.

During and immediately after laser treatment, the formation, localization and lifespan of intratissue bubbles were monitored by video. Typically, between three and seven bubbles occurred as fluorescent sites along a 320-µm line. Using microsecond beam dwell times per pixel, a maximum mean diameter of 5 µm of intrastromal bubbles with a mean lifespan of 1.8 ± 0.3 s were recorded. The relatively long lifespan corresponds to observations by others who noted long-lived gas-filled bubbles (oxygen, hydrogen, methan) compared to short-lived cavitation bubbles. The possibility to remove intratissue material without destruction of the epithelial surface is demonstrated in Fig. 16.

Conclusions

These preliminary studies clearly show that nJ NIR femtosecond pulses at TW/cm2 intensities of nonamplified compact MHz lasers have the potential for

highly-precise, intratissue processing. No significant collateral damage by optical or optomechanical effects, such as intratissue self-focusing and photodisruption, were seen, in contrast to amplified NIR femtosecond laser surgery with µJ pulse energies.

Using fs 80 MHz laser pulses at pJ pulse energies, high-resolution optical tomography of the cornea can be performed, based on two-photon excited NAD(P)H autofluorescence and SHG formation in the collagen structures.

It should be possible to build a compact femtosecond laser system for both diagnostic and therapeutical eye procedures. Further studies, including in-vivo animal studies, still have to be conducted in order to discover the optimum pulse energy, pulse length, and spot size for refractive intracorneal surgery.

Acknowledgments

We would like to thank Oliver Krauss for his experimental contributions to this chapter. Helmut Hörig performed the EM studies, Mrs Möller and Mrs Hitschke the histology. This work was supported by the German Science Foundation (DFG, KO1361/10-3), the Ministry of Science, Research and Art of the State of Thuringia, and the German Ministry of Science, Research and Technology (BMBFT 01ZZ0105).

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