Chapter 1      6

Introduction:      6

Chapter 2      8

Optical coherence tomography (OCT) :      8

2.1      Imaging the human retina by OCT in ophthalmology:      8

Chapter 3      10

Basic principles of OCT imaging:      10

3.1      Optical interferometry and time domain OCT:      10

3.2       Low-coherence interferometry:      11

3.3      Spectral or Fourier domain OCT:      12

3.4      Axial and transversal resolution:      12

3.5      Polarization-sensitive OCT:      13

Chapter 4      15

The human eye:      15

4.1       Structure and function of the human retina:      15

4.2      The role of rods and cones:      17

4.3      The retinal pigment epithelium (RPE):      17

4.4      Bruch´s membrane:      18

4.5      Signal processing within the retina:      18

4.6      Retinal blood supply:      18

Chapter 5      19

Age-related macular degeneration (AMD):      19

5.1      Prevalence of AMD:      19

5.2      Risk factors for AMD:      19

5.3      Histopathology of drusen:      20

5.4      Appearance of drusen in fluorescein angiography:      21

5.5      Atrophic/dry AMD:      22

5.5.1      Symptoms:      22

5.5.2      Clinical findings:      23

5.9      Geographic atrophy:      23

5.6      Neovascular AMD:      24

5.6.1      Symptoms:      24

5.6.2      Clinical findings:      24

5.6.3      Angiographic characteristics of neovascular AMD:      26

5.7      Differential diagnoses of neovascular AMD:      28

5.8      Treatment of AMD:      28

Chapter 9      30

Conclusion:      30

References:      31

Curriculum vitae      41

Summary/Précis:

Age-related macular degeneration (AMD) is a chronic disease leading to severe visual deterioration and impairment of daily life. Two types of AMD are distinguished: Wet or neovascular AMD and dry AMD. Dry AMD is characterized by drusen (small yellowish retinal deposits) and atrophy of the retinal pigment epithelium (RPE); causal treatment is currently not available.

Wet AMD is marked by abnormal growth of blood vessels within the macular area (spot of sharpest sight) leading to fluid accumulation (edema) and retinal bleeding as well as scar formation. If the disease is not treated severe irreversible visual impairment will result. Treatment with new agents that are administered intravitreally and act against vascular endothelial growth factor (VEGF) are now available, impeding disease progression in many cases.

The present book gives a comprehensive overview about all aspects of AMD, useful for affected patients and interested readers and covers causes, epidemiological facts, diagnostics (emphasizing the topic optical coherence tomography-OCT) and treatment strategies of the disease, based on extensive literature research.

Imprint:

© 2018 Vienna, Austria, by Christopher Schütze; The present book is based on the doctoral thesis Schütze et al., 2014 found in the list of references.

Chapter 1

Introduction¹:

AMD is a disease affecting the macula, the part of the eye responsible for central vision. The disease mainly affects people over 50 years of age. Early and late characteristics of the disorder can be differentiated (Kanski et al., 2008b):

Early characteristics of AMD include discrete yellow deposits in the macular region (drusen), hyperpigmentation of the RPE, and well-circumscribed regions with RPE depigmentation. Late characteristics are presence of regions of circumscribed RPE atrophy (geographic atrophy,-GA) with exposed choroidal vessels in case of dry AMD. Pigment epithelial detachments (PED), or subretinal/sub-RPE neovascularization occurs in case of neovascular AMD (Kanski et al., 2008b).

The diagnosis of AMD disease is made clinically by ophthalmoscopy, fundus autofluorescence imaging (FAF), and fluorescein angiography (FA). Other diagnostic modalities include techniques like optical coherence tomography (OCT). FAF images autofluorescent retinal tissue (particularly metabolic by-products that are not metabolized adequately due to i.e. insufficient processing of phororeceptor outer segments and of other by-products) (Schmitz-Valckenberg et al., 2011). FA is an invasive diagnostic tool where fluorescein is injected into the patient´s vein and retinal changes can be observed using a special camera device.

Due to the ability of generating high-resolution images of retinal structures, the use of OCT has become increasingly important in clinical practice in the field of ophthalmology over the past years (Gabriele et al., 2011). Spectral domain (SD)-OCT devices generate high-resolution images of retinal morphology, reaching an axial resolution of up to 2.1µm in tissue (high-speed, ultrahigh-resolution Fourier domain OCT) (Wojtkowski et al., 2004).

In principle, OCT works similar to ultrasonography, however the technology does not apply sound during an investigation. Instead, the magnitude and echo time delay of backscattered light is measured on the basis of low coherence interferometry. Following scan acquisition, detailed cross-sectional analysis of retinal morphology is possible (Huang et al., 1991, Yun et al., 2003).

With the introduction of frequency domain imaging (Fercher et al., 1995), the acquisition of backscattered intensity signals has significantly improved (Wojtkowski et al., 2002, Choma et al., 2005, de Boer et al., 2003, Leitgeb et al., 2003, Nassif et al., 2004, Wojtkowski et al., 2005). A broad bandwidth light source and a spectrometer to record the interference signal are used (Wojtkowski et al., 2002, de Boer et al., 2003, Leitgeb et al., 2003, Wojtkowski et al., 2005).

Alternatively, in Swept Source (SS) OCT, a narrow-bandwidth source is swept through a broad range of frequencies, applying a photodetector unit (Choma et al., 2005, Zhang et al., 2005, Yun et al., 2004, Choma et al., 2003, Chinn et al., 1997, Lexer et al., 1997).

In SD-OCT imaging, axial scans are generated using the Fourier transform of the interference spectrum (Fercher et al., 1995, Drexler et al., 2012), thereby assuring fast A-scan collection (Gabriele et al., 2011). High detection sensitivity is gained by SD-OCT due to the fact that SD-OCT offers a high signal-to-noise ratio (de Boer et al., 2003, Choma et al., 2003, Leitgeb et al., 2003). Three-dimensional high-resolution images of the retina can be obtained by SD-OCT systems (Srinivasan et al., 2008). The majority of SD-OCT devices used in clinical practice today, is capable of gaining ~27,000 A-scans per second, reaching an axial resolution of ~5-6µm in tissue (Gabriele et al., 2011).

Besides detailed cross-sectional imaging of retinal morphology, SD-OCT provides the capability of detecting retinal layer thickness and retinal nerve fiber layer thickness in a quantitative

and qualitative way (Wojtkowski et al., 2005). Retinal diseases clinically investigated by SD-OCT include age-related macular degeneration (AMD), glaucoma, epiretinal membranes, central serous

chorioretinopathy, amongst others (Hee et al., 1995, Hee et al., 1995a, Hee et al., 1995b,

Hee et al., 1995c, Puliafito et al., 1995, Hee et al., 1996, Schuman et al., 1995, Schuman et al., 1995a, Gaudric et al., 1999, Chauhan et al., 2000, Massin et al., 2000, Massin et al., 2001, Spaide et al., 2002, Srinivasan et al., 2006, Kumagai et al., 2012).

Although intensity-based SD-OCT has substantial benefits in imaging retinal morphology and is capable of generating high-resolution retinal images, the method has limitations in the tissue-specific detection of retinal layers due to similar reflectivity of these anatomical structures (Ahlers et al., 2010). This is caused by the fact that SD-OCT has the exclusive ability to image backscattered light on an intensity basis (Ahlers et al., 2010, Chen et al., 2005).

With the recent implementation of polarization-sensitive OCT (PS-OCT), a new imaging method offering tissue-specific contrast visualization has been introduced and has been used in several recent studies (Yamanari et al., 2008, Pircher et al., 2006, Ahlers et al., 2010, Pircher et al., 2011, Götzinger et al., 2008, Götzinger et al., 2008a, Götzinger et al., 2005, Götzinger et al., 2009, Baumann et al., 2010, Schlanitz et al., 2011, de Boer et al., 1997, de Boer et al., 2002, Cense et al., 2004, Yasuno et al., 2010). PS-OCT has the ability of measuring the polarization state of backscattered light. The technique is based on the principle that certain tissues can change the polarization state of light (Pircher et al., 2011). Changes in polarization states are caused by different mechanisms regarding the interaction between light and exposed tissue (i.e. birefringence and depolarization) (Pircher et al., 2011). Birefringent tissue includes the ocular muscles, the sclera, tendons, the trabecular meshwork, the retinal nerve fiber layer (RNFL) (Cense et al., 2004, Götzinger et al., 2008, Baumann et al., 2007, Yamanari et al., 2008), or the corneal stroma (Lim et al., 2011, Götzinger et al., 2004). Further, Henle´s fiber layer (Brink and van Blokland, 1988) and scar tissue lead to birefringence. Depolarization is caused by melanin-containing structures (i.e. the RPE), due to intrinsic tissue-specific properties (Baumann et al., 2009, Pircher et al., 2011, Pircher et al., 2006).²

Chapter 2

Optical coherence tomography (OCT)³:

2.1      Imaging the human retina by OCT in ophthalmology:

OCT is a non-invasive and non-contact imaging modality capable of generating cross-sectional images of the retina, the optic nerve, and the vitreous. Further, anterior segment OCT is capable of imaging anatomical structures of the anterior section of the eye (i.e. the cornea, the anterior chamber, or the lens). OCT is analogous to ultrasonography, however the technology uses light waves for imaging instead of sound waves. During an examination the eye is scanned transversely using a light beam. Cross-sectional intensity-based images can be generated and analyzed by means of a false color or gray scale image. Using this technology, the echo time delay and the amount of reflected or backscattered light is measured using low coherence interferometry (Kanski et al., 2008).

Indications for retinal OCT imaging:

- Diagnostics of pathological alterations in the macular area such as macular holes, central serous retinopathy, epiretinal membranes, macular edema, or vitreomacular traction (Hee et al., 1995b, Hee et al., 1995c, Hwang et al., 2012, Hee et al., 1995a, Tsunoda et al., 2012, Kanski et al., 2008).

- Documentation of disease progression and treatment response, i.e. retinal thickness measurements in neovascular AMD and its development during treatment follow-up (Sulzbacher et al., 2013, Golbaz et al., 2011, Kanski et al., 2008).

- Differentiation of retinal detachment of longer duration and retinoschisis (Kanski et al., 2008).

- Retinal nerve analysis and evaluation of retinal nerve fiber layer (RNFL) thickness, i.e. in patients with glaucoma (Chauhan et al., 2012, Kanski et al., 2008).

There are several color and gray scales used for representing intensity-based retinal images in OCT imaging. To mention only one example, images acquired by i.e. Cirrus SD-OCT (Cirrus HD-OCT; Carl. Zeiss Meditec, Dublin, California, USA) are displayed in the following way: light colors usually represent highly reflective layers (red and white), whereas retinal structures with low reflectivity are visualized in dark colors (blue and black). Structures with intermediate reflectivity are illustrated in green color. The plexiform and nerve fiber layer are visualized in red, yellow, or light green color.

Using Cirrus SD-OCT imaging as a representative example, the inner (IPL) and outer plexiform layers (OPL) are represented in light green color. The inner (INL) and outer nuclear layers (ONL) usually appear in blue or black color. The junction between the outer and inner segments of photoreceptors is displayed as a thin red structure in the outer retina. High-resolution OCT (i.e. Cirrus OCT) is further capable of differentiating retinal structures like the external limiting membrane (ELM) and the ganglion cell layer (GCL) (Kanski et al., 2008) (Fig.1).

Fig.1 shows an SD-OCT image of a healthy retina, displaying a cross-sectional image of retinal layers.

In essence, the bandwidth of the light source used in OCT imaging determines the axial resolution (Drexler, 2008). Enhanced axial OCT resolution can be achieved when using broad bandwidth light or light sources with a low coherence length (Drexler, 2008a). Image resolution of modern OCT machines used in clinical practice ranges between 5-8μm (Drexler, 2008a).

Chapter 3⁴

Basic principles of OCT imaging:

OCT is an interferometric technique first introduced by Huang et al. in 1991 (Huang et al., 1991) that is used to detect the depth profile of layerd media (i.e. the retina). This imaging modality is capable of measuring the degree of reflection on an intensity basis. For this purpose a partially coherent light source with a short coherence length and a Michelson interferometer are used (Drexler et al. 2012a, Drexler et al. 2012b, Drexler et al. 2012c, Drexler et al. 2012d).

In principle, a mirror defers a replicate of the initial wave, which is subsequently overlaid with waves reflected by the investigated sample. Thereby, the depth profile of the investigated sample is contained in the resulting interferogram. The interferogram describes the measured intensity based on the movement of the reference mirror (in Time Domain-TD-OCT) (Drexler et al. 2012e).

3.1      Optical interferometry and time domain OCT:

In OCT imaging high-resolution measurements of tissue are performed by measuring the echo time delay of light backscattered from investigated tissue. In order to enable the measurement of optical echoes in OCT imaging, a correlation technique known as low-coherence interferometry is applied. Basically, this method allows for measuring distances to objects with high accuracy by measuring light that is reflected from these objects, and by correlating these measurements with light traveling a known reference path delay (Youngquist et al., 1987, Schuman et al., 2012, Born et al., 1999, Gilgen et al., 1989, Fercher et al., 1988).

Essentially, a light beam (light wave) is characterized by oscillating magnetic and electrical fields, varying periodically in space and time (Born, Wolf, 1999). Light is characterized by its wavelength and frequency (ν) spreading with a defined velocity. The velocity of light is determined by the medium in which it propagates. The definition of the velocity of light is c = 3 x 10⁸ m/sec in vacuum and is slower in media such as water, or for instance within the vitreous body. The velocity of light is defined by c/n, where n stands for the index of refraction of the respective medium. The axial delay in OCT imaging has to be scaled by 1/n due to the fact that it is correlated to the index of refraction, in order to determine the physical dimension (Schuman et al., 2012).

This principle is similar to ultrasound imaging where information about the sound-wave velocity is needed (Schuman et al., 2012).

When 2 light beams are emitted from an identical source and are combined, a phenomenon known as interference may occur, due to the fact that light is composed of oscillating magnetic and electrical fields. In this regard, constructive or deconstructive interference can occur. Constructive interference occurs when the fields of 2 light beams are in phase and therefore have a tendency to add. Consequently, this results in a light beam of greater intensity. On the other hand, when the 2 light beams are out of phase (by λ/2), this will result in cancellation, and the outcome is less light intensity. This phenomenon is termed destructive interference. The relative phase of the field´s oscillations is a decisive factor for them to add up constructively or deconstructively (Schuman et al., 2012a).

Light waves from 2 light beams can be added or interfered by an optical interferometer (Born, 1999) such as a Michelson interferometer (Fig.2). A beam splitter or a partially reflecting mirror contained in the device causes splitting of the incident light into 2 light beams. One of the beams functions as a reference beam, the other one serves as a sample beam. The reference mirror reflects the reference light beam, and the measurement light beam is back-reflected from the imaged sample/organ (i.e. the retina), thereby generating the sample light beam. The partially reflecting mirror of the interferometer is used in order to recombine the reference and the signal light beams. The length that is traveled in the reference arm will alter when the position of the reference mirror is changed. Subsequently, the time delay of the reference light will vary. The reference light may either add constructively or deconstructively (or any state in between) to the measurement light from the investigated sample. A photo detector unit then measures the intensity of the resulting output beam. Light with an electrical field summed up from the reference arm and the signal light beam is the final output at the interferometer (Schuman et al., 2012b).

3.2       Low-coherence interferometry:

In OCT imaging, the precise measurement of the absolute position of a structure within investigated tissue is fundamental. In order to fulfill this requirement, light with a short coherence length is used. Coherence is defined as a statistical property of light. Low-coherence light originates from an oscillating electric field, exhibiting fast statistical changes in phase, described as a function of space and time. Low-coherence light covers a large spectral wavelength range. There is an inverse proportionality of the coherence length to the bandwidth of the light. This means that the shorter the coherence length, the wider the spectrum´s wavelength range of the light source. (Drexler et al., 2001, Schuman et al., 2012b).

When using low-coherence light in an interferometer, it is only possible to gain interference if the path lengths of the signal and the reference path are closely matched. If this is not granted, the 2 beams are incoherent and subsequently no interference will occur (Schuman et al., 2012b).

For retinal OCT imaging, a superluminescent diode operating at ~840nm (near infrared wavelength) is typically used (bandwidth ~50nm). These properties enable an axial resolution of approximately 6μm in air. As light travels at less speed in tissue, the axial dimension has to be divided by the index of refraction in order to calculate axial resolution when imaging tissue, amounting up to a resolution of approximately 4.5μm. When using light sources with very short coherence lengths (equal to a broader bandwidth), the axial resolution of OCT imaging may be significantly improved (up to 1 to 3μm in tissue using laser light sources that have bandwidths between 100-200nm) (Drexler et al., 1999, Drexler et al., 2001, Povazay et al., 2002).

To summarize, when correlating backscattered light with the known delay of the reference arm accomplished by the interferometer, it is possible to precisely measure the magnitude and time delay of light. Therefore, as each depth profile is acquired as a function of time, this technique is termed time domain detection (Schuman et al., 2012c).

3.3      Spectral or Fourier domain OCT:

The axial scan depth profile is also determinable by analyzing the spectrum of the output from the interferometer, instead of evaluating the depth of backscattered light by assessing the reference delay (Schuman et al., 2012c). This method is known as spectral domain (SD) or Fourier domain detection since the year 1995 (Fercher et al., 1995) enabling faster imaging speeds (~50x to 100x faster than TD-OCT), and allowing for improved image quality. Significant advantages regarding sensitivity have been shown in SD-OCT imaging compared to TD-OCT technology (Choma et al., 2003, de Boer et al., 2003, Leitgeb et al., 2003). Further, 3-D volumetric datasets can be obtained. The majority of all commercially available and presently clinically used OCT systems are based on Fourier domain detection (Schuman et al., 2012d).

In principle, Fourier domain imaging works by encoding different spatial positions as different frequencies. Compared to TD technology, SD-OCT systems have a reference mirror in a fixed position, and the reference delay does not have to be scanned in order for gaining an axial depth profile. Instead, light that is backscattered from different axial depths is being detected simultaneously. A spectrometer is used in order to obtain the interference signal at the output of the interferometer. The spectrometer separates the backscattered light into its different frequency components applying an optical diffraction grating (Schuman et al., 2012d).

In order to obtain thickness information about an investigated sample in SD imaging, the Fourier transform is applied (Holz et al., 2013). In essence, the depth profile, which is backscattered, is calculated by applying the Fourier transform relationship between depth and frequency. Each modulation frequency of the interference spectrum corresponds to a certain depth in an A-scan (Wojtkowski et al., 2004, Schuman et al., 2012e, Davis, 2008). By using the Fourier transformation, a higher resolution and a better signal to noise ration is achieved as compared to TD detection (Schuman et al., 2012f, de Boer et al., 2003, Leitgeb et al., 2003, Choma et al., 2003).

The output interference spectra are recorded using a high-speed line scan camera (Schuman et al., 2012d) (axial scan repetition rate of 20,000 to 50,000 Hz) (Schuman et al., 2012e). Tremendous advancements in SD/Fourier domain detection regarding sensitivity are achieved, as all echoes of light are detected simultaneously and not sequentially as in TD imaging (Schuman et al., 2012f). Spectral/Fourier domain OCT operates ≈50 times faster than TD-OCT (Coscas et al., 2009).

3.4      Axial and transversal resolution:

In Fourier domain OCT imaging it is the coherence length of the light source that determines the axial resolution, analogous to TD-OCT (Schuman et al., 2012g).

The scanning optics are determining the transversal resolution, which is affected by the focal length of the respective imaging lens that focuses the sample beam onto the investigated sample, and further by the diameter of the beam in the sample path of the interferometer (Baumann, 2009b). Numerical aperture (NA) determines achieved resolution dependent on the properties of the scanning optics and the sample (Chelliyil et al., 2008, Reider et al., 1997).

Commercially used SD-OCT devices have a light source with a bandwidth of ~50nm and reach an axial resolution of approximately 5-6μm (Schuman et al., 2012h). Transverse resolution frequently used in commercially available SD-OCT devices is ~20μm, as there is a trade-off between