Other Imaging Techniques

3.3.1 Magnetic Resonance Imaging

Magnetic resonance imaging is a non-invasive high-resolution imaging technique. Magnetic reso-nance imaging is of interest in the context of this thesis as the developed techniques are in principle applicable to this imaging technique as well. Magnetic resonance imaging was recently used as a dynamic imaging modality to acquire sequences of two-dimensional cross-sectional scans of a beat-ing heart at a20mstemporal resolution.^{195, 196} The spatial resolution was1mm×1mmwith a field of view of25cm×25cm. Despite recent advances, dynamic imaging of three-dimensional volumes using magnetic resonance imaging at real-time speeds remains challenging.95, 131, 169 The penetration depth of magnetic resonance imaging is large, its spatial resolution about∼1mm³.

3.3.2 Computerized Tomography

Computerized tomography uses x-ray radiation with wavelengths in the order of10⁻⁷mto10⁻⁹m to penetrate biological tissue and produce tomographic scans. Computerized tomography is used for static, structural imaging of prepared hearts with very high spatial resolutions, large fields of view and penetration depths of many centimeters.¹⁶ Computerized tomography with micron resolution capabilities was used in this work to obtain three-dimensional structural scans of excised rabbit hearts, see figure 3.2 and section 4.4 in chapter 4. Because the technique employs ionizing electromagnetic radiation, it is a less attractive candidate for the use for electromchanical wave imaging as ultrasound or magnetic resonance imaging.

(a) (b)

Figure 3.2:Computerized tomography scan of formalin-fixed rabbit heart with µm-resolution: (a) volume rendering with steep opacity transfer function showing outer surface of heart, (b) combination of orthoslices showing ventricular walls and chambers and translucent grey surface of tissue extracted from image segmentation

3.3.3 Transillumination and Optical Tomography

Transillumination and so-called diffusive optical tomography, multiplicative optical tomography¹³⁸ and laminar optical tomography¹⁴² was introduced by Pertsovand co-workers.47, 75, 97, 138, 142, 204 It refers to a series of experimental imaging techniques that were developed with the intend to visual-ize intra- or transmural electrical activity in the heart employing optical techniques. The techniques relate to optical tomography, and employ similarly as fluorescence imaging light-tissue interactions and fluorophores. Due to the low penetration depths of light, however, these techniques were not suc-cessful in providing in-depth visualizations of transient wave phenomena inside and throughout the heart walls. To achieve higher penetration depths, high-intensity focused light was used and directed together with scanning schemes into the tissue. Electrical activity was reported to have been imaged inside the ventricular wall of a mouse heart.¹⁴² However, the single light point scanning scheme lead also to slow acquisition speeds, making it impossible to capture non-repetitive transient wave phenomena. In the context of imaging cardiac activity, transillumination and optical tomography techniques are purely scientific . No clinical applications were reported.

3.3.4 Electrode Recording

Electrode recordings are a standard tool to access the electrophysiological activity of cardiac tissue.

Spatio-temporal patterns of electrical activity can be recorded with hundreds of electrodes situated on the heart surface at high speeds with temporal resolutions<1ms. Common implementations are multi-electrode arrays¹⁷⁶ or ECGi.^{109, 137} However, while the spatial resolution of these methods is low and no transmural information is provided, it is likely that the presence of the electrodes modifies the patterns of electrical activity. In many cases, optical mapping techniques are favored.

3.3.5 Optical Coherence Tomography

Optical coherence tomography (OCT) is a high-speed tomographic imaging modality employing light and interferometry techniques to obtain structural information of optically penetrable tissue

Chapter 3. Imaging of Electromechanical Wave Activity in the Heart

volumes. OCT is typically used to obtain two-dimensional, cross-sectional or three-dimensional, volumetric data of biological tissue. OCT is especially suited for morphological tissue imaging due to its fine spatial resolution (<10µm) and its reasonable penetration depth of severalmm. The spatial resolution of OCT scans can be of comparable quality to histological cuts with sub micrometer resolutions. OCT is notably applied in ophthalmology where it is used to obtain detailed two- and three-dimensional structural scans of the eye, including cornea and retina, see Gora et al.¹⁷⁹ OCT was also shown to provide sufficient acquisition speeds to acquire cardiac motion appropriately and be able to penetrate the ventricular wall of a mouse heart.¹⁹⁹

Fundamentals

Analogous to ultrasound OCT uses the delay timeτ in which a light wave propagates through the tissue and is backscattered by sites in the tissue to determine the distances of those sites. However, due to the speed of light, the delay timeτ cannot be determined electronically, but is determined by interferometric techniques. OCT images are then generated based on the variation of optical tissue properties from different structures. An introduction to optical coherence tomography is given by Drexler et al.¹⁶² and others.66, 82, 124

The time delayτ is determined from the auto-correlation function Γ(τ) in the interference process of an electromagnetic light wave E with itself. The basic setup of an OCT imaging system is a Michelson interferometer. A coherent, broad-bandwidth laser beam is split into a reference and a probe beam by a beamsplitter. The probe beam E1 is directed into a tissue sample from where it is reflected back and reintroduced into the optical path and recombined with the reference beam.

The reference beamE₂ stays in the Michelson interferometer. Both electromagnetic wave fieldsE₁ andE2 interfere and we obtain a superposition of the two electromagnetic waves E1 andE2. The intensityI_D ∼E² on a detector is given by:

ID = < E1∗

E2 >

= I1+I2+Re{< E₁^∗E2 >}

= I₁+I₂+Re{Γ(τ)}

The result is an interference pattern which depends on the complex auto-correlation functionΓ(τ).

The electromagnetic light waveE₁ having traveled into the tissue and back experienced a delay of τ = ^2d_c compared to the reference light waveE2 that stayed in the interferometer. Actually, both electromagnetic waves are the same electromagnetic wave field as they were delivered by the same light source and we setE₁=E₂=E:

I_D = < E^∗E >

= 2I+Re{Γ(τ)}

That is, a part of the electromagnetic light wave E experiences a temporal shift τ and interferes with itself. The intensity on a detector depends on the time delayτ. The complex auto-correlation

function, also referred to as complex self-coherence function,⁹²is given by:

Γ(τ) = lim

T→∞

1 T

T

Z2

−^T₂

E^∗(t)E(t+τ)dt

For a harmonic, monochromatic, linear waveE(t) =E₀e^−iωt the complex self-coherence function Γ(τ) = |E₀|²e^−iωt depends harmonically on the time delay τ. Furthermore, the intensity on the detectorID(τ) = 2I(1 + cos(ωτ))depends harmonically on the time delayτ which depends in turn on the distancedthat the probe beam needed to travel through the tissue. When translating the end mirror of the reference arm of the Michelson interferometer and subsequently shifting the reference light wave against the probe light wave we would repeatedly obtain destructive and constructive in-terference alternating periodically with the wavelength of the light. This means, that with completely coherent light one would not be able to obtain depth resolved tissue information by scanning through the tissue. OCT generally employs broad bandwidth light with a short coherence length lc. This results in an interference signal on the detector with gaussian input:

I_D ∼ Re{Γ(τ)}

= Re{e^−iω0∆τe^−σ

2ω∆τ2

2 }

The auto-correlation function of a gaussian function is again a (modulated) gaussian function. Con-sequently, when employing light with a short coherence lengthlcwe obtain constructive interference only when the length of the reference arm matches the distancedthat the light travelled in the probe arm. Consequently, when scanning the reference arm we can determine the distance of the site in the tissue that reflected the probe beam back into the Michelson interferometer. The distance is indicated by the one distance at which constructive interference occurs.

Technological Implementations

OCT imaging systems are presently realized in a variety of technological implementations. The ma-jority of OCT imaging systems employ single scanning point or confocal scanning schemes. They use a single focused laser beam to construct a cross-sectional B-scan or a volumetric scan from sub-sequent, adjacent A-scans. The beam is directed into the tissue via galvanometer scanners which can be programmed to scan arbitrary scanning patterns. A time-domain based OCT system em-ploys a mechanically translated end mirror in the reference arm to temporally shift the light against itself and to scan the different delay times τ and distances d in axial direction. A-scan rates of f_A≈0.1kHz−10kHzare typically achieved with time-domain based OCT and acquisition speeds are limited by the mechanical scanning technique.

Parallel or full-field OCT employs a different scanning scheme together with time-domain based OCT. Instead of scanning a single laser beam through the tissue the sample is full-field illuminated and en face imaged with a 2-dimensional CCD detector, hence acquiring many neighboring sites in parallel during one mechanical axial scan. Newer technological implementations of OCT imaging systems employ different, faster axial data acquisition techniques. In particular, the mechanical axial scanning technique is avoided.

Chapter 3. Imaging of Electromechanical Wave Activity in the Heart

Left Ventricle

Atrium 1mm

Figure 3.3:Mouse heart volumetric scan acquired with optical coherence tomography: high-speed FDML-OCT system operating in the near-infrared at wavelengths of λ = 1300nm. Due to the long wavelength the beam was found to be able to penetrate almost the entire thickness of the ventric-ular wall (indicated by red line) of the mouse heart.

In Fourier-domain based OCT systems the axial depth information is encoded in frequency space rather than in time. The axial depth information can be calculated via the Wiener-Khintchine the-orem between the auto-correlation function and the spectral power density by a Fourier-transform of spectra acquired when light with different wavelengths is backscattered from the tissue. One implementation of a Fourier-domain OCT imaging system is Swept-Source (SS) OCT. In SS-OCT the frequency is encoded in time as a scanning light source repeatedly sweeps over a broad band-width spectrum and the backscattered light is recorded sequentially by a photodiode. Instead of mechanically scanning with the reference arm through the tissue depth in axial direction, a frequency bandwidth is scanned through and its single frequencies come to a constructive interference at dif-ferent tissue depths. The wavelength having caused the constructive interference can be correlated with the spatial distancedof the site having backscattered the light. One sweep of the light source corresponds to one A-scan. A-scan acquisition speeds can be increased dramatically as the sweeping speed of the light source is faster than the mechanical translation of the reference arm. Another im-plementation of a Fourier-domain OCT imaging system is Spectral-domain (SD) OCT. In SD-OCT the full bandwidth of the broad bandwidth light source is sent into the probe arm at once and the single frequencies are spatially separated by a dispersive element before they are captured by a linear detector (CCD/CMOS) array. Broadband interference is acquired with spectrally separated detectors from which each detector captures a frequency range corresponding to a spatial depth range in axial direction. Accordingly, the number of used elements of the detector determines the axial spatial reso-lution. Fourier-domain OCT imaging systems operate with a single scanning point scanning scheme.

Performance: Speed and Penetration Depth

A-scan acquisition speeds can be increased dramatically using Fourier-domain OCT compared to time-domain based OCT as the read-out rates of the line scan camera detectors determine the A-scan rates of SD-OCT imaging systems. SD-OCT imaging systems typically operate with A-scan rates offA ≈10kHz−100kHz. The latest development in Fourier-domain OCT imaging is the use of FDML (Fourier Domain Mode Locking) lasers as light sources.¹²⁶ FDML lasers are ideal rapidly

sweeping light sources for OCT applications. They combine high repetition rates of conventional mode-locking lasers at a constant output power with a spectral chirp, and allow higher sweeping rates than conventional swept-sources. Consequently, FDML SS-OCT imaging systems are very promising candidates for high-speed applications of OCT imaging. Single scanning point high-speed OCT systems based on FDML technology can achieveA-scan repetition rates offA = 300kHz.

This translates to frame rates of about f_{f ps} = 100−1000Hz when scanning in two-dimensional imaging mode. The penetration depth of OCT is limited to a few millimeters in optically dense tissue as cardiac muscle tissue.

Clinical Applications

Clinical applications of OCT are limited to cardiovascular tomography. Intravascular OCT is a cost effective and more practical alternative to intravascular echocardiography.¹⁵² Transmural imaging of the heart is limited to single applications in cardiac research, see.⁹⁹ Elastographic imaging was demonstrated using OCT together with speckle-tracking techniques.⁶³However, elastographic imag-ing was not applied to cardiac tissue. The functional use of OCT in neuronal tissue was demon-strated by.⁹⁴ It was observed that optical scattering increases in electrically stimulated axons. Con-sequently, OCT is able to detect optical changes in cells and tissue during physiological processes.

OCT typically employs near-infrared light as scattering effects are less with longer wavelengths and the achieved penetration depth into optically dense tissue is larger. The optimal wavelengths for applications in cardiology range from 800nm−1500nm and are referred to as the therapeutical window.¹⁵²

3.3.6 Other Imaging Techniques

There is a bewildering landscape of imaging techniques and it is possible to identify many more tech-niques which have a certain relevance with respect to this work. Magnetic particle imaging uses the distribution of superparamagnetic iron oxide particles to acquire dynamic three-dimensional tomo-graphic data of cardiac tissue.^{117, 187} Temporal resolutions of∼20mswith three-dimensional fields of view of20mm×12mm×17mmwith a spatial resolution sufficient to resolve all heart chambers have been reported. Spatial resolutions are in the order of∼1mm³. Photoacoustic imaging is a hy-brid imaging modality consisting of a combination of ultrasound and optical imaging. It employs the photoacoustic effect in which non-ionizing laser pulses are used to deposit heat energy in tissue and to stimulate a transient thermoelastic expansion with a consequential wideband ultrasonic emission.

Acquisiton speeds are therefore mostly bound to the acquisition speeds of the ultrasound detection system. B-scan acquisition speeds off_B = 166Hzwith axial and lateral resolutions of25µmand 70µmrespectively and penetration depths as deep as severalmmwere reported.¹⁸⁵ Terahertz imag-ing employs electromagnetic light to penetrate matter. Terahertz imagimag-ing is much less affected by scattering due to its long wavelengths and would likely penetrate much further than infrared light into tissue. However, the high absorption of terahertz radiation by water limits the penetration depth to<1.5mmin biological tissues with high water contents. As of today, terahertz imaging is limited to applications in dermatology and dentistry.¹⁰¹ A variety of optical tomography imaging modalities have been developed, some of them with the particular aim to visualize the electrophysiology within the 3-dimensional cardiac muscle. Confocal microscopy is a typical example of a widely used opti-cal tomography imaging modality, which provides high spatial resolutions of just a few micrometers.

As this technique only illuminates one point in the tissue at a time, 2-dimensional or 3-dimensional imaging requires scanning over a regular pattern. This is a typical feature of most optical tomography

Chapter 3. Imaging of Electromechanical Wave Activity in the Heart

imaging modalities. However, confocal microscopy typically penetrates much less deeply, that is in the order of< 100µm, into tissue than the desired penetration depths of severalmm. This limits its relevance to applications in transmural imaging of the heart. Macroconfocal imaging uses a com-bination of confocal microscopy and laser scanning techniques to assemble raster images of tissue volumes. Acquisition times for acquiring the scans is in the order of minutes.