Download Simultaneous Dark-Bright Field Swept Source OCT for ultrasound

Simultaneous Dark-Bright Field Swept Source OCT for ultrasound
detection
Cedric Blattera, Branislav Grajciara, Boris Hermanna, Robert Huberb, Wolfgang Drexlera,
Rainer A. Leitgeb*a
a
Center of Medical Physics and Biomedical Engineering, Medical University Vienna,
Waehringerstr.13, A-1090 Vienna, Austria;
b
Lehrstuhl für BioMolekulare Optik, Ludwig-Maximilians-Universität München, Oettingenstraße
67, 80538 München, Germany
ABSTRACT
We introduce a swept source FDOCT imaging system that allows measuring simultaneously the reflected light and
scattered light (bright field) and the scattered light only (dark field) in two different channels through separate Gaussian
and Bessel detection. Specular reflections can then be used to obtain knowledge about the sample time evolution with
high SNR for phase analysis. Based on this configuration, we provide a proof-of principle study for resolving ultrasound
pulse trains with high temporal resolution on surfaces, which potentially provides a novel phase sensitive all optical
detection scheme for the combination of OCT with photoacoustic imaging.
Keywords: Dark field imaging, Bessel beam, Extended focus, FDML Swept Source, Multichannel detection,
Photoacoustic, Ultrasound, Phase sensitive.
1. INTRODUCTION
Fourier domain optical coherence tomography (FDOCT) combines high-speed imaging with high sensitivity and axial
resolution [1]. FDOCT has raised strong interest as biomedical imaging tool and is currently applied in many different
fields. Apart from its original ophthalmology application, it is for example used in dermatology or cardiology. For skin
imaging the sample is usually placed against a glass window in order to increase the stability during the measurement
and to maintain the sample position relative to the focus plane. The large refractive index change associated with the
transition glass-sample produces a strong reflex that might saturate the detector and reduce the dynamic range and
sensitivity. It is especially critical when using an analog-to-digital converter with a limited dynamic range, as used for
high-speed application. The use of index-matching gel might not be possible for specific applications and the associated
reflection attenuation might not be sufficient because of the strong sensitivity of OCT. The glass window is then usually
tilted to avoid the direct reflection. However, this solution is not optimal as the focus is axially displaced in the sample
during the lateral scanning. Moreover, the effective depth imaging range is reduced and different depths in the sample
experience different sensitivity, potentially influencing the interpretation of en-face views. An efficient solution is dark
field illumination as was recently shown by using Bessel beams to realize an extended focus (xf) OCT system operating
at 1300nm [2-4]. Dark field OCT imaging rejects back-reflected light, while back-scattered light remains unaffected. It is
also possible to measure the light back-reflected from the sample, in a second channel called bright field. It contains
addition information about the sample. The detection of both channels offers new measurement and contrast
perspectives. Furthermore simultaneous measurement permits natural co-registration and similar timing between
channels. In this paper, we focus on the use of that configuration for the particular case of a sample placed against a glass
slide perpendicular to the detection. This condition is optimal for imaging in the dark field channel for the previously
mentioned reasons. In the bright field channel, the reflection that occurs at the glass-tissue interface is measured with a
high SNR. A high SNR provides good phase stability, when the system operates under shot noise limited regime as
opposed to a condition where it would be limited by mechanical vibrations. There are several applications that require a
precise measurement of the phase. We propose to use this phase information for the optical measurement of vibration. It
has several applications, like tissue elastography or photoacoustic (PA) imaging for example. In particular, the
combination of PA imaging with OCT is interesting as they provide complementary information.
*[email protected]
Optical Coherence Tomography and Coherence Domain Optical Methods in Biomedicine XVI,
edited by Joseph A. Izatt, James G. Fujimoto, Valery V. Tuchin, Proc. of SPIE Vol. 8213, 82131M
© 2012 SPIE · CCC code: 1605-7422/12/$18 · doi: 10.1117/12.911443
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PA or ultrasound imaging, both require high bandwidth to provide good axial resolution. Typical values are in the tens of
MHz, depending on how much penetration depth is wished to be traded against resolution. This frequency range is still
above the fastest light sources available. It means that it is not possible to measure phase differences between A-Scans,
in a similar fashion like Doppler imaging. Another approach, based on the assumption that the signal shows periodic
behavior, which could be true for the PA signal over several pulses, consist in using an interlaced or gated approach.
However, it is usually not efficient to increase the bandwidth too much. It requires also assumption of the spectrum of
the signal. It was shown how a Fabry-Perrot filter in contact with the sample can be optically interrogated to measure
acoustic waves [5]. However, a non-contact measurement would be preferred. This was demonstrated by low-coherence
tomography (LCI) [6]. This method does not allow for depth resolved measurement. We introduce a new ultra-high
speed technique based on swept source (SS) technology and the intra-sweep analysis.
SS OCT imaging shows increasing interest over the Spectral Domain (SD) technique because of ultrahigh-speed, low
sensitivity roll-off and the absence of fringe wash-out. In case of SS, the interference signal is measured in time. In
particular, the phase difference over the k spectrum between reference and sample can be resolved over time, for
example by taking an Hilbert Transform of the interference fringes. It means that any time change of the path length
difference, for example caused by vibration, will affect the phase. A-Scan rates of current SS require electronic
bandwidth over 100MHz to over GHz, in term of optical detector, usually dual-balanced detectors and analog-to-digital
converters. This offers sufficient time resolution to resolve vibration in the MHz range. We are therefore analyzing the
phase within an A-Scan duration.
Fourier Domain Mode-Locking (FDML) laser is a particular interesting swept source as it provides high phase stability
due to pseudo stationary operation [7]. Furthermore, it usually operates with a 100% duty cycle that permits to avoid any
dead time for the phase analysis.
This work is a first demonstration of simultaneous OCT imaging and detection of high frequency vibration, ultrasound
(US) using an optical phase sensitive method.
2. MATERIAL AND METHODS
2.1 Optical Setup
The optical setup is shown in Fig. 1a. The light source is a FDML sweeping at 55kHz, centered at 1310nm with a
sweeping range of about 140nm, and followed by a buffering stage making delayed copies of the original sweep to
achieve up to 8 times speed multiplication, giving a 440kHz A-scan rate.
Fig. 1: a. Optical scheme. SS: Swept-source FDML followed by buffering stage, CIRC: circulator, FC: Fiber coupler, PC: polarization
controller, DC: dispersion compensation, REF: Reference arm, A: Axicon lens, M: Mirror, GALVO: Beam steering device, SAMPLE:
sample arm, DBD: Dual-balanced detector, CH A & B: input channels of the DAQ. US: ultrasound transducer. Red: Bessel detection.
b. Energy distribution in the objective back-focal plane and conjugated planes.
The simultaneous measurement in a second channel requires few changes to the extended focus system that we
previously described [2-4]. Dark field imaging requires different numerical aperture (NA) for illumination and detection
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(see Fig. 1b). As opposed to our previous system, we perform the illumination through a low NA standard Gaussian
mode. In this bright field channel, both back-reflected and back-scattered light will be detected. A circulator (CIRC)
followed by a 50/50 fiber coupler (FC) used for a Michelson interferometer is the configuration for measurement in
channel B. Only Back-scattered light will be detected by a Bessel mode, leading to the Dark field regime. A Bessel mode
results from a ring distribution in the objective back-focal plane. This can be created in the Fourier plane of a lens placed
after an Axicon (A). A relay is required to conjugate it to the fast axis of the steering system (GALVO) and to the
objective back-focal plane so that the phase relation can be preserved. At the first conjugated position, in the first 4f
system, illumination and detection are properly spatially separated. There, a mirror (M) smaller than the detection ring
permits to deflect only the illumination light and thus offers an energy efficient way to decouple the detection from the
illumination path. In channel A (red path), dark field operates with a Mach-Zehnder interferometer that consists in 2
50/50 fiber couplers is used to measure the scattered light through the axicon. A beam-splitter is used to split the
reference in the 2 channels after reflection on a common mirror. The path length difference of both channels can be
adjusted. The resulting interferences are measured with dual-balanced detectors (350 Mhz, PDB130C, Thorlabs). The
dual mode detection system is easily implemented with swept source OCT, since most of DAQs offer two independent
acquisition channels. The signal is digitalized with 8bit DAQ (1GS/s, ATS9870, Alazartech). The acquisition, as well as
simple processing, was performed with Labview 2010 (National Instruments).
In general the advantage of such combined system lies in the availability of a clear surface reflex offered by the Gaussian
mode together with structural imaging with optimal sensitivity and SNR offered by the dark field Bessel mode.
2.2 Processing
The signal processing, including k remapping follows what we previously described [2-4].
Fig. 2: Processing steps to detect on a glass slide interface producing a reflex, a vibration pulse occurring during A-Scan 30. a:
Interferogram of the M-Scan, b: Tomogram of the M-Scan, c: Filtered tomogram of the M-Scan, d: Phase of the filtered M-Scan, e:
FFT of the phase, f: Profile along the frequency bin corresponding to the delay of the reflex, g: Spectrogram of A-Scan 30, h:
Spectrogram of A-Scan 40.
The method that we propose for resolving the acoustic wave of several MHz on the surface at the reflex level is based on
high-speed SS OCT and FDML technology. Clearly the A-scan rate of 440kHz is too low to resolve the US pulse
oscillations. Also the time resolution of 1/440000 sec=2,2µs for the US wave corresponds to only 3.4mm axial resolution
assuming an average propagation speed in tissue of 1500m/s. However, the spectrum in SS OCT is acquired over time at
a speed of up to 1GHz. The pulse can then be extracted from the spectral phase obtained via a Hilbert transform method
together with a short time FFT (STFT) along the spectral phase over time. In the following, we analyze the phase for a
single spatial point over time, equivalent to an M-Scan.
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The signal at the glass slide interface acquired with the Gauss-Gauss mode (Bright field channel) is analyzed to measure
vibrations introduced by an ultrasound pulse. The use of OCT for vibration measurement permits to select the interface
of interest, while rejecting other influences. The processing steps are shown in Fig. 2. We measure M-Scans at a fixed
lateral position (Fig. 2a). We calculate the tomogram with the FFT of the remapped interferogram (Fig. 2b). The
tomogram shows the 2 reflections from the glass slide interfaces. We then filter one peak (Fig. 2c) and perform an
inverse FFT to calculate the spectral phase over time for each A-Scan (Fig. 2d) corresponding to the extracted signal
peak. At this point the oscillating signal associated to the vibration could already be accessed by removing the linear
phase ramp along k(t) due to the group or path length delay. We chose to calculate the FFT of the spectral phase signal
along k(t) for assessing the modulation frequency content (Fig. 2e). For a pulse with a limited bandwidth, it also
improves the sensitivity. We can then plot a profile along time or over the A-Scans for a frequency bin corresponding to
the delay of the reflex (Fig. 2f). This permits to locate the pulse but only with one A-Scan time accuracy. The central
step to improve the time accuracy is to calculate the short time Fourier transform (STFT) of the spectral phase (Fig. 2d)
of the A-Scan. The improvement result from the small window length.
3. RESULTS
The system was tested by using an ultrasound transducer pushing the sample against the glass slide. The ultrasound
transducer’s position is visible in Fig. 1a. It induces vibration on the glass slide after propagation through the tissue. The
repetition rate of the acoustic pulse was limited to 2kHz by the pulse driving the transducer. The A-Scan used was
110kA-Scan/s (period of 9µs) and the signal was digitalized at 250MSamples/s. The difference of the rate between
ultrasound and OCT permits to measure several A-Scans per pulse with only one containing the acoustic pulse. In our
case, it represents 54 A-Scans that produce 1 M-Scan.
3.1 OCT imaging
The lateral point spread functions (PSF) at the confocal position were measured for both channels by using iron oxide
nano-particles suspended in resin. In channel A, the central lobe of the Gauss-Bessel PSF is clearly visible together with
the first side lobe that is much lower in intensity (see Fig. 3 left). It has 6µm diameter that keeps constant lateral
extension over more than 1.5mm in depth. In channel B, the Gaussian spot has a size of about 15 µm (see Fig. 3 right).
Fig. 3: lateral PSFs at the confocal position measured with a nanoparticle. Left: Gauss-Bessel in channel A, right: Gauss-Gauss in
channel B. Scale bar denotes 10µm. The optics of the illumination path were selected to obtain a Bessel beam distribution central lobe
size (1/e2) of 5.5µm over a depth range of 1.6mm and a Gaussian spot of about 15µm.
As a test sample, we imaged chicken breast placed against the glass slide. Figure 4 shows tomograms taken with both
channels simultaneously. In the Gauss-Bessel mode or Dark field channel, the reflection of the glass plate fits into the
dynamic range and doesn´t degrade the signal structure (Fig. 4a). This is not case in the other (standard OCT) mode, the
bright field channel, where the sensitivity is lower (Fig. 4b). Indeed, the reference power has to be lowered so that the
signal from the reflection fits into the dynamic range. Furthermore the reflection increases also the noise floor. However
we can take advantage of the high SNR of the peak associated to the reflex to measure the phase.
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Fig. 4: Tomogram of chicken breast. a: Gauss-Bessel resolving the structure (Dark field in channel A), b: Gauss-Gauss with a strong
reflex from the glass plate (Bright field in channel B). Scale bare denotes 250µm. No averaging was applied.
3.2 Vibration measurement
The back side of the chicken breast was excited with a pulse of about 2µs. The thickness of the sample was about 5mm.
We first analyze the vibration over the M-Scan. After averaging over several pulses, a clear difference appears between
A-Scans. The pulse reached the glass slide during the 30th A-Scan.
Fig. 5: FFT of the phase over the M-Scan. It shows that a vibration occurred during the 30th.
It shows that the pulse can clearly be measured after propagation through the tissue. The time resolution is however
limited here by the A-Scan duration, in our case 9 µs. By calculating the STFT, the time accuracy can be improved. The
effect of the pulse introduces a small change in frequency. The time accuracy and the accuracy in general are then given
by the size of the time increment of the moving window and the size of the window. Differences of spectrograms can be
taken to highlight the difference. Figure 6 shows the result for the bin corresponding to the reflex delay for two
spectrograms differences, with and without US signature, (Spectrogram (S) for A-Scan 30 minus S40) and (S41-S40). It
clearly outlines the time evolution of the US pulse for the white line (S30-40).
Fig. 6: Difference of the STFT of the phase between the A-Scan 30 and A-Scans 40 in white and between A-Scan 41 and A-Scan 40 in
red plotted for the frequency bin corresponding to the reflex delay.
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4. DISCUSSION AND CONCLUSION
We present a new OCT imaging system that allows measuring simultaneously the reflected light and scattered light and
the scattered light only in 2 different channels, leading to bright and dark field imaging respectively. With a sample
placed against a glass slide perpendicular to the detection channel, dark field OCT imaging with extended focus (xfOCT) can be performed. The image quality is not impaired by the glass slide. The bright field channel offers a clear
surface reflex with high SNR. Because both channels are simultaneously acquired and are using the same illumination
beam, they are naturally co-registered and provide complementary and supplementary information about the sample. We
proposed to analyze the phase of the interface. Due to the high sensitivity of OCT and the large refractive step that
occurs at the interface between glass and tissue, a good SNR is obtained from the interface even if it is not placed at the
confocal position. Furthermore, as we can separately adjust the reference delay for both channels, the interface can be
placed close to the zero delay for optimal phase sensitivity.
The phase can be for example used to measure vibration or acoustic waves. This has further application for PA imaging.
We introduced a new method that allows for resolving high-frequency vibration, which is necessary to provide enough
acoustic axial resolution. It is based on SS technology by analyzing the phase within an A-Scan. Current electronics
bandwidth required for SS imaging are similar to the one required for PA imaging. It means that the measurement of PA
signal with OCT requires no modification to the measurement system. Furthermore, this optical phase sensitive method,
that provides a novel all optical detection scheme for the combination of OCT with PA imaging, allows for non-contact
measurement which is particularly interesting for specific applications.
A disadvantage of the proposed method is a low phase sensitivity at the border of the A-Scan because of smaller signal.
The effect of windowing has to be further investigated.
A high phase sensitive is generally required to provide enough sensitivity, in particular for in-vivo imaging. We believe
that OCT high-speed measurement is of paramount importance in order to be in a frequency domain where mechanical
vibrations are low. The current system provides enough phase stability for single spatial point measurement. The effect
of galvo scanners, required to perform 2 and 3D imaging will be investigated.
Acknowledgements
Funding from the European FP7 HEALTH program is acknowledged (grant no. 201880, FUN OCT). We thank Pu Zou
and Aart-Jan Verhoef from the Technical University Vienna for the kind help with the pulse laser as well as Wolfgang
Wieser, Christoph Eigenwillig, Benjamin Biedermann and Thomas Klein for the support with the FDML Laser.
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