This study was funded by Doctoral Scholarships from The University of Auckland (awarded to JD&#160;and MC), Sir Charles Hercus Health Research Fellowships (20/011 and 21/116) from the Health Research Council of New Zealand (awarded to J-CH to KT, respectively), a Doctoral Scholarship awarded by the National Heart Foundation (awarded to AA), Marsden Fast-Start grants (UOA1504 and UOA1703) from the Royal Society of New Zealand (awarded to J-CH and KT, respectively), and a James Cook Research Fellowship from the Royal Society of New Zealand (awarded to AT). The original development of this instrument was funded by a Marsden grant (11-UOA-199) from the Royal Society of New Zealand (awarded to AT and PN).

<ol>
	<li>Han, J. -. C., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Energetics+of+stress+production+in+isolated+cardiac+trabeculae+from+the+rat.">Energetics of stress production in isolated cardiac trabeculae from the rat.</a> American Journal of Physiology. Heart and circulatory physiology. 299 (5), 1382-1394 (2010).</li><li>Ter Keurs, H. E. D. J., Rijnsburger, W. H., Van Heuningen, R., Nagelsmit, M. J. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Tension+development+and+sarcomere+length+in+rat+cardiac+trabeculae.+Evidence+of+length-dependent+activation.">Tension development and sarcomere length in rat cardiac trabeculae. Evidence of length-dependent activation.</a> Circulation Research. 46 (5), 703-714 (1980).</li><li>Shen, X., Cannell, M. B., Ward, M. L. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Effect+of+SR+load+and+pH+regulatory+mechanisms+on+stretch-dependent+Ca2++entry+during+the+slow+force+response.">Effect of SR load and pH regulatory mechanisms on stretch-dependent Ca2+ entry during the slow force response.</a> Journal of Molecular and Cellular Cardiology. 63, 37-46 (2013).</li><li>Dowrick, J. M., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=The+slow+force+response+to+stretch:+Controversy+and+contradictions.">The slow force response to stretch: Controversy and contradictions.</a> Acta Physiologica. 226 (1), 13250 (2019).</li><li>Stuyvers, B. D. M. Y., Miura, M., Ter Keurs, H. E. D. J. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Diastolic+viscoelastic+properties+of+rat+cardiac+muscle;+involvement+of+Ca2+.">Diastolic viscoelastic properties of rat cardiac muscle; involvement of Ca2+.</a> Advances in Experimental Medicine and Biology. 430, 13-28 (1997).</li><li>Tang, E. J. L. P., Laven, R. C., Hajirassouliha, A., Nielsen, P. M. F., Taberner, A. J. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Measurement+of+displacement+in+isolated+heart+muscle+cells+using+markerless+subpixel+image+registration.">Measurement of displacement in isolated heart muscle cells using markerless subpixel image registration.</a> Conference Record - IEEE International Instrumentation and Measurement Technology Conference. , (2019).</li><li>Bers, D. M. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Cardiac+excitation-contraction+coupling.">Cardiac excitation-contraction coupling.</a> Nature. 415 (6868), 198-205 (2002).</li><li>Cheuk, M. L., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=A+method+for+markerless+tracking+of+the+strain+distribution+of+actively+contracting+cardiac+muscle+preparations.">A method for markerless tracking of the strain distribution of actively contracting cardiac muscle preparations.</a> Experimental Mechanics. 61 (1), 95-106 (2020).</li><li>Lippok, N., Coen, S., Nielsen, P., Vanholsbeeck, F. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Dispersion+compensation+in+Fourier+domain+optical+coherence+tomography+using+the+fractional+Fourier+transform.">Dispersion compensation in Fourier domain optical coherence tomography using the fractional Fourier transform.</a> Optics Express. 20 (21), 23398 (2012).</li><li>Cheuk, M. L., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Four-Dimensional+imaging+of+cardiac+trabeculae+contracting+in+vitro+using+gated+OCT.">Four-Dimensional imaging of cardiac trabeculae contracting in vitro using gated OCT.</a> IEEE Transactions on Biomedical Engineering. 64 (1), 218-224 (2017).</li><li>Ritland, H. N. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Relation+between+refractive+index+and+density+of+a+glass+at+constant+temperature.">Relation between refractive index and density of a glass at constant temperature.</a> Journal of the American Ceramic Society. 38 (2), 86-88 (1955).</li><li>Tuchina, D. K., Bashkatov, A. N., Genina, E. A., Tuchin, V. V. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Quantification+of+glucose+and+glycerol+diffusion+in+myocardium.">Quantification of glucose and glycerol diffusion in myocardium.</a> Journal of Innovative Optical Health Sciences. 8 (3), (2015).</li><li>Taberner, A., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=A+dynamometer+for+nature's+engines.">A dynamometer for nature's engines.</a> IEEE Instrumentation and Measurement Magazine. 22 (2), 10-16 (2019).</li><li>Arganda-Carreras, I., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Trainable+Weka+Segmentation:+A+machine+learning+tool+for+microscopy+pixel+classification.">Trainable Weka Segmentation: A machine learning tool for microscopy pixel classification.</a> Bioinformatics. 33 (15), 2424-2426 (2017).</li><li>Jiang, Y., Julian, F. 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The authors have nothing to disclose.

In this study, we present a configuration that enables the assembly of three optical systems combining brightfield, fluorescence, and OCT imaging to gather data from an actively contracting ex vivo cardiac trabecula (Figure 1 and Figure 2). Such an orchestrated integration is possible due to the design of the measurement chamber (Figure 3) to enable the orthogonal arrangement of the OCT to the brightfield-fluorescence axis. The muscle-mounting system plays an equally important role in the success of simultaneous quantifications of key indices in characterizing cardiac muscle excitation-contraction dynamics. Its novelty resides in enabling muscle-scanning procedures with no apparent disturbance to the mechanical performance of the muscle (Figure 6). With the combined imaging configuration and motorized-hook system for force measurement, this system can evaluate regional heterogeneity in the Ca2+ transient, displacement, and sarcomere length, together with macroscopic geometric information of a contracting trabecula throughout the time course of the twitch (Figures 7 and Figure 8).
Given the ubiquity of brightfield-epifluorescence imaging systems within cardiac research laboratories, the reproduction of these results can be achieved with some minor hardware considerations. Here, we present the image-processing toolkit for combining brightfield-epifluorescence and OCT, which is essential in analyzing the underlying contractile heterogeneity. The integration of the OCT requires an unobstructed optical path, while the gated imaging requires an external trigger line between the stimulus and both the OCT and brightfield imaging camera, and muscle mounting hooks that are capable of moving the sample throughout the measurement chamber. The required post-processing software and methods are freely available. In particular, the segmentation software used, WEKA14, is open-source. The technique of markerless tracking of material points8, sarcomere length, gated volumetric imaging10, and mesh generation codes are likewise accessible and can be made available on request from the corresponding author.
Muscle viability, optimal loading of Fura-2, and image focus are the three pillars that form the foundations of a successful experiment. Using a dissection solution containing BDM to prevent contracture, transportation of the muscle in a syringe, continuous oxygenation of the solution, and preparing new experimental solutions on the day of an experiment, all contribute to a high muscle viability rate. Before loading the trabecula with Fura-2AM, autofluorescence must be collected for each condition one is interested in studying as it can have a significant effect on the measured Ca2+ transient15. Oxygenation of Fura-2AM loading solution is complicated by the necessary inclusion of the surfactant pluronic-F127 to aid dye loading. To combat the resultant excess bubble formation caused by this surfactant, a small drop of anti-foam in the loading solution allows the user to increase the oxygenation rate, thereby improving the chance that the trabecula maintains functional viability throughout the loading process. Finally, the imaging focus must be uniform along muscle length to maximize the signal to noise ratio of the brightfield and fluorescence information.
There are two limitations to consider with the methods presented here. First is the spatial resolution of the fluorescence microscope. While the spatial resolutions of the OCT and brightfield imaging are high, the resolution of the fluorescence microscope is limited to the integral of the fluorescence from the volume captured within a 540 &#181;m by 540 &#181;m imaging window. There is scope to increase the spatial resolution of the fluorescence microscope by using a high-gain charge-coupled device camera, instead of a PMT, to capture the fluorescence signal at the expense of the signal to noise ratio16. Second is the diameter of the trabecula that can be studied in terms of measurable sarcomere length and geometric&#160;depth. The windowed-FFT approach for computing sarcomere length exploits the benefit of improved spatial resolution but is associated with reduced robustness (Figure 8D). In cases where turbid or large-diameter trabeculae are to be studied, the resolvability of the FFT will be greatly reduced due to the reduced contrast associated with the sarcomeric banding in larger tissue samples. Likewise, within the OCT, the back-reflections from an imaging depth of greater than 300 &#181;m will be too weak to be resolved during the segmentation stage. Hence, our technique is limited to trabeculae of a diameter less than 300&#160;&#181;m. However, it is not recommended to study large diameter samples as there may be problems with diffusive oxygenation of the muscle core during high rates of stimulation17.
Our method enables the assessment of ionic mechanical function in association with muscle geometry in healthy and diseased muscle, providing a powerful approach to understanding cardiac muscle physiology, pathophysiology, and pharmacology. The image-processing pipeline outlined here extracts data that will be pivotal for garnering a deeper understanding of contractile heterogeneity. One avenue to fully realize the potential of such a rich dataset is in the construction of mathematical models that integrate and interpret these data, and to make predictions that can be tested experimentally using our device.

Superfusion of isolated cardiac muscle tissue preparations is a standard and widely used protocol for studying cardiac ionic activation and mechanics1. In particular, the isolation of trabeculae, rod-like structures from the ventricular walls, has enabled assessment of phenomena including the length-dependent activation of contraction2, stretch-dependent response of contraction3,4, and diastolic viscoelasticity5 of cardiac tissue. Ter Keurs, the initiator of this technique of superfusing isolated trabeculae, initially used a combination of fluorescence imaging for Ca2+ measurements and laser diffraction for determining sarcomere lengths2,5. Since these early studies, it has become increasingly common to extract sarcomere length information with a greater spatial resolution using 2D fast Fourier transform (FFT)-based techniques6 on brightfield microscopy images. The two imaging systems allow a partial assessment of the underlying relationship between Ca2+ release and sarcomere length-dependent force production.
Cardiac muscle is striated, with the visible banding associated with an underlying series of contractile units consisting of thick and thick filaments. The interaction of these constituent filaments that make up sarcomeres underlies force generation, which commences as follows:&#160;a&#160;depolarizing electrical signal, or action potential, causes voltage-dependent L-type Ca2+ channels in the cell membrane to open;&#160;the ensuing cellular influx of Ca2+ induces the release of Ca2+ from the sarcoplasmic reticulum (SR), an intracellular Ca2+ store, in a process known as Ca2+-induced Ca2+ release7; this sudden increase in intracellular Ca2+ concentration from nanomolar to micromolar range enables force production to occur;&#160;Ca2+ pumps continuously extrude Ca2+ out of the cytosol back into the SR and extracellular compartment;&#160;when the intracellular Ca2+ concentration returns to the nanomolar range, force production ceases, and the muscle relaxes. During force production, the constituent thick and thin filaments slide over one another. The sarcomere length dictates the relative extent of overlap and, therefore, the potential for force production of the muscle macroscopically.
In this paper, we extend these fluorescence-brightfield imaging techniques to include optical coherence tomography (OCT). OCT utilizes the physical principle of interference and is capable of garnering the geometric deformation of tissue to understand muscle contractile heterogeneity8. Our device (Figure 1) uses a spectral-domain OCT (SD-OCT) system. In SD-OCT, a beam splitter splits the light from a broadband short coherence-length superluminescent diode into reference and measurement arms. The reference arm contains a fixed-mirror, and the measurement arm contains a 2D-galvanometer to steer the light. Light backscattered from the sample is collected and interferes with the reflected light in the reference arm to form an interference pattern. The depth information is encoded in the frequency of the spectral fringe. To extract the information, the signal is passed through a spectrometer and an inverse FFT is applied to the result. The corresponding 1D signal represents the structures at different depths, corresponding to changes in refractive index9 (A-scan). By steering the laser in a single axis, one can construct a cross-section of the sample of interest (B-scan) and, similarly, by repeating the process in a step-wise pattern in the remaining axis, a three-dimensional image can be generated (C-scan). By extension, one can collect a series of B-scans at a single slice for a repeating time-varying subject based on an external trigger and repeat to generate a three-dimensional scan, representing a time-varying planar image10.
In integrating the three imaging systems, we have considered the following two principles. First, imaging sensors should not detect light from an alternative imaging modality, and second, the physical design should contain free space for at least three simultaneous imaging planes. To address the first requirement, the brightfield microscope uses a 660 nm wavelength LED to illuminate the sample in an inverted configuration. The fluorescence microscope is in an epifluorescence configuration where the same objective lens is used for both excitation and collection of the emitted light. The excitation light has a wavelength of between 340 nm and 380 nm, and a photomultiplier tube (PMT) measures the emitted light at a wavelength of 510 nm. A pair of dichroic mirrors enable these two optical paths to share the same physical footprint without interfering with the opposite measurement (Figure 2). Finally, the OCT uses broadband (100 nm spectral width) light with a central wavelength of 840&#160;nm, distinct from the other two modalities. Due to the low-coherence nature of the light used for OCT, any scattered light from the brightfield-fluorescence sources will not contribute to the interference pattern that encodes depth information. For the second requirement, the housing design for the capillary tube has accessible optical pathways to the anterior, inferior, and superior planes of the sample. During experiments, two platinum hooks hold a trabecula within a capillary tube perfused with oxygenated Krebs-Henseleit (KH) solution. The galvanometer head of the OCT is oriented orthogonally to the brightfield-fluorescence imaging pathway to take advantage of the third orthogonal optical plane (Figure 3).
This paper outlines the design considerations for building a device capable of simultaneously imaging calcium, sarcomere length, and muscle geometry. To demonstrate these measurement capabilities, we describe the process of isolating a ventricular trabecula, the preparation of the necessary buffer solutions, along with the critical steps involved in the handling and fluorescence loading of an ex vivo trabecula. Finally, this paper outlines the processes required to translate the dataset into more useful visualizations.

<table border="1" fo:keep-together.within-page="1" fo:keep-with-next.within-page="always">
	<tbody>
		<tr>
			<td>2,3-Butanedione monoxime</td>
			<td>Acros Organics</td>
			<td>150375000</td>
			<td></td>
		</tr>
		<tr>
			<td>20&#215; microscope lens</td>
			<td>Nikon</td>
			<td>CFI Super Fluor 20&#215;</td>
			<td>NA 0.75</td>
		</tr>
		<tr>
			<td>2D Galvanometer</td>
			<td>Thorlabs</td>
			<td>GVSM002/M</td>
			<td></td>
		</tr>
		<tr>
			<td>50-50 beam splitter</td>
			<td>Thorlabs</td>
			<td>FC850-40-50-APC</td>
			<td></td>
		</tr>
		<tr>
			<td>90-10 beam-splitter</td>
			<td>Thorlabs</td>
			<td>TW850R2A2</td>
			<td></td>
		</tr>
		<tr>
			<td>Analogue input module</td>
			<td>National Instruments</td>
			<td>NI-9205</td>
			<td>Records the PMT signal at 200 kHz</td>
		</tr>
		<tr>
			<td>Brightfield imaging light source</td>
			<td>CoolLED</td>
			<td>PE-2</td>
			<td>660&#160;nm LAM</td>
		</tr>
		<tr>
			<td>Broadband light source</td>
			<td>Superlum</td>
			<td>Broadlighter-840</td>
			<td></td>
		</tr>
		<tr>
			<td>CaCl2</td>
			<td>Sigma-Aldrich</td>
			<td>C4901</td>
			<td></td>
		</tr>
		<tr>
			<td>Cameralink card</td>
			<td>National Instruments</td>
			<td>NI-1429</td>
			<td>Brightfield imaging frame grabber</td>
		</tr>
		<tr>
			<td>Carbogen 5</td>
			<td>BOC</td>
			<td>Gas code: 181</td>
			<td></td>
		</tr>
		<tr>
			<td>Condensor lens</td>
			<td>Nikon</td>
			<td>LWD 0.52</td>
			<td></td>
		</tr>
		<tr>
			<td>D(+)-Glucose</td>
			<td>Merck</td>
			<td>108337</td>
			<td></td>
		</tr>
		<tr>
			<td>DAQ</td>
			<td>National Instruments</td>
			<td>NI-6259</td>
			<td>Triggers the galvanometer movement</td>
		</tr>
		<tr>
			<td>Dichroic mirror 1</td>
			<td>Semrock</td>
			<td>FF409-Di03</td>
			<td></td>
		</tr>
		<tr>
			<td>Dichroic mirror 2</td>
			<td>Semrock</td>
			<td>FF552-Di02</td>
			<td></td>
		</tr>
		<tr>
			<td>Diffraction grating</td>
			<td>Wasatch Photonics</td>
			<td>1200 lines/mm @840 nm</td>
			<td></td>
		</tr>
		<tr>
			<td>Dimethyl sulfoxide</td>
			<td>Sigma-Aldrich</td>
			<td>276855</td>
			<td></td>
		</tr>
		<tr>
			<td>Direct-Q 3 UV System</td>
			<td>Merck Millipore</td>
			<td>ZRQSVR3WW</td>
			<td>Distilled water machine</td>
		</tr>
		<tr>
			<td>Dry bath</td>
			<td>Corning</td>
			<td>6875-SB</td>
			<td>LSE digital dry bath</td>
		</tr>
		<tr>
			<td>FIJI</td>
			<td>ImageJ</td>
			<td></td>
			<td>Open-source image processing software</td>
		</tr>
		<tr>
			<td>Fura-2AM pentapotassium salt</td>
			<td>Thermofisher</td>
			<td>F14186</td>
			<td></td>
		</tr>
		<tr>
			<td>Hardware FPGA card</td>
			<td>National Instruments</td>
			<td>NI-7813R</td>
			<td>Also controls the triggering of the brightfield capture</td>
		</tr>
		<tr>
			<td>Heparin</td>
			<td>Pfizer</td>
			<td>61024</td>
			<td></td>
		</tr>
		<tr>
			<td>HEPES</td>
			<td>PanReac AppliChem</td>
			<td>A1069</td>
			<td></td>
		</tr>
		<tr>
			<td>Inverted microscope</td>
			<td>Nikon</td>
			<td>TI-DH illumination pillar</td>
			<td></td>
		</tr>
		<tr>
			<td>Isofluorane</td>
			<td>MedSource</td>
			<td>VAPDRUGISO250</td>
			<td></td>
		</tr>
		<tr>
			<td>KCl</td>
			<td>Sigma-Aldrich</td>
			<td>P9541</td>
			<td></td>
		</tr>
		<tr>
			<td>KH2PO4</td>
			<td>Sigma-Aldrich</td>
			<td>P5655</td>
			<td></td>
		</tr>
		<tr>
			<td>Line-scan camera</td>
			<td>Basler</td>
			<td>spL2048-70km</td>
			<td>Spectrometer camera</td>
		</tr>
		<tr>
			<td>Magnetic stirrer</td>
			<td>IKA</td>
			<td>3810000</td>
			<td>RCT basic</td>
		</tr>
		<tr>
			<td>Matlab</td>
			<td>Mathworks</td>
			<td></td>
			<td>Data processing code</td>
		</tr>
		<tr>
			<td>MgCl2</td>
			<td>Sigma-Aldrich</td>
			<td>M8266</td>
			<td></td>
		</tr>
		<tr>
			<td>MgSO4.7H2O</td>
			<td>Sigma-Aldrich</td>
			<td>M1880</td>
			<td></td>
		</tr>
		<tr>
			<td>NaCl</td>
			<td>Sigma-Aldrich</td>
			<td>71376</td>
			<td></td>
		</tr>
		<tr>
			<td>NaH2PO4.2H2O</td>
			<td>Sigma-Aldrich</td>
			<td>71505</td>
			<td></td>
		</tr>
		<tr>
			<td>NaHCO3</td>
			<td>Sigma-Aldrich</td>
			<td>S6014</td>
			<td></td>
		</tr>
		<tr>
			<td>OCT FPGA card</td>
			<td>National Instruments</td>
			<td>NI-1483R</td>
			<td></td>
		</tr>
		<tr>
			<td>Oxygen tank</td>
			<td>BOC</td>
			<td>Gas code: 100D</td>
			<td></td>
		</tr>
		<tr>
			<td>pH meter</td>
			<td>Mettler Toledo</td>
			<td>MP220</td>
			<td></td>
		</tr>
		<tr>
			<td>Photomultiplier tube</td>
			<td>Hamamatsu</td>
			<td>H7422-20</td>
			<td></td>
		</tr>
		<tr>
			<td>Powerload</td>
			<td>Thermofisher</td>
			<td>P10020</td>
			<td></td>
		</tr>
		<tr>
			<td>Superluminescent diode</td>
			<td>Broadlighter</td>
			<td>D-840</td>
			<td></td>
		</tr>
		<tr>
			<td>Transimpedance amplifier</td>
			<td>Custom</td>
			<td></td>
			<td></td>
		</tr>
		<tr>
			<td>Tris(hydroxymethyl)amino-methane</td>
			<td>Sigma-Aldrich</td>
			<td>252859</td>
			<td></td>
		</tr>
		<tr>
			<td>Wistar rat</td>
			<td>Vernon Jansen Unit</td>
			<td></td>
			<td>8 &#8211; 10 weeks</td>
		</tr>
		<tr>
			<td>Xenon arc lamp</td>
			<td>Sutter Instrument</td>
			<td>DG-4</td>
			<td>Lambda DG-4</td>
		</tr>
	</tbody>
</table>

simultaneous brightfield, fluorescence, and optical coherence tomographic imaging of contracting cardiac trabeculae ex vivo

In cardiac muscle, intracellular Ca2+ transients activate contractile myofilaments, causing contraction, macroscopic shortening, and geometric&#160;deformation. Our understanding of the internal relationships between these events has been limited because we can neither &#39;see&#39; inside the muscle nor precisely track the spatio-temporal nature of excitation-contraction dynamics. To resolve these problems, we have constructed a device that combines a suite of imaging modalities. Specifically, it integrates a brightfield microscope to measure local changes of sarcomere length and tissue strain, a fluorescence microscope to visualize the Ca2+ transient, and an optical coherence tomograph to capture the tissue&#39;s geometric&#160;changes throughout the time-course of a cardiac cycle. We present here the imaging infrastructure and associated data collection framework. Data are collected from isolated rod-like tissue structures known as trabeculae carneae. In our instrument, a pair of position-controlled platinum hooks hold each end of an ex vivo muscle sample while it is continuously superfused with nutrient-rich saline solution. The hooks are under independent control, permitting real-time control of muscle length and force. Lengthwise translation enables the piecewise scanning of the sample, overcoming limitations associated with the relative size of the microscope imaging window (540 &#181;m by 540 &#181;m) and the length of a typical trabecula (&#62;2000 &#181;m). Platinum electrodes at either end of the muscle chamber stimulate the trabecula at a user-defined rate. We exploit the stimulation signal as a trigger for synchronizing the data from each imaging window to reconstruct the entire sample twitching under steady-state conditions. Applying image-processing techniques to these brightfield imaging data provides tissue displacement and sarcomere length maps. Such a collection of data, when incorporated into an experiment-modeling pipeline, will provide a deeper understanding of muscle contractile homogeneity and heterogeneity in physiology and pathophysiology.

In order to capture the regional Ca2+ and brightfield information for the entire length of the trabecula presented here, seven muscle positions were required. Figure 6 suggests that the twitch force was undisturbed by this motion, revealing that there was no position dependence of the active force production.
B-scans collected using optical coherence tomography at a rate of 100 Hz were segmented using the ImageJ plugin WEKA14 (Figure 7A). Each cross-section appears distorted due to the difference between the lateral (10 &#181;m) and depth (1.73 &#181;m (in myocardium)) resolutions. This distortion was corrected by scaling the image&#39;s depth axis by the lateral resolution-depth resolution ratio. Figure 7B,C demonstrate that after scaling the raw C-scan of the trabecula it is approximately cylindrical in geometry. The reflection of the measurement chamber wall can sometimes overlap with the muscle data (Figure 7A,B), but the segmentation software can be trained to account for this (Figure 7D,E). Once segmented, the cross-sectional area along the length of the muscle can be calculated throughout the twitch (Figure 7F). Note that this particular trabecula has a small appendage branching from it. The motion of the branch is evident ~0.75&#160;mm along the trabecula. Finally, the segmented images can be converted into meshes to aid the construction of geometric&#160;models (Figure 7G).
Imaging data captured at each of the different trabecula positions at a rate of 100 fps were stitched together to create a single complete image of the trabecula (Figure 8A). The resolution of these images is 0.535 &#181;m/pixel. The use of linear weighting functions in the overlapping regions of neighboring windows aids visualization and minimizes the impact of the vignette present in the brightfield images. To measure the fluorescent signal, a 540 &#181;m by 540 &#181;m window of the trabecula is cyclically illuminated with 340 nm, 365 nm, and 380 nm wavelength light at a rate of 600 Hz. The ratio of the emitted fluorescence associated with the 340 nm and 380 nm excitation light correlates to the intracellular calcium after the trabecula has been loaded with Fura-2. As this measurement is a ratio, the effective measurement rate is 200 Hz. The averaged (n = 10) intracellular Ca2+ transients from each window are aligned with the region they were imaged (Figure 8B). While the peak of the transients appear reasonably consistent, the diastolic [Ca2+]&#160;is lower within the region between 900 &#181;m and 1800 &#181;m along the trabecula. Similarly, the results of the displacement tracking (Figure 8C) and sarcomere length (Figure 8D) calculations also indicate the presence of regional variability. The markerless tracking technique used is able to process the displacement of each pixel, given enough contrast. When mapping the distribution of sarcomere lengths in post, a cross-correlation area of 128 pixels by 128 pixels (~67 &#181;m by 67 &#181;m) was used to calculate regional sarcomere length. This area encapsulates approximately 29 sarcomeres when close the sample is close to optimal sarcomere length. The step-size (in both the x- and y-directions) between the centroid of each cross-correlation window was set to 50 pixels (~26 &#181;m) for processing these data. The suitability of sarcomere length estimates were tested based on the width and amplitude of the Gaussian fit to the FFT signal. These conditions were not met in the muscle region between 0 &#181;m and 500 &#181;m so no sarcomere length information could be computed there. Given the associated displacements, it is likely that the sarcomeres in this region elongated during the contractile phase. In keeping with this speculation, the average sarcomere lengths on the right-hand side of the trabecula shorten during that period. By combining the information provided by each of the panels, it appears that the region with the largest cross-sectional area does not produce the most significant force. On the assumption that the regional variation of Ca2+ transients has an approximately smooth gradient, Figure 8B indicates that the largest amplitude Ca2+ transient occurs somewhere between 1300 &#181;m and 1600 &#181;m along the trabecula. The displacement map indicates that the region that undergoes the least motion aligns well with the peak Ca2+ transient. However, this region has the smallest cross-sectional areas of the sample. With these data in mind, one could infer that this region generated the most stress.
<img alt="Figure 1" class="xfigimg" src="/files/ftp_upload/62799/62799fig01.jpg" /> 
Figure 1: Annotated image of the cardiomyometer. Each of the main optical components is outlined. The inset contains a close-up, rear-view, of the microscope objective in situ, below the measurement chamber. <a href="https://www.jove.com/files/ftp_upload/62799/62799fig01large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 2" class="xfigimg" src="/files/ftp_upload/62799/62799fig02.jpg" /> 
Figure 2: Optical pathway for simultaneous brightfield and fluorescence microscopy. The illumination source for the fluorescence microscope is a Xenon arc lamp, the output of which cyclically switches between 340 nm, 365 nm, and 380&#160;nm light. The arc lamp output path contains a dichroic mirror with a cut-off wavelength of 409 nm that reflects the UV light onto a mirror that directs the light into a fluorescence microscope objective. The lens focuses the excitation light onto the sample and collects the emitted light, which has a longer wavelength of 510 nm. This emitted light passes through the first dichroic mirror, but not the second, as it has a cut-off wavelength of 552 nm. A field lens then focuses the reflected light onto the sensor of the PMT. Meanwhile, the illumination source (660 nm LED) for the brightfield microscope is located above the sample. The transmitted light is focused onto the sample by a condenser lens, and the 20&#215; fluorescence objective captures the resultant transmission image. The wavelength used for brightfield illumination exceeds the cut-off wavelength of each dichroic mirror, so it passes through both of them before the image is focused onto the sensor of a CMOS camera. <a href="https://www.jove.com/files/ftp_upload/62799/62799fig02large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 3" class="xfigimg" src="/files/ftp_upload/62799/62799fig03.jpg" /> 
Figure 3: Measurement chamber holder design. (A)&#160;Isometric view of the measurement chamber holder with optical paths overlaid. Brightfield illumination occurs from the superior surface (z-axis); fluorescence illumination occurs from the inferior surface (z-axis), and the measurement arm OCT signal is orthogonal to the other illumination axis (y-axis). During the experiment, two platinum hooks hold a trabecula within a glass capillary tube that functions as the measurement chamber. Voice-coil motors control each hook and their positions are measured using laser interferometry. The current position is compared with a user-defined set point and, using a PID controller encoded within an FPGA, the error is minimized. (B)&#160;Measurement chamber in situ with the brightfield illumination turned on. The back-view is shown in the Figure 1 inset. (C)&#160;Schematic of the superfusate flow through the measurement chamber holder. Superfusate enters the rear of the block and flows in the direction indicated by arrows. The upstream and downstream electrodes establish the field stimulation for eliciting the contraction of a mounted trabecula in the measurement chamber. Blue shading indicates the regions where the superfusate flows during an experiment. <a href="https://www.jove.com/files/ftp_upload/62799/62799fig03large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 4" class="xfigimg" src="/files/ftp_upload/62799/62799fig04.jpg" /> 
Figure 4: Front panel of the image acquisition and control software. (A) Brightfield imaging user interface. (B) OCT imaging user interface. (C)&#160;Hardware control user interface. <a href="https://www.jove.com/files/ftp_upload/62799/62799fig04large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 5" class="xfigimg" src="/files/ftp_upload/62799/62799fig05.jpg" /> 
Figure 5: Trabecula dissection and mounting protocol. (A)&#160;Langendorff-perfused rat heart in the dissection chamber. (B) The same heart with the atria removed. Dashed lines indicate the excision trajectory to open the ventricles. (C)&#160;An opened heart to expose the interior of both ventricles. The dashed box indicates the region where trabeculae are typically found. (D)&#160;Excised right-ventricular wall region (the same as indicated by the dashed box in C). Dashed lines highlight three trabeculae. (E)&#160;A trabecula selected from the three in D. (F)&#160;The trabecula from panel E with the wall tissue removed. (G)&#160;The isolated trabecula in the end of a 1 mL syringe. (H)&#160;The trabecula in the mounting chamber. (I)&#160;The trabecula mounted between two platinum hooks. (J)&#160;The trabecula, mounted between hooks, within the measurement chamber (Figure 3B). The green spot is an artefact from the first dichroic filter. (K) Secondary angle of the mounted trabecula within the measurement chamber. The distance between the trabecula and the microscope objective lens is approximately 1 mm. <a href="https://www.jove.com/files/ftp_upload/62799/62799fig05large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 6" class="xfigimg" src="/files/ftp_upload/62799/62799fig06.jpg" /> 
Figure 6: Position dependence of the force measurement. The force produced by the muscle from each of the imaging positions (n = 7) overlaid. The average active force production was 0.527 mN &#177; 0.003 mN, time to 50 % contraction 77.1 ms &#177; 0.3 ms, and time to 50 % relaxation 328.1 ms &#177; 0.9 ms (all data are presented as the mean &#177; SE). <a href="https://www.jove.com/files/ftp_upload/62799/62799fig06large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 7" class="xfigimg" src="/files/ftp_upload/62799/62799fig7v2.jpg" /> 
Figure 7: OCT imaging analysis. (A)&#160;Example of WEKA segmentation. The segmented cross-section of the muscle is highlighted in red, and the background is highlighted in green. (B)&#160;Superior view of the raw C-scan data of a trabecula. The bright angled line towards the top of the image is the reflection of the measurement chamber wall. (C)&#160;Lateral view of the raw C-scan data of a trabecula. (D) Superior view of the segmented OCT data. (E) Lateral view of the segmented OCT data. (F)&#160;The cross-sectional area along the length of the trabecula (x-axis) through time (y-axis). The average cross-sectional area along the muscle length was 0.0326 mm2 &#177; 0.0005 mm2 (mean &#177; S.E.) (G)&#160;A mesh of the trabecula. The mesh has been approximately aligned with the cross-sectional area plot of panel F. <a href="https://www.jove.com/files/ftp_upload/62799/62799fig7v2lage.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 8" class="xfigimg" src="/files/ftp_upload/62799/62799fig08.jpg" /> 
Figure 8: Brightfield and fluorescence imaging analysis. (A)&#160;Stitched image (seven imaging windows) of the trabecula. (B)&#160;Ca2+&#160;transients along the length of the trabecula. (C)&#160;Average x-displacement from each imaging window. A positive displacement represents a motion to the right and negative to the left. (D)&#160;Average sarcomere lengths from each imaging window that had the necessary image contrast. <a href="https://www.jove.com/files/ftp_upload/62799/62799fig08large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
Table 1: Table of Solutions <a href="https://www.jove.com/files/ftp_upload/62799/Tables_of_solutions.xlsx" target="_blank">Please click here to download this Table.</a>

Watch this Scientific Journal Video about Simultaneous Brightfield, Fluorescence, and Optical Coherence Tomographic Imaging of Contracting Cardiac Trabeculae Ex Vivo at JoVE.com

Simultaneous Brightfield, Fluorescence, and Optical Coherence Tomographic Imaging of Contracting Cardiac Trabeculae Ex Vivo

This protocol presents a collection of sarcomere, calcium, and macroscopic geometric&#160;data from an actively contracting cardiac trabecula ex vivo. These simultaneous measurements are made possible by the integration of three imaging modalities.

Simultaneous Brightfield, Fluorescence, and Optical Coherence Tomographic Imaging of Contracting Cardiac Trabeculae Ex Vivo

In cardiac muscle, intracellular Ca2+ transients activate contractile myofilaments, causing contraction, macroscopic shortening, and ...

simultaneous-brightfield-fluorescence-optical-coherence-tomographic

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JoVE Journal

Bioengineering

3.2K Views.  University of Auckland. This protocol presents a collection of sarcomere, calcium, and macroscopic geometric data from an actively contracting cardiac trabecula ex vivo. These simultaneous measurements are made possible by the integration of three imaging modalities.

Video: Simultaneous Brightfield, Fluorescence, and Optical Coherence Tomographic Imaging of Contracting Cardiac Trabeculae Ex Vivo

High Resolution 3D Imaging of Ex-Vivo Biological Samples by Micro CT

Non-destructive volume visualization can be achieved only by tomographic techniques, of which the most efficient is the x-ray micro computerized tomography (&mu;CT).
High resolution &mu;CT is a very versatile yet accurate (1-2 microns of resolution) technique for 3D examination of ex-vivo biological samples1, 2. As opposed to electron tomography, the &mu;CT allows the examination of up to 4 cm thick samples. This technique requires only few hours of measurement as compared to weeks in histology. In addition, &mu;CT does not rely on 2D stereologic models, thus it may complement and in some cases can even replace histological methods3, 4, which are both time consuming and destructive. Sample conditioning and positioning in &mu;CT is straightforward and does not require high vacuum or low temperatures, which may adversely affect the structure. The sample is positioned and rotated 180&deg; or 360&deg;between a microfocused x-ray source and a detector, which includes a scintillator and an accurate CCD camera, For each angle a 2D image is taken, and then the entire volume is reconstructed using one of the different available algorithms5-7. The 3D resolution increases with the decrease of the rotation step. The present video protocol shows the main steps in preparation, immobilization and positioning of the sample followed by imaging at high resolution.

High Resolution 3D Imaging of Ex-Vivo Biological Samples by Micro CT

Non-destructive volume visualization can be achieved only by tomographic techniques, of which the most efficient is the x-ray micro computerized tomography ( CT).

Non-destructive volume visualization can be achieved only by tomographic techniques, of which the most efficient is the x-ray micro computerized ...

Three-dimensional Optical-resolution Photoacoustic Microscopy

Optical microscopy, providing valuable insights at the cellular and organelle levels, has been widely recognized as an enabling biomedical technology. As the mainstays of in vivo three-dimensional (3-D) optical microscopy, single-/multi-photon fluorescence microscopy and optical coherence tomography (OCT) have demonstrated their extraordinary sensitivities to fluorescence and optical scattering contrasts, respectively. However, the optical absorption contrast of biological tissues, which encodes essential physiological/pathological information, has not yet been assessable. 
The emergence of biomedical photoacoustics has led to a new branch of optical microscopy optical-resolution photoacoustic microscopy (OR-PAM)1, where the optical irradiation is focused to the diffraction limit to achieve cellular1 or even subcellular2 level lateral resolution. As a valuable complement to existing optical microscopy technologies, OR-PAM brings in at least two novelties. First and most importantly, OR-PAM detects optical absorption contrasts with extraordinary sensitivity (i.e., 100%). Combining OR-PAM with fluorescence microscopy3 or with optical-scattering-based OCT4 (or with both) provides comprehensive optical properties of biological tissues. Second, OR-PAM encodes optical absorption into acoustic waves, in contrast to the pure optical processes in fluorescence microscopy and OCT, and provides background-free detection. The acoustic detection in OR-PAM mitigates the impacts of optical scattering on signal degradation and naturally eliminates possible interferences (i.e., crosstalks) between excitation and detection, which is a common problem in fluorescence microscopy due to the overlap between the excitation and fluorescence spectra. 
Unique for optical absorption imaging, OR-PAM has demonstrated broad biomedical applications since its invention, including, but not limited to, neurology5, 6, ophthalmology7, 8, vascular biology9, and dermatology10. In this video, we teach the system configuration and alignment of OR-PAM as well as the experimental procedures for in vivo functional microvascular imaging.

Optical-resolution photoacoustic microscopy (OR-PAM) is an emerging technology capable of imaging optical absorption contrasts in vivo with cellular resolution and sensitivity. Here, we provide a visualized instruction on the experimental protocols of OR-PAM, including system configuration, system alignment, typical in vivo experimental procedures, and functional imaging schemes.

Optical microscopy, providing valuable insights at the cellular and organelle levels, has been widely recognized as an enabling biomedical technology. As ...

Optical Frequency Domain Imaging of Ex vivo Pulmonary Resection Specimens: Obtaining One to One Image to Histopathology Correlation

Lung cancer is the leading cause of cancer-related deaths1. Squamous cell and small cell cancers typically arise in association with the conducting airways, whereas adenocarcinomas are typically more peripheral in location. Lung malignancy detection early in the disease process may be difficult due to several limitations: radiological resolution, bronchoscopic limitations in evaluating tissue underlying the airway mucosa and identifying early pathologic changes, and small sample size and/or incomplete sampling in histology biopsies. High resolution imaging modalities, such as optical frequency domain imaging (OFDI), provide non-destructive, large area 3-dimensional views of tissue microstructure to depths approaching 2 mm in real time (Figure 1)2-6. OFDI has been utilized in a variety of applications, including evaluation of coronary artery atherosclerosis6,7 and esophageal intestinal metaplasia and dysplasia6,8-10.
Bronchoscopic OCT/OFDI has been demonstrated as a safe in vivo imaging tool for evaluating the pulmonary airways11-23 (Animation). OCT has been assessed in pulmonary airways16,23 and parenchyma17,22 of animal models and in vivo human airway14,15. OCT imaging of normal airway has demonstrated visualization of airway layering and alveolar attachments, and evaluation of dysplastic lesions has been found useful in distinguishing grades of dysplasia in the bronchial mucosa11,12,20,21. OFDI imaging of bronchial mucosa has been demonstrated in a short bronchial segment (0.8 cm)18. Additionally, volumetric OFDI spanning multiple airway generations in swine and human pulmonary airways in vivo has been described19. Endobronchial OCT/OFDI is typically performed using thin, flexible catheters, which are compatible with standard bronchoscopic access ports. Additionally, OCT and OFDI needle-based probes have recently been developed, which may be used to image regions of the lung beyond the airway wall or pleural surface17.
While OCT/OFDI has been utilized and demonstrated as feasible for in vivo pulmonary imaging, no studies with precisely matched one-to-one OFDI:histology have been performed. Therefore, specific imaging criteria for various pulmonary pathologies have yet to be developed. Histopathological counterparts obtained in vivo consist of only small biopsy fragments, which are difficult to correlate with large OFDI datasets. Additionally, they do not provide the comprehensive histology needed for registration with large volume OFDI. As a result, specific imaging features of pulmonary pathology cannot be developed in the in vivo setting. Precisely matched, one-to-one OFDI and histology correlation is vital to accurately evaluate features seen in OFDI against histology as a gold standard in order to derive specific image interpretation criteria for pulmonary neoplasms and other pulmonary pathologies. Once specific imaging criteria have been developed and validated ex vivo with matched one-to-one histology, the criteria may then be applied to in vivo imaging studies. Here, we present a method for precise, one to one correlation between high resolution optical imaging and histology in ex vivo lung resection specimens. Throughout this manuscript, we describe the techniques used to match OFDI images to histology. However, this method is not specific to OFDI and can be used to obtain histology-registered images for any optical imaging technique. We performed airway centered OFDI with a specialized custom built bronchoscopic 2.4 French (0.8 mm diameter) catheter. Tissue samples were marked with tissue dye, visible in both OFDI and histology. Careful orientation procedures were used to precisely correlate imaging and histological sampling locations. The techniques outlined in this manuscript were used to conduct the first demonstration of volumetric OFDI with precise correlation to tissue-based diagnosis for evaluating pulmonary pathology24. This straightforward, effective technique may be extended to other tissue types to provide precise imaging to histology correlation needed to determine fine imaging features of both normal and diseased tissues.

Optical Frequency Domain Imaging of Ex vivo Pulmonary Resection Specimens: Obtaining One to One Image to Histopathology Correlation

A method to image ex vivo pulmonary resection specimens with optical frequency domain imaging (OFDI) and obtain precise correlation to histology is described, which is essential to developing specific OFDI interpretation criteria for pulmonary pathology. This method is applicable to other tissue types and imaging techniques to obtain precise imaging to histology correlation for accurate image interpretation and assessment. Imaging criteria established with this technique would then be applicable to image assessment in future in vivo studies.

Lung cancer is the leading cause of cancer-related deaths1. Squamous cell and small cell cancers typically arise in association with the conducting ...

Integrated Photoacoustic Ophthalmoscopy and Spectral-domain Optical Coherence Tomography

Both the clinical diagnosis and fundamental investigation of major ocular diseases greatly benefit from various non-invasive ophthalmic imaging technologies. Existing retinal imaging modalities, such as fundus photography1, confocal scanning laser ophthalmoscopy (cSLO)2, and optical coherence tomography (OCT)3, have significant contributions in monitoring disease onsets and progressions, and developing new therapeutic strategies. However, they predominantly rely on the back-reflected photons from the retina. As a consequence, the optical absorption properties of the retina, which are usually strongly associated with retinal pathophysiology status, are inaccessible by the traditional imaging technologies.
Photoacoustic ophthalmoscopy (PAOM) is an emerging retinal imaging modality that permits the detection of the optical absorption contrasts in the eye with a high sensitivity4-7 . In PAOM nanosecond laser pulses are delivered through the pupil and scanned across the posterior eye to induce photoacoustic (PA) signals, which are detected by an unfocused ultrasonic transducer attached to the eyelid. Because of the strong optical absorption of hemoglobin and melanin, PAOM is capable of non-invasively imaging the retinal and choroidal vasculatures, and the retinal pigment epithelium (RPE) melanin at high contrasts 6,7. More importantly, based on the well-developed spectroscopic photoacoustic imaging5,8 , PAOM has the potential to map the hemoglobin oxygen saturation in retinal vessels, which can be critical in studying the physiology and pathology of several blinding diseases 9 such as diabetic retinopathy and neovascular age-related macular degeneration.
Moreover, being the only existing optical-absorption-based ophthalmic imaging modality, PAOM can be integrated with well-established clinical ophthalmic imaging techniques to achieve more comprehensive anatomic and functional evaluations of the eye based on multiple optical contrasts6,10 . In this work, we integrate PAOM and spectral-domain OCT (SD-OCT) for simultaneously in vivo retinal imaging of rat, where both optical absorption and scattering properties of the retina are revealed. The system configuration, system alignment and imaging acquisition are presented.

Photoacoustic ophthalmology (PAOM), an optical-absorption-based imaging modality, provides the complementary evaluation of the retina to the currently available ophthalmic imaging technologies. We report the using of PAOM integrated with spectral-domain optical coherence tomography (SD-OCT) for simultaneous multimodal retinal imaging in rats.

Both the clinical diagnosis and fundamental investigation of major ocular diseases greatly benefit from various non-invasive ophthalmic imaging ...

Wideband Optical Detector of Ultrasound for Medical Imaging Applications

Optical sensors of ultrasound are a promising alternative to piezoelectric techniques, as has been recently demonstrated in the field of optoacoustic imaging. In medical applications, one of the major limitations of optical sensing technology is its susceptibility to environmental conditions, e.g. changes in pressure and temperature, which may saturate the detection. Additionally, the clinical environment often imposes stringent limits on the size and robustness of the sensor. In this work, the combination of pulse interferometry and fiber-based optical sensing is demonstrated for ultrasound detection. Pulse interferometry enables robust performance of the readout system in the presence of rapid variations in the environmental conditions, whereas the use of all-fiber technology leads to a mechanically flexible sensing element compatible with highly demanding medical applications such as intravascular imaging. In order to achieve a short sensor length, a pi-phase-shifted fiber Bragg grating is used, which acts as a resonator trapping light over an effective length of 350 &#181;m. To enable high bandwidth, the sensor is used for sideway detection of ultrasound, which is highly beneficial in circumferential imaging geometries such as intravascular imaging. An optoacoustic imaging setup is used to determine the response of the sensor for acoustic point sources at different positions.

Optical detection of ultrasound is impractical in many imaging scenarios because it often requires stable environmental conditions. We demonstrate an optical technique for ultrasound sensing in volatile environments with miniaturization and sensitivity levels appropriate for optoacoustic imaging in restrictive scenarios, e.g. intravascular applications.

Optical sensors of ultrasound are a promising alternative to piezoelectric techniques, as has been recently demonstrated in the field of optoacoustic ...

Imaging Denatured Collagen Strands In vivo and Ex vivo via Photo-triggered Hybridization of Caged Collagen Mimetic Peptides

Collagen is a major structural component of the extracellular matrix that supports tissue formation and maintenance. Although collagen remodeling is an integral part of normal tissue renewal, excessive amount of remodeling activity is involved in tumors, arthritis, and many other pathological conditions. During collagen remodeling, the triple helical structure of collagen molecules is disrupted by proteases in the extracellular environment. In addition, collagens present in many histological tissue samples are partially denatured by the fixation and preservation processes. Therefore, these denatured collagen strands can serve as effective targets for biological imaging. We previously developed a caged collagen mimetic peptide (CMP) that can be photo-triggered to hybridize with denatured collagen strands by forming triple helical structure, which is unique to collagens. The overall goals of this procedure are i)&#160;to image denatured collagen strands resulting from normal remodeling activities in vivo, and ii) to visualize collagens in ex vivo tissue sections using the photo-triggered caged CMPs. To achieve effective hybridization and successful in vivo and ex vivo imaging, fluorescently labeled caged CMPs are either photo-activated immediately before intravenous injection, or are directly activated on tissue sections. Normal skeletal collagen remolding in nude mice and collagens in prefixed mouse cornea tissue sections are imaged in this procedure. The imaging method based on the CMP-collagen hybridization technology presented here could lead to deeper understanding of the tissue remodeling process, as well as allow development of new diagnostics for diseases associated with high collagen remodeling activity.

Imaging Denatured Collagen Strands In vivo and Ex vivo via Photo-triggered Hybridization of Caged Collagen Mimetic Peptides

This procedure demonstrates in vivo near IR fluorescence imaging of collagen remodeling activities in mice as well as ex vivo staining of collagens in tissue sections using caged collagen mimetic peptides that can be photo-triggered to hybridize with denatured collagen strands.

Collagen is a major structural component of the extracellular matrix that supports tissue formation and maintenance. Although collagen remodeling is an ...

Fluorescence-quenching of a Liposomal-encapsulated Near-infrared Fluorophore as a Tool for In Vivo Optical Imaging

Optical imaging offers a wide range of diagnostic modalities and has attracted a lot of interest as a tool for biomedical imaging. Despite the enormous number of imaging techniques currently available and the progress in instrumentation, there is still a need for highly sensitive probes that are suitable for in vivo imaging. One typical problem of available preclinical fluorescent probes is their rapid clearance in vivo, which reduces their imaging sensitivity. To circumvent rapid clearance, increase number of dye molecules at the target site, and thereby reduce background autofluorescence, encapsulation of the near-infrared fluorescent dye, DY-676-COOH in liposomes and verification of its potential for in vivo imaging of inflammation was done. DY-676 is known for its ability to self-quench at high concentrations. We first determined the concentration suitable for self-quenching, and then encapsulated this quenching concentration into the aqueous interior of PEGylated liposomes. To substantiate the quenching and activation potential of the liposomes we use a harsh freezing method which leads to damage of liposomal membranes without affecting the encapsulated dye. The liposomes characterized by a high level of fluorescence quenching were termed Lip-Q. We show by experiments with different cell lines that uptake of Lip-Q is predominantly by phagocytosis which in turn enabled the characterization of its potential as a tool for in vivo imaging of inflammation in mice models. Furthermore, we use a zymosan-induced edema model in mice to substantiate the potential of Lip-Q in optical imaging of inflammation in vivo. Considering possible uptake due to inflammation-induced enhanced permeability and retention (EPR) effect, an always-on liposome formulation with low, non-quenched concentration of DY-676-COOH (termed Lip-dQ) and the free DY-676-COOH were compared with Lip-Q in animal trials.

Fluorescence-quenching of a Liposomal-encapsulated Near-infrared Fluorophore as a Tool for In Vivo Optical Imaging

The use of fluorophores for in vivo imaging can be greatly limited by opsonization, rapid clearance, low detection sensitivity and cytotoxic effects on the host. Encapsulation of fluorophores in liposomes by film hydration and extrusion leads to fluorescence quenching and protection which enables in vivo imaging with high detection sensitivity.

Optical imaging offers a wide range of diagnostic modalities and has attracted a lot of interest as a tool for biomedical imaging. Despite the enormous ...

Using Extraordinary Optical Transmission to Quantify Cardiac Biomarkers in Human Serum

For a biosensing platform to have clinical relevance in point-of-care (POC) settings, assay sensitivity, reproducibility, and ability to reliably monitor analytes against the background of human serum are crucial.
Nanoimprinting lithography (NIL) was used to fabricate, at a low cost, sensing areas as large as 1.5 mm x 1.5 mm. The sensing surface was made of high-fidelity arrays of nanoholes, each with an area of about 140 nm2. The great reproducibility of NIL made it possible to employ a one-chip, one-measurement strategy on 12 individually manufactured surfaces, with minimal chip-to-chip variation. These nanoimprinted localized surface plasmon resonance (LSPR) chips were extensively tested on their ability to reliably measure a bioanalyte at concentrations varying from 2.5 to 75 ng/mL amidst the background of a complex biofluid-in this case, human serum. The high fidelity of NIL enables the generation of large sensing areas, which in turn eliminates the need for a microscope, as this biosensor can be easily interfaced with a commonly available laboratory light source. These biosensors can detect cardiac troponin in serum with a high sensitivity, at a limit of detection (LOD) of 0.55 ng/mL, which is clinically relevant. They also show low chip-to-chip variance (due to the high quality of the fabrication process). The results are commensurable with widely used enzyme-linked immunosorbent assay (ELISA)-based assays, but the technique retains the advantages of an LSPR-based sensing platform (i.e., amenability to miniaturization and multiplexing, making it more feasible for POC applications).

This work describes a nanoimprinting lithography method to fabricate high-quality sensing arrays that work on the principle of extraordinary optical transmission. The biosensor is low-cost, robust, easy to use, and can detect cardiac troponin I in serum at clinically relevant concentrations (99th percentile cutoff &#8764;10-400 pg/mL, depending on the assay).

For a biosensing platform to have clinical relevance in point-of-care (POC) settings, assay sensitivity, reproducibility, and ability to reliably monitor ...

Switchable Acoustic and Optical Resolution Photoacoustic Microscopy for In Vivo Small-animal Blood Vasculature Imaging

Photoacoustic microscopy (PAM) is a fast-growing invivo imaging modality that combines both optics and ultrasound, providing penetration beyond the optical mean free path (~1 mm in skin) with high resolution. By combining optical absorption contrast with the high spatial resolution of ultrasound in a single modality, this technique can penetrate deep tissues. Photoacoustic microscopy systems can have either a low acoustic resolution and probe deeply or a high optical resolution and probe shallowly. It is challenging to achieve high spatial resolution and large depth penetration with a single system. This work presents an AR-OR-PAM system capable of both high-resolution imaging at shallow depths and low-resolution deep-tissue imaging of the same sample in vivo. A lateral resolution of 4 &#181;m with 1.4 mm imaging depth using optical focusing and a lateral resolution of 45 &#181;m with 7.8 mm imaging depth using acoustic focusing were successfully demonstrated using the combined system. Here,&#160;in vivo&#160;small-animal blood vasculature imaging is performed to demonstrate its biological imaging capability.

Switchable Acoustic and Optical Resolution Photoacoustic Microscopy for In Vivo Small-animal Blood Vasculature Imaging

Here a switchable acoustic resolution (AR) and optical resolution (OR) photoacoustic microscopy (AR-OR-PAM) system capable of both high resolution imaging at shallow depth and low resolution deep tissue imaging on the same sample in vivo is demonstrated.

Photoacoustic microscopy (PAM) is a fast-growing invivo imaging modality that combines both optics and ultrasound, providing penetration beyond the ...

Multimodal Volumetric Retinal Imaging by Oblique Scanning Laser Ophthalmoscopy (oSLO) and Optical Coherence Tomography (OCT)

While fluorescence imaging is widely used in ophthalmology, a large field of view (FOV) three-dimensional (3D) fluorescence retinal image is still a big challenge with the state-of-the-art retinal imaging modalities because they would require z-stacking to compile a volumetric dataset. Newer optical coherence tomography (OCT) and OCT angiography (OCTA) systems overcome these restrictions to provide three-dimensional (3D) anatomical and vascular images, but the dye-free nature of OCT cannot visualize leakage indicative of vascular dysfunction. This protocol describes a novel oblique scanning laser ophthalmoscopy (oSLO) technique that provides 3D volumetric fluorescence retinal imaging. The setup of the imaging system generates the oblique scanning by a dove tail slider and aligns the final imaging system at an angle to detect fluorescent cross-sectional images. The system uses the laser scanning method, and therefore, allows an easy incorporation of OCT as a complementary volumetric structural imaging modality. In vivo imaging on rat retina is demonstrated here. Fluorescein solution is intravenously injected to produce volumetric fluorescein angiography (vFA).

Multimodal Volumetric Retinal Imaging by Oblique Scanning Laser Ophthalmoscopy oSLO and Optical Coherence Tomography OCT

Here, we present a protocol to get a large field of view (FOV) three-dimensional (3D) fluorescence and OCT retinal image by using a novel imaging multimodal platform. We will introduce the system setup, the method of alignment, and the operational protocols. In vivo imaging will be demonstrated, and representative results will be provided.

While fluorescence imaging is widely used in ophthalmology, a large field of view (FOV) three-dimensional (3D) fluorescence retinal image is still a big ...