The working mechanism of OCT is akin to ultrasound (US) imaging [29,30]. In both techniques, waves are transmitted to the tissue under examination—light waves in the case of OCT and acoustic waves for US. These waves echo off the tissue structures, and the back-reflected waves from the tissue layers, which are interfaces between structures of different refractive indices, are detected. The echo delay is then measured to determine the depth at which the reflection occurred. Since OCT uses light waves, which travel significantly faster (i.e., 2×10⁵ times greater) than sound waves used in US, it is impossible to directly detect the train of back-reflected light with current photo-receivers, as they cannot resolve the very short time intervals between them. Instead, OCT utilizes a technique known as low-coherence optical interferometry, which is fundamental to all OCT implementations [29,30]. Optical interferometry is a method that uses the interference created by the superposition of 2 light waves with identical optical properties (e.g., amplitude and polarization), originating from the same light source. This technique is characterized by a coherence length, which is the physical distance over which interference is maintained. In OCT interferometry, the coherence length is only a few micrometers, due to the broad spectral bandwidth of the light source. This means that interference is observed only within a range of a few micrometers. Interference occurs only if the difference in the travel distances (optical path lengths) of the 2 light waves is within this short coherence length. Light that travels outside the coherence length will not interfere. Therefore, in a multi-layered sample with light transparency, it can be deduced that interferometry with low coherence can resolve interferences only if the optical paths are matched in length within the coherence length. This phenomenon of generating interference within the coherence limit is known as coherence gating [29,30].

OCT technology is implemented using low-coherence interferometers, which are categorized into 2 types based on the method of interference signal acquisition: time-domain OCT (TD-OCT) and Fourier-domain OCT (FD-OCT). TD-OCT represents the initial application of OCT technology, and Fig. 1 illustrates a typical schematic of a fiber-based TD-OCT system [31]. As shown in Fig. 1, an output beam from a broadband light source (40-100 nm at full width at half-maximum [FWHM] in spectrum) is split into 2 equal beams by a fiber coupler, directed to a reference arm and a sample arm, respectively. The light waves that are retro-reflected from the surface of the reference reflector (a reference mirror) and back-reflected from a sample (a human eye) are recombined at the fiber coupler, generating interference under the conditions described above. The resulting interference signal is detected by a photodetector, which converts the optical signal into an electrical signal. If the reference reflector is displaced in the depth (z) direction as shown in Fig. 1, the optical path length (OPL) of the reference arm is modulated, and then the coherence gate is temporally positioned inside the sample. The single axial (z) movement of the reference reflector generates a series of interferometric signals contributed by the internal layers of the sample in depth, and these are consecutively recorded by the photodetector (U). The envelopes of the recorded interference signal intensities are extracted via post-signal processing (A), representing the reflectivity at each sample layer. Therefore, the depth profile of the sample reflectivity values provides structural information on the sample at the beam position. This depth-scanned reflectivity profile is referred to as an A-scan.

Furthermore, the sample beam is laterally scanned in the x-direction using a mechanical scanner, while the reference reflector is scanned in the depth (z) direction, collecting A-scans from different beam positions. This collection of A-scans, also known as a B-scan, forms a cross-sectional (x-z) OCT image. By incorporating another scanner that slowly steers the beam along the y-axis, B-scans can be obtained at various positions in the elevational y-direction. Ultimately, the acquisition of B-scans by a pair of scanners creates an OCT data cube, representing the 3D structural information of the sample.

The second-generation OCT is known as Fourier or frequency domain OCT (FD-OCT) [32]. Unlike TD-OCT, which relies on time modulation of the OPL in the reference arm, FD-OCT utilizes the spectral information from interference signals to generate A-scans without moving the reference arm. Essentially, the broadband interference is captured spectrally in the wavelength domain. FD-OCT is categorized into 2 methods: spectral-domain OCT (SD-OCT) and swept source OCT (SS-OCT). These methods are designed to record the interference spectrum (spectral interferogram), enabling the immediate calculation of the depth scan (A-scan) through a Fourier-transform, thanks to the Fourier relation [33]. This capability substantially improves the imaging speed and signal-to-noise ratio by eliminating the need for mechanical scanning of the OPL, while maintaining image resolution.

Fig. 2A presents a schematic of an SD-OCT system, which is similar to TD-OCT; however, it features a spectrometer instead of a photodetector. This spectrometer comprises a dispersive element, such as a grating, and a line-array charge-coupled device (CCD) camera. The interference light from the fiber is dispersed by the grating into different wavelength components, which are then projected onto the line-array CCD camera. The camera captures the spectral interferogram (U), which is a composite of the spectra of all interference signals produced between each sample layer in depth and the reference arm, which is in a fixed position. Therefore, applying a fast Fourier-transform (FFT) to the spectral interferogram enables the extraction of information from the full-depth scan (A-scan).

However, in SS-OCT, the spectral components are encoded by time, rather than spatial separation [34]. The setup of SS-OCT, depicted in Fig. 2B, is analogous to TD-OCT and SD-OCT systems. The key difference lies in the replacement of the broadband light source with a spectrally scanning light source capable of rapidly sweeping a very narrow line-width (<0.01 nm) across a wide range of wavelengths. This sweeping action emits successive single-frequency lights directed towards the interferometer. During a single sweep at a rate exceeding 100 kHz, interference is generated for the swept single-frequency light, allowing each wavelength component of the interference signal to be sequentially detected by a photodetector, thus reconstructing a spectral interferogram (U). Additionally, the interferogram simultaneously contains reflectivity information for all depth layers of the sample, which can be extracted by applying an FFT, resulting in a depth profile of the sample (A). For both SD- and SS-OCT systems, volumetric data acquisition is facilitated using a pair of x-y mechanical scanners.

In addition, full-field OCT (FF-OCT), a variant of ultrahigh-resolution TD-OCT, has been developed to obtain OCT images without lateral scanning [35]. Unlike TD- and FD-OCT, which focus the light beam to a single point on the sample, FF-OCT tightly focuses multiple beam spots in parallel across the sample, creating a full-field illumination effect [36]. The back-reflected light signals from the sample are detected using a 2D pixel array photodetector, such as an area CCD or CMOS camera. This approach to illumination eliminates the need for lateral mechanical scans in OCT imaging. Fig. 3A depicts a schematic of a standard FF-OCT system, which is based on a double-beam Michelson interferometer with a pair of high numerical aperture (NA) microscope objectives (referred to as a Linnik interferometer). The beam emitted from the broadband light source enters the interferometer via a Kohler illuminator, which is necessary to generate even illumination across the sample and to ensure that an image of the illumination source is not visible in the resulting CCD image [37]. The beam is then equally split by a beam splitter into 2 beams directed towards separate reference and sample arms, where they are tightly focused onto the surface of a reflector (mirror) attached to a piezo actuator and onto the sample by identical high NA (>0.3 NA) microscope objectives. Interference occurs when the OPLs between the reference arm and the sample layer in the focal plane are within the coherence length, producing an en face (x-y) interference image captured by a high-speed area CCD camera. To decode the reflectivity values at the sample layer (i.e., sample structural information) from the interference image, phase-shifting interferometry (PSI) is applied to FF-OCT. This involves varying the OPL difference by oscillating the reference mirror by a few hundred nanometers through sinusoidal motion of the piezo actuator, leading to time-varying changes in the interference signal intensities [38]. Consequently, the interference images display varying fringe patterns over time, captured by the CCD camera during a single modulation period of piezo motion. An en face (x-y) oriented tomographic image (a sample reflectivity image) is then extracted through simple arithmetic calculations between the acquired interference images [35]. Therefore, FF-OCT enables en face OCT imaging throughout the depth of tissues, achieving depths of hundreds of microns with an isotropic resolution of 1 μm or less, which is 10 times finer than the resolution of typical OCT. This high-resolution imaging technique is preferentially used to delineate microscopic tissue structures [39].

Recent advances in FF-OCT have enabled the imaging of individual cells and their nuclei within tissue, which are not always visible in conventional FF-OCT due to significant backscattering from the surrounding tissue, a dominant component of the FF-OCT signal [40]. This method utilizes time-dependent interferometric signals modulated by the rapid motion of biological scattering elements such as nuclei, providing additional contrast correlated with subcellular metabolic activity in tissues or organs. This variant, known as dynamic FF-OCT (D-FFOCT), is achieved using the typical FF-OCT setup, where modulation of the reference arm for PSI is no longer necessary, as shown in Fig. 3B. However, the interference signal can be modulated by the movements of scatterers within the sample itself. For example, if a single scattering particle in the sample moves uniformly along the vertical axis (i.e., z-axis) within the coherence gate, the linear z-motion of the scatterer induces a displacement of the OPL in the sample arm, resulting in a time variation in the interference signal intensity (refer to Fig. 3B). By calculating the standard deviation of the interference signal over time, the dynamic signal component caused by the moving scatterers can be isolated, while the signal from highly scattering static tissue is suppressed [40]. The ability of D-FFOCT to visualize dynamics at the subcellular scale is beneficial for identifying active intracellular features at short time scales without the need for exogenous contrast agents [40,41].

Unlike conventional microscopy, a unique advantage of OCT is that axial resolution and lateral resolution are decoupled from one another. Fig. 4 illustrates several OCT imaging parameters. The axial resolution is denoted as the resolving power to discern 2 discrete objects apartwith a difference in depth (i.e., the z-axis), which is an important imaging parameter for high-resolution OCT imaging. The axial resolution (Δz) is equivalent to the coherence length (l_c) of the light source, theoretically defined as:

where λ₀ and Δλ, and n are the central wavelength and the spectral bandwidth at full width at half maximum (FWHM) of the light source spectrum, and a refractive index (RI) of the sample, respectively [42]. Eq. (1) indicates that the axial resolution is fully dependent on the light source, which means that a wider spectrum ensures better axial (depth) resolution. Currently, the most commonly used light source in fiber-based OCT is a broadband superluminescent diode (SLD) with a bandwidth of up to 100 nm. With a center wavelength of 1,300 nm, the theoretical axial resolution in tissue (RI=1.4) is calculated to be 5.3 μm. Furthermore, sub-micrometer resolution can be achieved by using a much wider broadband light source at the shorter central wavelength [43]. However, the use of SLD is inappropriate for free space or parallel OCT (e.g., FF-OCT) due to its high spatial coherence [44], which generates strong speckles on the interference image due to multi-scattering induced crosstalk across the 2D pixel array of detector. Therefore, a temporally and spatially incoherent light source is imperative for free space OCT imaging, and spatiotemporally less coherent thermal light or white LED is a good choice for a light source to yield depth resolution less than 1 μm [37].

Meanwhile, the lateral resolution (Δx) of OCT, a resolving powerdefined as its ability to discern two 2 discrete objects apart in the transverse direction (x or y-axis), is equivalent to the size of the focal beam spot. This resolution adheres to the principles of conventional microscopy, which are governed by Abbe’s diffraction formula:

where f and d are the focal length of the microscope objective (OBJ) and the beam spot size, respectively. NA is the numerical aperture, a ratio of d to f. Eq. (2) indicates that the lateral resolution can be improved by increasing the NA. In other words, the use of an OBJ with a larger NA ensures better lateral resolution, although this requires a compromise with the imaging depth defined as the depth of focus, πΔx22λ.

Organoid research utilizing OCT has predominantly concentrated on retinal organoids, reflecting the common application of OCT in both ophthalmology research and clinical settings. In 2017, Browne et al. [45] pioneered the use of OCT imaging to explore structural changes in live retinal organoids throughout their development. Human pluripotent stem cell (hPSC)-derived retinal organoids replicate many aspects of embryonic retinal development, including the formation of a bi-layered optic cup and light-detecting photoreceptors [46]. These organoids are proposed as models for studying human retinal development and diseases, as well as for drug screening purposes. In their research, Browne et al. [45] employed a commercial high-resolution SD-OCT system (Spectralis; Heidelberg Engineering Inc., Carlsbad, CA, USA) to assess the microanatomic organization of hPSC-derived retinal organoids at various developmental stages. The specifications of the system include an 870 nm wavelength SLD light source, an axial resolution of 3.9 μm, a lateral resolution of 14 μm, a scanning depth of 1.9 mm, a maximum field of view (FOV) of 30×30 degrees, and a scanning speed of 40,000 A-scans per second. Additionally, the OCT system incorporates a confocal scanning laser ophthalmoscope with an 820 nm wavelength for fundus imaging. The organoids were imaged while suspended in 0.1 mL of culture media within 0.2 mL polypropylene tubes, positioned in the imaging plane of the device. Fig. 5 shows OCT images of the retinal organoid at different stages (from 46 to 151 days in culture [DIC]), revealing the maturation of the organoid [45]. For instance, OCT displays a distinct superficial hypo-reflective band (indicated by arrowheads in Fig. 5F) present at the later stage (151 DIC). This low signal band appears as a thin cellular structure, sandwiched between hyper-reflective cellular structures. The outer rim of the band is estimated to be approximately 25 μm thick, coinciding with a lamellar structure observed on histology (Fig. 5G) of the same organoid at 151 DIC.

Similar morphological alterations in retinal organoids were also identified using OCT by another research group in 2019 [47], as shown in Fig. 6. This group employed the high-resolution SD-OCT system Telesto-II from Thorlabs Inc. to monitor extensive areas of retinal organoid cultures. The 3D imaging capability of this system offers comprehensive insights into the internal anatomy and the overall shape and surface topography of the organoids. The system utilizes an SLD light source at 1300 nm with a broad spectral bandwidth (>80 nm at FWHM), achieving an imaging depth of 2.6 mm in water, with axial and lateral resolutions of 5.5 μm and 13 μm, respectively. It operates at a scanning speed of 76,000 A-scans per second. As shown in Fig. 6, OCT delineated characteristic morphological features in 3D retinal organoids during development: an open bowl-like contour for stage 1 (35 DIC) organoids (Fig. 6A-C), and an enclosed spheroidal shape with a hollow core for stage 2 (54 DIC) organoids (Fig. 6D-F). OCT images of stage 3 (250 DIC) organoids revealed more complex structural elements, including hair-like surface appendages and alternating high and low reflectance layers near the outer rim (Fig. 6G-I), similar to the OCT images at 151 DIC (Fig. 6C and 6F). Histological images of stage 1, 2, and 3 organoids confirmed the differences in cellular stratification and interior structure implied by OCT (Fig. 6J-L). Overall, OCT not only described the distinct morphological characteristics, but also uncovered additional anatomical features, distinguishing organoids from each of the 3 stages.

McLelland et al. [48] utilized OCT in their study on retinal organoids to explore whether sheets of hESC-derived retinal organoids could differentiate, integrate, and enhance visual function when transplanted into an immunodeficient rat model with severe retinal degeneration (RD). RD, including conditions like age-related macular degeneration, is a progressive retinopathy resulting from genetic mutations and/or environmental or pathological damage to the retina. This leads to irreversible vision loss due to the permanent loss of photoreceptors [49]. While the transplantation of fetal retinal sheets into patients with RD has been shown to improve vision, scaling this approach to a clinical level is challenging due to ethical and supply constraints. However, using sheets from retinal organoids as a transplant material may offer a scalable, effective, and safe therapeutic alternative for vision restoration in RD [50]. Therefore, the study focused on assessing the efficacy of transplanting hESC-derived retinal sheets into rats with severe RD. For the preparation of retinal sheets, RPE-free rectangular sheets measuring 1.0-1.7×0.6 mm were dissected from hESC-derived retinal organoids between days 30 and 65 of differentiation. During transplantation, a small incision was made posterior to the pars plana, and the dissected retinal organoid sheet was carefully inserted into the subretinal space of the left eye of a rat in advanced stages of RD, where most rod photoreceptors had been lost. A commercial SD-OCT system designed for preclinical research (Bioptigen Envisu R2200; Bioptigen, Durham, NC, USA) was used to monitor the transplanted rats (n=25), with imaging conducted every 1 to 2 months starting 2 weeks post-surgery and continuing up to 9.5 months. This system, optimized for posterior imaging of small animals, provides an axial resolution of less than 2 μm in tissue and an acquisition rate of 32,000 A-scans per second. Fig. 7 shows the OCT imaging result of transplant development and growth post-surgery [48]. Fig. 7A is a fundus image of the RD posterior eye with a transplant (dashed area), and representative OCT images (Fig. 7B-E) in the transplant show that the retinal organoid sheets were subretinally deposited below the degenerated host retina. Notably, the formation of hyper-reflective photoreceptor rosettes, which are light-sensing structures indicated by arrows, was observed within the transplants at various days post-surgery, indicating that the transplanted sheets had differentiated, integrated, and begun functioning as photoreceptors. Additionally, improvements in visual function in the transplants were assessed using optokinetic testing and electrophysiological recording in the superior colliculus [51]. In this study, OCT proved to be a valuable tool for longitudinally monitoring the growth of human transplants derived from ESCs in vivo.

The first demonstration of FF-OCT imaging of retinal organoids was achieved by the Grieve group in 2020 [52]. They employed D-FFOCT to noninvasively capture both the structural and functional aspects of living hPSC-derived retinal organoids. This technique produced colored images with endogenous contrast related to organelle motility, featuring subcellular spatial resolution and millisecond temporal resolution. The FF-OCT system utilized in this research was a custom-built time-domain FF-OCT equipped with a broadband 660 nm LED source, achieving an axial resolution of 1.7 μm. High-magnification water-immersion objectives (40×, 0.8 NA) were used, providing a lateral resolution of 0.5 μm and a limited FOV measuring 320×320 μm². To generate a D-FFOCT image, 512 frames of interferograms produced at the coherence gate and superimposed in the focal plane were sequentially captured by a high-speed complementary metal-oxide-semiconductor camera (Quartz 2A750; Adimec, Eindhoven, The Netherlands, operating at 750 frames per second). These frames were subsequently processed to extract local fluctuations and render them in color [52].

Fig. 8 shows D-FFOCT images of a 28-day-old (28 days post coitum [DPC]) retinal organoid culture, corresponding to the optic vesicle stage during retinogenesis [52]. Fig. 8A is a D-FFOCT image in which image brightness is linked to fluctuation amplitude (strength of organelle oscillation). Meanwhile, Fig. 8B is a colored D-FFOCT image, in which the color is linked to fluctuation speed (frequency of organelle oscillation), ranging from blue (slow) to red (fast) through green (in between). A 3D reconstruction of the organoid’s dynamics is depicted in Fig. 8C, along with a sub-volume in Fig. 8D, highlighting the layered internal organization of retinal progenitor cells (5 μm in diameter). A cross-section is shown in Fig. 8E, in which the elongated shape of cells is visible. Red arrows in the images indicate surface cells exhibiting fast dynamics. Fig. 8F presents a series of time-lapse D-FFOCT images obtained at 50 μm depth of the same organoid to examine its temporal evolution over a period of 3 hours. The images in the top row show changing cell dynamics near a clear boundary (dotted line) between 2 distinct types of cells. The cells on the right side of the boundary exhibit faster and stronger dynamics, suggesting a differentiation process towards specific retinal lineages. This work demonstrated the utility of D-FFOCT for monitoring the time course of retinal organoid developmental processes at the single-cell level, a task that has been challenging for other OCT technologies.

Since 2020, several studies have utilized OCT in research on other organoids. For example, Deloria and colleagues explored the use of OCT in analyzing human placenta-derived trophoblast organoids (TB-ORGs) [53]. These organoids, developed recently, mimic the developmental program of the early human placenta, an extraembryonic organ essential for supporting the fetus during intrauterine life. TB-ORGs offer a highly reproducible and stable in vitro model for studying not only developmental but also physiological and pathophysiological processes in early pregnancy [54,55]. The researchers aimed to use 3D OCT to investigate pathophysiological processes in placental organoids, with the objective of modeling placental diseases and enhancing understanding of fetal-maternal interactions during placentation. They hypothesized that 3D OCT could reveal the internal structures of living placental organoids without the need for sample preparation, and correlate these observations with the organoids' differentiation status. For the 3D OCT imaging, the team employed a high-resolution SD-OCT system featuring a central wavelength of 846 nm and a 3 dB bandwidth of 133.3 nm, achieving an axial resolution of 2.48 μm in the organoid tissue. The lateral resolution was also measured at 2.48 μm. OCT images were acquired at a scan speed of 5000 A-scans per second.

Fig. 9 shows OCT imaging results of TB-ORG cultures (n=3) embedded in Matrigel. Fig. 9A and 9B display bright-field microscopy and corresponding OCT cross-sections of the same TB-ORG, respectively [53]. While the bright-field microscope image barely reveals the interior of the organoid, the OCT clearly visualizes micro-cavities within the organoid, which are characteristic of the internal structure of normal TB-ORGs. Additionally, OCT has uncovered morphological differences between the control TB-ORG and the TB-ORG treated with a p38 MAPK inhibitor. This inhibitor blocks the differentiation of syncytiotrophoblasts, which is essential for hormone production and transport functions. The top row in Fig. 9C exhibits control TB-ORG samples, in which both OCT and hematoxylin and eosin (H&E)-stained images reveal the existence of cavities in the central region of the organoids. In contrast, the p38 MAPK inhibitor-treated TB-ORG (bottom row in Fig. 9C) does not exhibit these cavities, as shown in both OCT and H&E-stained images. This work suggests that OCT imaging can be routinely included in organoid research to facilitate the monitoring of in situ samples.

In 2021, Gil and colleagues conducted a study using patient-derived cancer organoids (PCOs) as models for cancer drug screening [56]. They employed a custom SS-OCT system to rapidly assess both the volumetric growth and the drug response of these PCOs. The OCT system featured a swept-wavelength source from Axsun Technologies, Pittsburgh, PA, USA with a center wavelength of 1,040 nm, a bandwidth of 109 nm, and a wavelength sweeping rate of 100 kHz, which is equivalent to 100,000 A-scans per second. The system achieved axial and lateral resolutions of 6.5 μm and 16 μm, respectively. PCOs were created from 2 metastatic colorectal cancer lines derived from patients, labeled P1 and P2. These were resuspended in base medium and combined with Matrigel in a 1:1 ratio. The resulting cell-Matrigel mixture was then pipetted onto the glass surface of each well in a 24-well glass bottom plate. Over a period of 48 hours, longitudinal OCT imaging was performed using a 3D single-particle tracking technique on individual PCOs in culture. Organoid volume was subsequently estimated through quantitative analysis of the OCT data. The imaging results revealed that the application of metabolic inhibitors and cancer therapies—including cyanide, 2DG, cisplatin, and paclitaxel—to P1 and P2 led to a reduced volumetric growth rate in the PCOs compared to untreated control PCOs [56]. This study thus demonstrated the utility of OCT and 3D single-organoid tracking as effective tools for quantitatively evaluating the therapeutic impact on cancer organoids.

Very recently, Ming et al. [57] conducted a longitudinal OCT study on human heart organoids (hHOs) to characterize their morphology and motion. In the development of organoid models for human organs, hHOs have made relatively less progress due to the structural complexity of the human heart. Recently, hHOs derived from hPSCs have been developed and evaluated as a reproducible and robust in vitro model for recapitulating the human fetal heart [58,59]. The morphogenesis of heart-forming organoids is of great interest, as these organoids are expected to capture the development of heart-specific structures, including functional chambers, blood vessels, heart valves, and more. The aim of this study was to use OCT to investigate the morphogenesis of hHOs and further understand their beating patterns. The OCT setup used in this work was a customized SD-OCT system equipped with an SLD light source (SLD1325; Thorlabs Inc., Newton, NJ, USA) with a central wavelength of 1,320 nm and a spectral bandwidth of 110 nm. The axial and lateral resolutions were 4.9 μm and 7 μm, respectively, and the imaging speed was 20,000 A-scans per second. To further improve the axial resolution, the light source was replaced with a supercontinuum laser (SC450; Fianium, Bregnerodvej, Denmark) with a wavelength range from 1,175 to 1,420 nm. The broader bandwidth of the light source yielded an axial resolution of approximately 2.36 μm. OCT imaging was performed on hHO cultures daily from day 1 to day 22 and every other day from day 24 to day 30.

Longitudinal 3D OCT imaging captured dynamic morphological changes in the development of hHOs (n=32) over a 30-day period, as illustrated in Fig. 10A. Small cavities and large chambers within the hHO became visible between days 3 and 4, with their merging and enlargement observed from day 6 to day 16. Interestingly, from day 16 onward, both the number and sizes of the chambers decreased, and by day 30, the large chamber was no longer visible. However, the diameter and volume of the hHO significantly increased and plateaued starting from day 10. The development and remodeling of these chambers indicate the growth and maturation of the hHOs. Additionally, the beating of the hHO was detected through repetitive B-scans at the same location of the organoid, beginning between days 6 and 12. Fig. 10B shows a representative time series of B-scans capturing the chamber of interest during a beating cycle (contraction+relaxation). During contraction, the bottom chamber contracted and the valve-like structure (indicated by a red arrowhead) closed, and during relaxation, it reopened (Fig. 10B). The beating was quantified by measuring the distance between the upper and lower boundaries of the chamber across all frames in the time series (indicated by a red dashed box in Fig. 10B). The blue traces on the left in Fig. 10C represent the distance profiles measured from 2 heart organoids, indicating the hHOs’ beating patterns. The average beat rate was estimated to be 22.6±13.05 beats per minute (BPM), and the average interbeat interval (IBI) was 4.2±3.58 seconds for 6 hHOs. These waveforms closely resembled the typical form of calcium transients recorded from fluorescence Ca²⁺ imaging of the same organoids (blue traces on the right in Fig. 10C). The heart rate and IBI measured by OCT were in close agreement with those (beat rate: 22.0±7.21 BPM, IBI: 3.1±2.89 seconds) by Ca²⁺ imaging). These experimental results demonstrate that hHO models and OCT imaging hold significant promise for studying the human fetal heart. They provide a foundational platform for investigating human hearts at different developmental stages and exploring mechanisms of heart disease.