Previous works have demonstrated the feasibility of OCT to extract important information in the myocardium, including fiber orientation and identification of radiofrequency ablation lesions.12–14 These works were conducted in the 1300-nm optical window mainly for higher penetration depth in the myocardium, while many details were still buried in the OCT images. This may be due to the insufficient axial resolution as well as the relatively weak backscattering signal collected from the myocardium sample at this optical window. In order to characterize the pathology-related tissue types in the myocardium, higher resolution and better contrast of the OCT images are needed and can be achieved by moving to a shorter wavelength regime. Since OCT signals are generated based on the detection of backscattered light, another benefit to working in the shorter wavelength regime is that it favors stronger light scattering, which will provide more information about the structures inside the tissue. Moreover, the optical properties of different types of tissue in the shorter NIR regime are more diversified compared with the 1300-nm window. Working in this regime will enable functional extensions of the standard OCT imaging, such as spectroscopic OCT. Based on these reasons, most of the ultrahigh-resolution (UHR) works in OCT chose to employ light sources that were in the shorter NIR regime.15–24 In general, UHR OCT categorizes OCT systems with axial resolution lower than in air. The early UHR OCT systems were built in time domain (TD) configuration with the help of state-of-the-art femtosecond lasers to generate ultrashort pulses.15,16,25 After balancing the dispersion mismatch in the interference arms, an axial resolution as high as in air was reported.15 The high-lateral resolutions of these systems were achieved by adopting high-NA objectives in the sample arm, which was necessarily associated with a reduced depth of focus and eventually limited the imaging depth to around 0.5 mm. Later on, UHR OCT was implemented in spectral domain (SD) configurations in both bench-top systems17–21 and endoscopic systems.22,23 SD-OCT is known to have a sensitivity advantage over its TD counterparts owing to the paralleled detection in the SD,24,26 and is more preferable in practice. The key component of UHR SD-OCT is the spectrometer. In particular, due to the finite sampling area of a single detector pixel, fringe visibility was compromised at higher spatial frequency, which causes a sensitivity falloff at deeper imaging ranges. For example, Liu et al.19 demonstrated an axial resolution of in air in the 800-nm spectral window using a spectrometer with a spectral range of around 400 nm captured by a 2-k pixel line camera. Despite the superb axial resolution, the imaging depth was limited to 0.5 mm, mainly due to a low spectral resolution in the -domain. Alternatively, Yadav et al.20 reported a UHR OCT system with axial resolution and lateral resolution in air. A modified Czerny–Turner spectrometer was designed for a similar spectral range but with a detector of 8 k pixels. Though not all of the pixels were used to cover the bandwidth, it helped to extend the imaging depth to around 1.2 mm. However, the pixel size of led to an insufficient coupling efficiency and resulted in a lower SNR. Moreover, the limited spectral resolution as well as the inaccurate -domain resampling method for data interpolation led to a 6-dB falloff range around 0.5 mm. Recently, Xi et al.23 reported a UHR SD-OCT endoscope with axial resolution in air produced by a linear- spectrometer, in which a more complicated design and alignment of the optics were required to achieve a good performance. Still, the reported imaging depth was limited to around 1.2 mm with a total sensitivity falloff of 16 dB for the entire imaging range.