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Published in final edited form as: Ultrasonics. 2020 Oct 10;110:106263. doi: 10.1016/j.ultras.2020.106263

Ultrasonic Elastography to Assess Biomechanical Properties of the Optic Nerve Head and Peripapillary Sclera of the Eye

Xuejun Qian 1,2,#, Runze Li 1,2,#, Gengxi Lu 1,2, Laiming Jiang 2, Haochen Kang 1, K Kirk Shung 1, Mark S Humayun 1,2, Qifa Zhou 1,2
PMCID: PMC7736296  NIHMSID: NIHMS1637939  PMID: 33065466

Abstract

Purpose:

To quantitatively investigate both optic nerve head (ONH) and peripapillary sclera (PPS) biomechanical properties of porcine eyes through an ultrasonic elastography imaging system in response to both increasing and decreasing intraocular pressure (IOP).

Methods:

The Young’s modulus of the ONH and PPS were assessed using our high resolution ultrasonic imaging system which utilized a mechanical shaker to induce shear waves and an off-axis aligned 40 MHz needle transducer to track micron-level displacement along the direction of wave propagation. In this study, imaging on a total of 8 ex vivo porcine eyes preloaded with IOPs from 6 mmHg to 30 mmHg was performed. To have a better understanding of the effect of varying IOP on biomechanics, both increasing and decreasing IOPs were investigated.

Results:

The increase of the Young’s modulus of ONH (92.4 ± 13.9 kPa at 6 mmHg to 224.7 ± 71.1 kPa at 30 mmHg) and PPS (176.8 ± 14.3 kPa at 6 mmHg to 573.5 ± 64.4 kPa at 30 mmHg) following IOP elevation could be observed in the reconstructed Young’s modulus of the shear wave elasticity (SWE) imaging while the B-mode structural images remained almost unchanged. In addition, for the same IOP level, both ONH and PPS have a tendency to be stiffer with decreasing IOP as compared to increasing IOP.

Conclusions:

Our results demonstrate the feasibility of using our ultrasonic elastography system to investigate the stiffness mapping of posterior eye with high resolution in both increasing and decreasing IOPs, making this technique potentially useful for glaucoma.

Keywords: high frequency ultrasound elastography, shear wave propagation, optic nerve head, peripapillary sclera, glaucoma

I. Introduction

Glaucoma is a leading cause of irreversible blindness, and it is estimated that by 2020 the number of people suffering glaucoma will reach 80 million worldwide [1]. Although elevated intraocular pressure (IOP) is the primary risk factor for the development of glaucoma, the mechanisms by which elevated IOP eventually leads to damage and loss of neural function are still unclear [2]. It is also unclear how sensitivity to IOP varies and interacts with other risk factors for glaucoma. The optic nerve head (ONH) is the primary site of damage in glaucoma, and the mechanical properties of the adjacent peripapillary sclera (PPS) are known to strongly influence the stresses and strains of the ONH where the retinal ganglion cell axons exit the eye [3, 4]. Specifically, elevated IOP may cause the scleral canal opening to widen circumferentially and the ONH to bow posteriorly, resulting in axon dysfunction, death, or activation of detrimental cellular response [5, 6]. It is therefore important to characterize the mechanical properties of the posterior eye under various IOP conditions [7].

Due to their location in the posterior segment of the eye, it is difficult to investigate the biomechanical properties of ONH and PPS in a non-invasive manner. Tensile testing [8, 9] and atomic force microscopy (AFM) nanoindentation [10] are two commonly used methods to measure the biomechanical properties of the ONH and PPS. Both methods require targeted tissue to be dissected from the eye globe into stripes, after which cryosections are further prepared for AFM testing. To maintain the complex physiological loading conditions of the eye in its three dimensional configuration, in vitro inflation testing of the posterior sclera tissue has been developed as an improvement to the uniaxial strip test methodology [11]. However, with this technique, the deformation force must still be manipulated by time-dependent IOP changes invasively, which impedes potential application.

Elastography is an emerging imaging modality to quantify biomechanics in soft tissue, mostly in a non-invasive manner. For ophthalmology, the more often reported elastography techniques such as magnetic resonance elastography (MRE) [12] and conventional ultrasound elastography (UE) [13] are not suitable mainly because of their low spatial resolution. Recently developed optical coherence elastography (OCE) techniques, utilizing optical coherence tomography (OCT) [14] to detect the propagation of induced elastic waves, have wide application in characterizing biomechanics in cornea [15], lens [16] and retina [17]. However, owing to the limited penetration depth, OCE systems lack the ability to image the deeper ONH region underneath the lamina cribrosa and non-transparent sclera tissue. It is notable that high frequency ultrasound has become an indispensable technique for ophthalmic imaging owing to its natural advantage of balanced spatial resolution and penetration depth [18, 19].

To investigate the biomechanical properties of posterior eye, especially for IOP-related deformation, some research studies have developed ultrasound imaging-based methods. Specifically, Alam et al. [20] implemented the sonoelastic Doppler ultrasound method to investigate the relations of mechanical resonance frequency (which is directly related to tissue elasticity) of the eye with elevated IOP using various models. Pavlatos et al [21] developed a high frequency ultrasound speckle tracking method to investigate regional deformation of the ONH and PPS during IOP elevation on ex vivo porcine eyes. Later, Ma et al [22] implemented the same technology using human donor globes. However, in their setup, the displacement vector must be calculated between images acquired at different IOP levels (i.e., the deformation force was generated by the changing IOP), which may not be feasible in practice. To generate a more controllable and stable deformation, an external pushing force generated by an air-pulse, acoustic radiation force (ARF) or mechanical shaker is preferred. Among these pushing strategies, the mechanical shaker is preferred given the requirement for generating sufficient deformation in posterior locations and also due to potential safety issues [23]. In addition to the pushing method, previously reported results only present either regional displacements or strain, which are all qualitative parameters of the tissue biomechanics. Without calibration of the applied force, the absolute Young’s modulus or elasticity cannot be conclusively determined.

In this study, we developed and implemented a mechanical shaker based ultrasonic elastography technique that has the capability to provide a quantitative estimation of tissue biomechanical properties via shear wave elasticity (SWE) imaging technology with high spatial resolution. The goal of this study is to quantify and map the mechanical properties of both ONH and PPS in response to either increasing or decreasing IOP for the purpose of better understanding the mechanism of development of Glaucoma.

II. Materials and methods

A. Experimental Setup

A schematic diagram of the experimental setup and synchronized timing sequence is shown in Figure 1. Due to strict U.S Food and Drug Administration (FDA) 510k standards for ophthalmic exposure, a mechanical shaker (mini-shaker type 4810; Bruel & Kjaer, Duluth, Georgia, USA) was used here as the pushing source in place of the low frequency high power ultrasound transducer used in our previous studies [18, 19, 24]. To precisely track tissue motion and elastic wave propagation caused by the mechanical shaker, a 40 MHz needle transducer was designed and fabricated in this study.

Fig. 1.

Fig. 1.

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Schematic diagram of the shaker induced ultrasonic elastography system and the synchronized timing sequence.

To acquire SWE imaging, the high frequency needle transducer was mounted on a 3-axis translation motorized linear stage (SGSP33-200, OptoSigma Corporation, Santa Ana, CA, USA) for mechanical scanning. Owing to the small size of the needle transducer aperture which is less than 0.4 mm in diameter, the pushing force generated by the mechanical shaker is applied almost orthogonally to the imaging subject, meaning only shear waves are generated. In order to track shear wave propagation, the shaker was fixed at the target position while the needle transducer was controlled by a linear stage which was moved based on the designed distance between the positions of the mechanical shaker and needle transducer. To sufficiently cover the ONH and PPS region, the step size and scanning distance were set to 42 μm and 8.4 mm, respectively. In addition, to avoid any issues with vibration of the needle during movement, the time delay between successive positions was set to 100 ms, including the data acquisition time and extra wait time.

During the experiment, an arbitrary function generator (AFG 3252C, Tektronix, Beaverton, OR, USA) generating a single pulse signal with 500 μs pulse duration was connected to a power amplifier (Type 2718, Bruel & Kjaer, Duluth, Georgia, USA) which was itself coupled to transmit an amplified signal to the mechanical shaker to induce tissue deformation. It is important to note that an impulsive pushing force is benefit to generate broadband shear wave while reducing the deformation of the tissues [25], Thus, based on our previous study [26], 500 μs pulse duration was selected here as a balance between the high temporal resolution (transient pushing) and stability of the shaker (i.e., too short pulse duration will cause incorrect pushing pattern of the shaker). In addition, the pulse amplitude was set to 600 mV in order to induce sufficient pushing force which can ensure that the displacement of the shear wave was detectable within the imaging region of the ONH and PPS. The needle transducer was set in conventional pulse-echo mode and was driven by a pulser/receiver (JSR500, Ultrasonics, Pittsford, NY, USA) with a pulse repetition frequency (PRF) of 20 kHz. After 20–80 MHz analog band-pass filtering (Mini-Circuits, Brooklyn, NY, USA) to remove signal contamination, the radiofrequency (RF) ultrasonic data was captured using a 12-bit digitizer (ATS9360, Alazartech, Montreal, QC, Canada) at a sampling rate of 1.8 GHz and stored for off-line analysis. To reduce system jitter, the same clock was used to synchronize the digitizer, pulser/receiver, and arbitrary function generator. To record the initial tissue position at each scanning position, the mechanical shaker was excited 50 μs after the needle transducer started to acquire data. At each tracking location, a total of 100 A-lines were acquired for each M-Mode image. All parameter settings were kept constant for all measurements.

B. Post-processing and data analysis

Data analysis was performed using MATLAB 2018a software (The MathWorks, Natick, MA, USA). At each M-mode, the first tracking A-line was served as the reference, and then dynamic tissue displacements were calculated using 1-D nonnalized cross-correlation [27] with a symmetric search region and 1.5λ window size (λ is the wavelength of the tracking transducer) between the reference A-line and the rest of A-lines in the M-mode. By repeating such procedures along lateral positions, we achieved a 3D matrix (expressed as lateral positions, depth positions and time) with displacement values as its index.

As demonstrated by previous studies [18, 28], using a small propagation distance to reconstruct the local shear wave speed (SWS) is preferred to sustain a high resolution SWS map. However, both ONH and PPS are associated with high elasticity, especially at high IOP. Therefore, in this study, a relatively long lateral propagation distance interval (756 μm) was selected to quantify each pixel of the SWS image mapping with a good balance between precision and resolution.

More specifically, we first converted the 3D matrix into the 2D spatiotemporal map (propagation distance versus time shifts curve) at each depth position where the wave propagation distance was measured by accumulation of lateral positions and the time shift was defined as the time to reach the peak displacement at each dynamic displacement. Then, the original spatiotemporal map was 2D interpolated to a finer grid size using a spline function. Next, the SWS at each pixel of image was estimated by applying a linear regression to a subset of 756 μm interval (lateral propagation distance) in the interpolated spatiotemporal map. By repeating the linear regression procedures to these subsets with a 42 μm (the moving step size of the needle transducer) distance shift each time, we achieved SWS image mapping. The final reconstructed SWS speed images were generated after applying a 3 × 3 median filter so as to increase the signal-to-noise-ratio (SNR). To quantify the Young’s modulus imaging mapping, we used the well-established equation (1) to convert SWS to Young’s modulus in each pixel,

E=3·ρ·cs2 (1)

where ρ presents the tissue density and cs is the SWS.

C. Phantoms and Biologic tissue preparation

Gelatin (Gelatin G8-500, Fisher Scientific, USA) based tissue-mimicking phantoms with equal concentration of silicon carbide powder (S5631, Sigma-Aldrich, St.Louis, MO, USA) and sound scatters were fabricated. The stiffness of each phantom was controlled by gelatin concentration. Three phantom types were made, including two homogeneous phantoms with different stiffness and a side-by-side phantom with two different stiffness levels in a single structure. The stiffness of the two homogeneous phantoms was tested using the gold standard – uniaxial mechanical testing (Model 5942, Instron Corp., MA, USA), and the results showed 8.1 ± 0.8 kPa for the soft phantom and 25.4 ± 1.9 kPa for the stiff phantom (the mean and standard deviation were determined by measuring each phantom five times), respectively.

Eight fresh, unscalded porcine eyes were obtained from a local service-oriented company (Sierra Medical Science, Inc., Whittier, CA, USA) within 24 hours of death in this study. The extraocular tissues were trimmed away, and the optic nerve was cut to be flush with the outer surface of the PPS. Eyes were mounted on a customized holder with the ONH and PPS exposed on the top. The lens and vitreous were removed and two ports were inserted into the eye chamber through the corneal limbus. One port was connected to the balanced saline solution bag set at various heights to manipulate IOP. The other port was connected to a pressure sensor (Model SPR-524, Millar Inc, TX, USA) to read the true IOP inside the chamber. In this study, five IOPs at both increasing and decreasing levels were investigated, including 6 mmHg, 12 mmHg, 18 mmHg, 24 mmHg and 30 mmHg.

III. Results

A. Phantom results

The tissue mimicking phantoms were first used to verify the stability and accuracy of the shaker-induced ultrasonic elastography system. Figure 2 (a–b) shows the spatiotemporal displacement maps of the two homogeneous phantoms at a certain depth location (randomly selected depth at the region of interest). As can be observed in Fig. 2, owing to homogeneity of the phantom, constant generated SWS resulted in a linear relationship between the shear wave propagation distance and the propagation time. Using linear regression of the data in the spatiotemporal displacement maps, the SWS of the soft phantom and stiff phantom were calculated to be 1.6 ± 0.1 m/s, 2.8 ± 0.3 m/s, respectively. The corresponding Young’s modulus values calculated using equation (1) are consistent with the gold standard value reported above, and demonstrate that our imaging system has the ability to accurately determine the absolute Young’s modulus. In addition to showing spatiotemporal information calculated at a certain depth, the uniform color mapping depicted in Fig. 3 (b,d) also indicates consistency of the obtained Young’s modulus values along the depth direction as expected for the two homogenous phantoms.

Fig. 2.

Fig. 2.

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The spatiotemporal displacement map of (a,b) the homogenous phantoms and (c) the side-by-side phantom. SWSs is the shear wave speed of soft phantom while SWSh indicates the stiffer phantom. The inflection point in (c) is the boundary between the soft and stiff part of the side-by-side phantom.

Fig. 3.

Fig. 3.

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The B-mode image and the reconstructed Young’s modulus SWE map of the homogenous phantoms with, (a,b) soft phantom, (c,d) stiff phantom, and (e,f) the side-by-side phantom.

To further validate the system’s performance in a non-uniform subject, the side-by-side phantom described above was next tested. The measured spatiotemporal displacement map and the reconstructed Young’s modulus maps are plotted in Fig. 2(c) and Fig. 3(f), respectively. The calculated SWS of the soft part of the side-by-side phantom is consistent with the results obtained for the soft homogenous phantom, while the stiff region also corresponds well to the earlier results. In addition, the location of the inflection point in Fig. 2(c) is consistent with the boundary location in Fig. 3(f).

As indicated by Fig. 3, the imaging field of view (FOV) determined by the maximum depth of the uniformly induced force displacement is over 2 mm based on both soft and stiff phantoms. Therefore, the effective FOV of our elastography imaging system covers the required depth of imaging of the posterior segment eye thickness for our intended subject, ONH and PPS.

B. ONH and PPS response to intraocular pressure

Figure 4 shows the 2D cross-sectional B-mode images and corresponding SWE images of the posterior segment of a porcine eye. A total of five different IOP levels were investigated, including one low IOP (6 mmHg), two normal physiological IOP (12 mmHg and 18 mmHg) and two high IOP (24 mmHg and 30 mmHg).

Fig. 4.

Fig. 4.

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The B-mode image and the reconstructed Young’s modulus map of the posterior pole of the porcine eye including ONH and PPS under five different IOPs − 6 mmHg, 12 mmHg, 18 mmHg, 24 mmHg and 30 mmHg. The second column shows the posterior pole biomechanics in an increased IOP while the third column shows that in a decreased IOP.

The differences in the appearance of the B-mode images obtained at each IOP level are barely discernable (i.e., subtle changes in thickness or curvature). In other words, it is difficult to provide quantitative measurements by relying solely on the B-mode images. In examining Fig. 4 where the color-coded Young’s modulus in the SWE images represents stiffer regions in red and softer areas in blue, the intensity of SWE imaging increases rapidly with IOPs, which indicated that SWE imaging is far more sensitive to changes in IOP than gray-scale B-mode. In addition, both ONH and PPS have a tendency to become stiffer at a higher IOP. This change in elasticity was noted to be more prominent in the PPS than it was in the ONH.

C. Statistical analysis

To statistically determine the relationship between the Young’s modulus of ONH / PPS with increasing / decreasing IOP, eight porcine eyeballs were used. For each eyeball, we calculated one value for ONH and one for PPS. More specifically, the average Young’s modulus of the central ONH contributed to one value. The left PPS and right PPS were averaged to represent the overall Young’s modulus for the PPS. The reconstructed Young’s modulus values were all expressed as mean ± standard deviation (eight eyes).

With increasing IOP, the estimated Young’s modulus values for the PPS were 176.8 ± 14.3 kPa at 6 mmHg, 201.3 ± 13.2 kPa at 12 mmHg, 230.6 ± 16.3 kPa at 18 mmHg, 451.5 ± 100.5 kPa at 24 mmHg, and 573.5 ± 64.4 kPa at 30 mmHg, respectively. With respect to the ONH, the corresponding reconstructed Young’s modulus values were 92.4 ± 13.9 kPa at 6 mmHg, 112.7 ± 18.6 kPa at 12 mmHg, 134.9 ± 18 kPa at 18 mmHg, 158.3 ± 22.8 kPa at 24 mmHg, and 224.7 ± 71.1 kPa at 30 mmHg, respectively. It is generally observed that PPS is stiffer than ONH and the stiffness of PPS increases more rapidly than that of ONH.

For decreasing IOP, the reconstructed Young’s modulus of PPS and ONH were 662 ± 78 kPa / 344.7 ± 58 kPa at 30 mmHg, 480 ± 12.7 kPa / 253.2 ± 2.9 kPa at 24 mmHg, 392.4 ± 27.5 kPa / 244.6 ± 23.3 kPa at 18 mmHg, 297.5 ± 17 kPa / 198.7 ± 27.3 kPa at 12 mmHg, 237.3 ± 5.6 kPa / 157.9 ± 13.6 kPa at 6 mmHg.

The differences between the various IOP levels were evaluated by one-way ANOVA. The statistical analysis showed that changes in Young’s modulus were statistically significant at different IOP values at either increased or decreased IOP. In this study, P < 0.05 was considered to be a significant difference. In Figure 5, the averaged Young’s modulus of ONH and PPS at different IOP levels are plotted. Due to the large tissue deformation variation and non-linearity, we fit the data using a 2nd order polynomial function as opposed to a linear fitting function.

Fig. 5.

Fig. 5.

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The relationship between the biomechanics (Young’s modulus) of the ONH and PPS under different IOPs.

IV. Discussion

We have presented the mapping of stiffness for both ONH and PPS for increasing and decreasing IOP using a shaker-induced ultrasonic elastography system. We used a high frequency and broad bandwidth needle transducer, and therefore our system has the ability to detect micro-level axial displacement and provide imaging mapping with high spatial resolution. The high spatial resolution of our system enables the acquisition of accurate spatial-temporal maps for subsequent post-processing. The deep penetration depth of the ultrasound signal covers the whole posterior segment of the eye, which enables the elastography system to provide a comprehensive stiffness mapping of the ONH and PPS.

To precisely induce localized shear wave propagation, most previous ultrasound elastography systems or optical coherence elastography systems utilized either an air pulse or acoustic radiation force generated by an ultrasonic transducer as an external pushing source. However, both of these are not suitable in this study, because the air pulse approach has a low bandwidth of induced mechanical waves and slow relaxation times [25], while there is strict FDA regulation of ultrasound in ophthalmology (i.e., to induce a detectable deformation on ONH and PPS, a large power is essentially required to be applied on the ultrasound transducer, resulting in a significantly higher mechanical index and acoustic intensity). Therefore, a mechanical shaker which has previously been investigated on human subjects clinically [23] was preferred in this study to vibrate the tissue to initiate elastic wave propagation. The effect of different pushing duration was discussed in our previous work [26]. In this study, 500 μs pulse duration was used as well to ensure a long propagation distance of the induced elastic wave under the transient pushing.

Two homogeneous tissue mimicking phantoms and a homogeneous side-by-side phantom were first used to validate the feasibility of our imaging system. By applying linear regression to a mapping of the spatiotemporal positions, Young’s modulus values at each spatial domain were reconstructed. The reconstructed Young’s modulus values via shear wave speed were comparable and matched the gold standard mechanical test. All these results demonstrate that the shaker-induced ultrasonic elastography system can provide reliable measurement of the elasticity distribution of the imaged subject, which is in a similar geometry in ONH and PPS (i.e., side-by-side structure).

Investigation of biomechanical properties such as Young’s modulus of ONH and PPS with elevated IOP has been conducted by other groups through tensile testing [8] and by using an AFM [10]. More specifically, a range from 10 to 36 MPa and ~ 17 kPa stiffness level have been reported, respectively. It should be noted that both ONH and PPS need to be cut into stripes for the tensile test, and micro-meter level specimens are required in AFM testing. In other words, the difference of geometry in these studies results in the large range of the measured Young’s modulus. To better describe the stiffness mapping of the ONH and PPS, non-invasive imaging of the target region of interest (ROI) in an intact eye structure is preferred. Pavlatos et al. [21] and Ma et al. [22] implemented an ultrasound speckle tracking method to measure the deformation of the posterior segment of porcine and human eyes, respectively. However, in their studies, only strain imaging was provided, and they therefore lacked quantitative measurement of the Young’s modulus. By contrast, our previous OCE work on the backside of the porcine eye [29] has shown that the Young’s modulus of ONH ranges from 180 kPa to 450 kPa with an increasing IOP from 10 mmHg to 30 mmHg. Despite the fact that PPS measurements were lacking with our OCE system, the measured Young’s modulus values for the ONH are comparable with the results in this study.

Jan et al. quantified the collagen crimp or micrometer-scale waviness of lamina cribrosa (LC) and PPS with varying IOP [30–32]. Their studies experimentally revealed that collagen fiber waviness decreased with elevated IOP, which in turn caused the fibers to straighten and stretch to become recruited, leading to stiffening of the tissue. Our experimental results which showed that tissue elasticity went up with the increasing of IOP, were consistent with those previous studies. In addition, our findings indicated that both ONH and PPS become stiffer when IOP drops from a high pressure. It is possible because that the tissue collagen crimp does not return to its original form when IOP returns to its initial level following elevation. As a consequence, the overall posterior segment of the eye at decreased IOP becomes stiffer than that of increased IOP.

Moreover, collagen fibers uncramp when stretched, which implies that the posterior ocular tissues are initially compliant at low stretch level but increase rapidly in stiffness at a higher deformation level [33]. The non-linear relationship between ocular tissue stretch and IOP elevation has previously been observed experimentally in the posterior sclera and ONH [34, 35]. In this study, we successfully demonstrated the trend of Young’s modulus of ONH and PPS with varying IOPs, and confirmed its non-linear properties. However, the 2nd order polynomial function fitting used in this study might not be rigorous because of only 8 eyes involved. More samples in the future study would be helpful to develop a more accurate relationship between Young’s modulus of ONH/PPS and IOPs.

There are a few limitations in this study. First, despite the good agreement between the reconstructed Young’s modulus and gold standard in the phantom study, the group shear wave speed (using equation (1) to reconstruct the Young’s modulus) based reconstruction method may not be accurate, and may lead to some bias in the ocular tissue studies because of boundary conditions and the ratio of ONH/PPS thickness to the shear wavelength. In addition, viscosity is another parameter that is crucial for biomechanical characterization of tissues, including ONH and PPS. A detailed analysis of mechanical mode propagation in a bounded medium (i.e., ONH and PPS) is preferred, however, it is beyond the scope of this manuscript. Future studies will develop advanced models to quantify both tissue elasticity and viscosity by evaluating dispersion of shear wave propagation. Second, our study shows that both ONH and PPS tend to be stiffer with increased IOP rather than decreased IOP. This is because the collagen fibers remain straight and therefore lead to local stiffening. Since only 8 porcine eyes were investigated, more samples should be tested in the future in order to obtain a comprehensive understanding of stiffness variations of the ONH and PPS with glaucoma.

V. Conclusions

In summary, we have shown proof of principle of using a shaker-induced ultrasonic elastography system to characterize biomechanical properties of the ONH and PPS quantitatively via the direct indicator: Young’s modulus. The performance of our imaging system was first calibrated in homogeneous, side-by-side gelatin tissue mimicking phantoms, then tested on porcine eyeballs ex vivo. With the IOP elevation, the reconstructed Young’s modulus of ONH raised from 92.4 ± 13.9 kPa at 6 mmHg to 224.7 ± 71.1 kPa at 30 mmHg non-linearly, while PPS showed a similar tendency from 176.8 ± 14.3 kPa at 6 mmHg to 573.5 ± 64.4 kPa at 30 mmHg. It is generally observed that at the same IOP level, the posterior segment of the eye at decreased IOP is stiffer than that of increased IOP. Overall, simultaneously investigating the link between the biomechanical properties of the ONH / PPS with both increased and decreased IOP may provide new insights to better understand the progression of the glaucoma. The proposed ultrasonic elastography method might be a powerful tool to assess both morphological and biomechanical properties of ocular tissue.

  • To investigate the biomechanical properties of the posterior eye under different intraocular pressures.

  • High frequency ultrasound elastography was used for experiments.

  • The performance of the developed elastography system was validated on different types of phantoms and then tested on porcine eyeball ex vivo.

  • The stiffness relationship between optic nerve head and peripapillary sclera may provide better understanding of the Glaucoma.

VI. ACKNOWLEDGEMENT

The authors acknowledge the help of Mr. Robert Wodnicki for manuscript revision and editing.

This work was supported by the National Institutes of Health (NIH) under grant R01EY026091, R01EY028662, R01EY030126 and NIH P30EY029220. Unrestricted departmental grant from research to prevent blindness.

Footnotes

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