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**|**Biomed Opt Express**|**v.3(5); 2012 May 1**|**PMC3342182

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- Abstract
- 1. Introduction
- 2. Principle and tool design
- 3. Results
- 4. Discussion
- 5. Conclusion
- References and links

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Biomed Opt Express. 2012 May 1; 3(5): 1062–1076.

Published online 2012 April 19. doi: 10.1364/BOE.3.001062

PMCID: PMC3342182

Xuan Liu,^{1,}^{*} Iulian I. Iordachita,^{2} Xingchi He,^{3} Russell H. Taylor,^{2} and Jin U. Kang^{1}

Received 2012 March 27; Revised 2012 April 16; Accepted 2012 April 16.

Copyright ©2012 Optical Society of America

This is an open-access article distributed under the terms of the Creative Commons Attribution-Noncommercial-No Derivative Works 3.0 Unported License, which permits download and redistribution, provided that the original work is properly cited. This license restricts the article from being modified or used commercially.

This article has been cited by other articles in PMC.

During vitreoretinal surgery, the surgeon manipulates retinal tissue with tool-to-tissue interaction forces below the human sensory threshold. A force sensor (FS) integrated with conventional surgical tools may significantly improve the surgery outcome by providing tactile feedback to the surgeon. We designed and built a surgical tool integrated with a miniature FS with an outer diameter smaller than 1 mm for vitreoretinal surgery based on low-coherence Fabry–Pérot (FP) interferometry. The force sensing elements are located at the tool tip which is in direct contact with tissue during surgery and the FP cavity length is interrogated by a fiber-optic common-path phase-sensitive optical coherence tomography (OCT) system. We have calibrated the FS's response to axial and lateral forces and conducted experiments to verify that our FS can simultaneously measure both axial and lateral force components.

Microsurgery involves the manipulation of very small anatomical structures. It is a series of manual manipulations performed predominantly using visual feedback, typically with the aid of a surgical microscope or other magnification device. In addition to visual feedback, tactile feedback can improve surgical outcome significantly. However, in microsurgical procedures such as vitreoretinal surgery, tool-to-tissue interaction forces are usually below human perception thresholds. Gupta et al. reported that 75% of tool-to-tissue interaction forces in vitreoretinal surgery were less than 7.5 mN in magnitude and that only 19% of force events at this magnitude were felt by the surgeon [1]. Although various tactile feedback or “sensory substitution” schemes (e.g., [2–6]) may be used to provide feedback to the surgeon about these forces to help prevent potential tissue damage in freehand or robotic assisted surgery, a suitable force sensor (FS) must be incorporated into the surgical tools. In vitreoretinal surgery, the interaction forces between the tool and the sclera (white part) of the eye are often equal to or larger than the forces between the tool tip and the tissue being manipulated [7,8]. Thus, in addition to high sensitivity in force measurement, vitreoretinal microsurgery also requires using small size sensors that can be integrated into sub-millimetric instrument shafts. These considerations make the design of suitable force sensors for vitreoretinal microsurgery very challenging.

Previously, a MEMS technique and miniature strain gauge have been implemented in the development of micro force sensing tools for retinal microsurgery, such as 1 degree of freedom (DOF) pick-like FS developed by Gupta et al. [1], multi-axis MEMS FS developed by Kim et al. [9], microgripper-integrated FS developed by Menciassi et al. [10], 3 DOF FS developed by Berkelman et al. [11,12] which had been incorporated later into a hand-held instrument and used to measure forces in retinal tasks [13]. Most of the above-mentioned sensors, with the exception of that of Gupta et al. [1], are too large to be incorporated within a surgical tool that will be placed inside the eye during vitreoretinal microsurgery.

Force sensors based on fiber optics can be very small, mechanically stable, immune to electrical noise, and sterilizable [14]. Therefore they are ideal for incorporating into conventional surgical tools and for use as sensory substitution for tactile feedback in vitreoretinal microsurgery. Various force sensing schemes based on fiber optic techniques have been investigated over the last several decades. For example, FS that measures the change of optical power due to beam deflection has been developed for minimally invasive surgery [15,16]. However, such FS cannot be directly used in vitreoretinal microsurgery due to the large diameter (about several to ten millimeters) and low force sensitivity (several to tens of millinewtons). In our previous work [8], we incorporated fiber Bragg grating (FBG) sensors into the shafts of microsurgical instruments. Although this approach produces very high force sensitivity in directions that are perpendicular to the tool shaft (lateral directions), the axial rigidity of both the tool and fibers makes achieving high axial force sensitivity with FBG sensors a challenging problem. Fabry–Pérot interferometer (FPI) type sensors have also been used in bio-applications, including surgical force sensing [17,18]. The immediate advantage of FPI type micro force sensing is that it can potentially achieve a better axial sensitivity than that of FBG FS.

In this study, we designed and built a fiber optic FS consisting of Fabry–Pérot interferometer (FPI) cavities at the distal end of a fiber probe for potential applications in vitreoretinal surgery. In our FPI-FS, the length of FP cavity changed with force and we used a common-path Fourier domain optical coherence tomography (CP-FD-OCT) to interrogate the change of cavity length by phase-sensitive detection. The simple and compact common-path OCT configuration not only allows the use of optical probes with arbitrary lengths [19,20], but also provides improved phase stability to achieve sub-nanometer sensitivity in displacement measurement [21] and therefore to achieve high force sensitivity. Moreover, OCT also facilitates simultaneous measurement of multiple FP cavity lengths [22]. In our FS, we multiplexed three FPIs together to simultaneously obtain signals from three different channels. This enabled us to measure force in three dimensions with components both along and perpendicular to the tool shaft, which, to the best of our knowledge, has not been reported before.

In this manuscript, we present the designing and fabrication of our FPI-FS (Section 2). Principle of phase-sensitive detection and multidimensional force measurement are presented in Section 3. In Section 4, we summarize the results of tool calibration and show experimental results of measured force. The final sections are discussion and conclusion.

Figure 1(a) shows the system schematic of FPI-FS for vitreoretinal microsurgery. The FP cavities that sense the forces are located at the tip of a conventional surgical tool—a surgical pick in this case—which is inserted into an eye to perform surgical tasks during vitreoretinal surgery. Detail of the FPI-FS probe is shown on the left side of Fig. 1(b). A CP-FD-OCT system is used to interrogate the length of the cavities by phase-sensitive detection. More details of our CP-FD-OCT system can be found in our previous work [23]. In brief, as shown on the right side of Fig. 1(b), it is a spectral-domain OCT system operating in 800 nm wavelength range with a ~6μm axial resolution and a ~3 mm imaging depth. The real-time 1-kHz data acquisition and processing was implemented with C++ in a Dell Precision T7400 computer.

(a) System schematic of FPI FS; (b) detailed schematic of the system including FPI-FS and CP OCT interrogation; (c) cross section of tool shaft with fiber embedded; (d) CAD model for the Nitinol flexure; (e) CAD model of the FPI-FS probe without the tool **...**

In our FPI-FS system, a superluminescent emission diode (SLED, Exalos Inc.) is used as broadband source (BS) to illuminate three FPIs which consist of a shared end mirror and three lead-in single-mode fibers (SM800-5.6-125, Thorlabs, 125 μm cladding diameter). The mirror, which serves as the end reflector of the FPIs, was fabricated by polishing the surface of a stainless steel wire having 0.5mm diameter and 2.5mm length. We cleaved the fiber tips to right angle so the light is partially reflected at the fiber tip. Optical fields reflected by the fiber tips (*E*_{1}* _{i}*,

As depicted in Fig. 1(b), the force-sensing part of the tool contained three optical fibers and a mirror, which form three FPIs with different cavity lengths. Other components of the FS include a tool shaft, a flexure, and a tool tip (pick). The tool shaft is made of a piece of titanium wire that is strong, lightweight, and biocompatible. The wire is 0.5 mm in diameter and 50 mm in length. Three square section channels (0.15 mm × 0.15 mm) were machined (laser cut) into the surface of the wire along the tool axis, 2π/3 from one another in a circle, as shown in Fig. 1(c). Fibers were embedded and fixed into the channels.

The mirror and tool shaft are bound by a flexure made of Nitinol using a laser cut technique. The flexure deforms proportionally to the force applied and therefore induces change in the lengths of the FP cavities that will be used to calculate the force. We chose to use Nitinol because of its super-elastic property which enables the tool to have a measurable deformation with a force in the order of a millinewton. Nitinol has modulus of elasticity ranging from 41 × 10^{3} MPa to 75 × 10^{3} MPa; therefore, the fabricated tool has extremely high force sensitivity. As shown in Fig. 1(d), the flexure—with 0.80mm outer diameter (OD) and 0.60mm inner diameter (ID)—has five layers; each layer consists of three 40-micrometer struts; and between adjacent layers there is a 100 micrometer slot. The multi-layered configuration can further increase the elasticity of the flexure in the axial direction and therefore achieve a high sensitivity for force measurement. The total length of the flexure is 3 mm, with 1 mm between the distal end and the first slot. The pick was made using a stainless steel SS304 hypodermic tube, 21 gauge (OD = 0.80mm, ID = 0.55mm). The tip was bevel cut (at 75 degrees) and bent at around 45 degrees relative to the axis. The pick total length is 2.5mm, with a 1.5mm length bevel section. Loctite medical instant adhesive (Henkel, Dusseldorf, Germany) was used to bond all the parts together: first the mirror with the pick, next the flexure, and finally the tool shaft. The optical fibers were embedded into slots machined to the tool shaft. The distance between the fiber tips and the polished metal surface was manually adjusted before finally bonding the fibers with the tool shaft. It is worth mentioning that we made the FP cavities to have different lengths. This allows us to detect the length change of each cavity respectively. The CAD model for the FPI-FS probe integrated with the pick is shown in Fig. 1(e). For calibration and future manual or robotic use, the force sensing probe is attached to a custom-made handle, as shown in Fig. 1 (f). Figure 1(g) is a photo of the tool pictured below a U.S. quarter.

Assuming that the polished metal surface reflects all the incident light and that the fiber tip reflectivity, *r,* is about 4% according to Fresnel equation [24], we calculated the finesse of the FPI to be about 1.25. The round trip loss of optical power in the FP cavity is larger than 4% due to beam divergence; therefore, the actual finesse of FPIs are even lower than 1.25. Moreover, we have measured the fringe visibility of interferograms obtained from the FPIs and the values are very small (0.39, 0.047 and 0.049, for three FPIs, respectively). Therefore, in our multiplexed FPIs, signals due to higher order reflections are small and therefore can be neglected in the following analysis.

As shown in Fig. 1(b), we denote the optical field reflected by the tip of the *i*^{th} fiber as E_{1}* _{i}*. We denote the optical field reflected by the shared end mirror and collected by the

$${L}_{i}={L}_{i,0}+\delta {l}_{i}\left(F\right).$$

(1)

In Eq. (1), *L _{i,0}* is the length of

Since the round trip optical path length traveled within the *i*^{th} fiber, *l _{i},* varies significantly from other fibers (much more than several millimeters), E

$$S\left(k\right)=\eta \left[{\displaystyle \sum _{i=1}^{3}{\left|{E}_{1i}\left(k\right)\right|}^{2}}+{\displaystyle \sum _{i=1}^{3}{\left|{E}_{2i}\left(k\right)\right|}^{2}+}{\displaystyle \sum _{i=1}^{3}2\mathrm{Re}\left({E}_{1i}\left(k\right){E}_{2i}{}^{*}\left(k\right)\right)}\right]$$

(2)

In Eq. (2), *η* indicates the responsive coefficient of the system; * denotes complex conjugate; Re() refers to the real part of a complex signal. The third term of the right-hand side of Eq. (2) is the interference term. OCT signal, *I*_{OCT}(*z*), can thus be obtained by performing inverse Fourier transform on *S*(*k*):

$${I}_{OCT}\left(z\right)={F}^{-1}\left[S\left(k\right)\right]$$

(3)

Denoting the axial point spread function of the OCT system as *h*(*z*), which equals *F*^{−1}(|*E*_{0}(*k*)|^{2}), we can rewrite Eq. (3) as Eq. (4) with approximation when the bandwidth is relatively small:

$$\begin{array}{l}{I}_{OCT}\left(z\right)=\text{}\eta h\left(z\right){\displaystyle \sum _{i=1}^{3}\left[{\left|r{a}_{i}\right|}^{2}+{\left|{r}_{2i}{a}_{i}\right|}^{2}\right]}\\ \text{}+\eta {\displaystyle \sum _{i=1}^{3}\left[r{r}_{2i}^{*}{a}_{i}^{2}h\left(z-2{L}_{i,0}\right){e}^{j2{k}_{0}\delta {l}_{i}}\right]}\\ \text{}+\eta {\displaystyle \sum _{i=1}^{3}\left[{r}^{*}{r}_{2i}{a}_{i}^{2}h\left(z+2{L}_{i,0}\right){e}^{-j2{k}_{0}\delta {l}_{i}}\right]}\end{array}$$

(4)

The first term on the right-hand-side (RHS) of Eq. (4) is the auto-correlation term. The second and third terms in the RHS of Eq. (4) represent coherence peaks due to the interference between two end reflectors of each FPI. The second and third term in the RHS of Eq. (4) are symmetry about the zero delay because the Fourier transform is performed on real spectral data.

As indicated by Eq. (4), the change of FPI cavity length δ*l _{i}* can be extracted from

$$2{k}_{0}\delta {l}_{i}={\mathrm{tan}}^{-1}\left\{\frac{\mathrm{Im}\left[{I}_{OCT}\left({L}_{i,0}\right)\right]}{\mathrm{Re}\left[{I}_{OCT}\left({L}_{i,0}\right)\right]}\right\}-{N}_{i}\pi $$

(5)

$$\delta {l}_{i}=\frac{{\lambda}_{0}}{4\pi}{\mathrm{tan}}^{-1}\left\{\frac{\mathrm{Im}\left[{I}_{OCT}\left({L}_{i,0}\right)\right]}{\mathrm{Re}\left[{I}_{OCT}\left({L}_{i,0}\right)\right]}\right\}-{N}_{i}\frac{{\lambda}_{0}}{4}$$

(6)

In Eqs. (5) and (6), *N _{i}* is the integer that makes (δ

$$\delta {l}_{i}\left(F\right)={d}_{i}-{d}_{i,0}$$

(7)

When the variation of FP cavity length is large—such that *ϕ _{i}* jumps from -π/2 directly to π/2 or from π/2 directly to -π/2—a phase unwrapping technique can be used to obtain continuous phase or displacement [21].

When a small force is applied to the tool, the change in FP cavity length is extremely small (in the order of a nanometer) compared to the axial resolution of our OCT system—even when the super-elastic material Nitinol is used to fabricate the flexure. Therefore it cannot induce any observable change in the amplitude of OCT signal. Nevertheless, the small displacements can lead to a measurable phase shift of the interference spectrum. In other words, the small deformation can be extracted from the phase of OCT signal at the coherence peaks as shown in Eq. (6), which is demonstrated in the following simulation. In this simulation, we assumed *L*_{1,0} = 100μm, *L*_{2,0} = 150μm, and *L*_{3,0} = 200μm; the light source has central wavelength of 800nm and FWHM bandwidth of 60nm. Superposed spectral interference signals from three FP cavities are simulated and shown in Fig. 2(a)
according to Eq. (2). Performing inverse Fourier transform on this interference signal leads to complex-valued spatial domain OCT signal *I*_{OCT}(*z*), which contains three coherence peaks at *z* = *L*_{1}, *L*_{2}, and *L*_{3} corresponding to the length of the three FP cavities. The amplitude of OCT signal is shown in Fig. 2(b). With force exerted, the cavity length changed proportionally to the force. In this simulation, we assumed *δl*_{1} = 20nm, *δl*_{2} = 40nm and *δl*_{3} = 60nm. However, such small change in cavity length cannot be seen directly from the amplitude of OCT signal described in Fig. 2(b). On the other hand, as shown in Fig. 2(c) which is the central part of the interference spectrum, change in cavity length in the order of nanometer shifts the original spectrum (black) to the red one. To demonstrate that the small displacement can induce measurable phase change to the complex-valued OCT signal at the coherence peak, we show complex OCT signals at the 1st, 2nd, and 3rd coherence peaks in the complex plane using a polar coordinate system in Fig. 2(d). Black symbols and red symbols represent OCT signals with and without additional displacement *δl*_{i}. Clearly, the vector that represents the complex OCT signal rotates due to phase shift induced by *δl*_{i}.

(a) Superposition of spectral interference signals from three FP cavities; (b) OCT signal with three coherence peaks corresponding to three FP cavities; (c) central part of the interference spectrum: with (red) and without (black) additional displacement **...**

In Fig. 2, we also show the actual spectrum obtained from the multiplexed FPIs of our FS as Fig. 2(e). We performed Fourier transform on the interferogram and obtained OCT signal shown in Fig. 2(f). Three peaks in Fig. 2(f) come from interference between end reflectors of the three FP cavities and the locations of these peaks indicate the length of FP cavities.

As demonstrated in the three-dimensional Cartesian coordinate of Fig. 1(b), a force (*F*) along an arbitrary direction can be decomposed into axial and lateral components. In the force coordinate attached to the tool tip, *z* axis or axial direction is along the tool shaft. We denote force along the tool shaft direction as axial force *F _{z}*.

As shown in Figs. 3(a)
and 3(b), with an axial force applied to the tool, the flexure deforms and the change of FP cavity length is denoted as *δl*_{i,z} which equals (*F _{z}D*)/(

(a) FS at neutral; (b) axial force changes the length of flexure and changes the FP cavity length; (c) lateral force leads flexure to bend and changes the FP cavity length; (d) diagram shows the calculation of cavity length induced by lateral force; (e) **...**

According to Fig. 3(d), the change of FP cavity length *δl*_{i,}* _{l}* due to lateral force can be expressed as [-2

$${b}_{i}=\sqrt{{x}_{i}^{2}+{y}_{i}^{2}}\mathrm{cos}\left(\theta +{\epsilon}_{i}\right)$$

(8)

In Eq. (8), *ε _{i}* = tan

$$\begin{array}{l}\delta {l}_{i}=\frac{{F}_{z}D}{{A}_{0}E}+\sqrt{{x}_{i}^{2}+{y}_{i}^{2}}\mathrm{cos}\left(\theta +\epsilon \right)\frac{{F}_{l}{D}^{2}}{2EI}\\ \text{}=\frac{D}{{A}_{0}E}{F}_{z}+{x}_{i}\frac{{D}^{2}}{2EI}{F}_{l}\mathrm{cos}\theta -{y}_{i}\frac{{D}^{2}}{2EI}{F}_{l}\mathrm{sin}\theta \end{array}$$

(9)

It is clear in Eq. (9) that the displacement induced by lateral force has a sinusoidal dependency on θ.

Denoting *F _{l}*sinθ and

$$\delta {l}_{i}={A}_{ix}{F}_{x}+{A}_{iy}{F}_{y}+{A}_{iz}{F}_{z}$$

(10)

After introducing *F _{x}* and

Although for a given amplitude of lateral force *F _{l}*,

$${F}_{l}=\sqrt{{F}_{x}^{2}+{F}_{y}^{2}}$$

(11)

With Eqs. (7) and (10), the relationship between *δl _{i}* and

$$\delta l=AF$$

(12)

$$\delta l=\left[\begin{array}{l}{d}_{1}-{d}_{01}\\ {d}_{2}-{d}_{02}\\ {d}_{3}-{d}_{03}\end{array}\right];\text{A=[A1xA1yA1zA2xA2yA2zA3xA3yA3z];F=[FxFyFz].}$$

Values of elements in matrix ** A** can be obtained from a calibration procedure shown in detail in the following section. Once

$$F={A}^{-1}\delta l$$

(13)

As shown in Eq. (12) and Eq. (13), it is critical to subtract *d*_{0}* _{i}* from the displacement

In our axial calibration experiment as illustrated in Figs. 4(a)
and 4(b), we attached different numbers of testing weights to the FS and maintained the tool shaft in the gravity direction to keep *F _{x}* and

(a) Photo of axial calibration setup; (b) schematic of axial calibration setup; (c) displacement versus axial force (first FP cavity); (d) displacement versus axial force (second FP cavity); (e) displacement versus axial force (third FP cavity).

Using the complex-valued OCT signal acquired from our CP-FD-OCT, we calculated the displacements at the three coherence peaks from the phase of OCT signal with Eq. (6) using *N _{i}* = 0. The measured displacements are equal to change in the length of FP cavities. With different numbers of testing weights i.e. axial forces, we obtained

In the lateral calibration setup as shown in Figs. 5(a)
and 5(b), we maintained the tool shaft (*z* axis of the force coordinate) perpendicular to the gravity direction so that the force induced by the gravity of the testing weights had only a lateral component in the force coordinate that was attached to the tool tip; in other words, *F _{z}* = 0. As indicated by Eq. (9), the displacement induced by lateral force not only depends on the magnitude of lateral force exerted to the FS, but also depends on θ, which essentially is the direction of lateral force in

(a) Photo of lateral calibration setup; (b) schematic of lateral calibration setup; (c)–(h) results from lateral calibration: (c)–(e) show displacements measured from three FP cavities at different θ with different lateral forces. **...**

In our lateral calibration experiments, we attached testing weights to the tool; we rotated the tool circumferentially to obtain different values of θ and took the displacement measurements. We repeated the above procedures with different lateral forces by attaching different numbers of testing weights to the FS. The experimental results of our lateral calibration are shown in Figs. 5(c)–5(e).

Displacements from the 1st, 2nd, and 3rd FP cavities are indicated by circles in Figs. 5(c)–5(e), in which data points with different colors were obtained with different lateral forces (numbers of testing weights are shown as legends in these figures). Figures 5(c)–5(e) clearly indicate that the displacements induced by lateral force vary sinusoidally as the tool orientation θ for a given lateral force, which is consistent with Eq. (9). Sinusoidal fittings of the experimental results are also shown as dashed curves in Figs. 5(c)–5(e) to further verify the functional dependency of displacement on θ. In addition, results in Figs. 5(c)–5(e) also show that the phase of the displacement's sinusoidal dependency on θ varies for different FP cavities. This is because the lead-in fiber for the *i*^{th} FP cavity takes a different transversal coordinate (*x _{i}*,

To obtain *A _{ix}* and

With *A** _{l}* obtained from linear regression of the lateral calibration data, we could calculate

In conclusion, we obtained the following coefficients as elements of our calibration matrix ** A**:

To validate that our FS can measure force with both axial and lateral components, we conducted the following experiment. In addition to the rotary stage in *x*-*y* plane, another rotary stage was used to change , the intersection angle between the tool shaft and the gravity direction, as shown in Figs. 7(a)
and 7(b). With the same load attached to the FS, different values of leads to different values of axial and lateral force. Here *G* is the gravity of the load which equals 8.2mN for this experiment. We used a rotation stage (as indicated by Rotation 1 in Figs. 7(a) and 7(b)) to change and therefore change *F _{x}*,

(a) Photo of 3D force measurement setup; (b) Schematic of 3D force measurement setup; (c) Measured and actual axial force; (d) Measured and actual lateral force; (e) Measured and actual axial force obtained by using the known load gravity as bias; (f) **...**

$${F}_{x}=G\mathrm{sin}\mathrm{\mathrm{cos}\theta ;{F}_{y}=G\mathrm{sin}\mathrm{\mathrm{sin}\theta ;{F}_{z}=G\mathrm{cos}}}$$

(14)

We repeated the measurements with θ = π/2 and θ = π. Axial and lateral forces obtained are shown as circles in Figs. 7(c) and 7(d), respectively. Red circles indicate data obtained with θ = π/2; blue circles indicate data obtained with θ = π.

It is worth mentioning that we biased the displacement reading when = π/2 and the FS was exerted with a non-zero force. As shown in Eq. (7) and discussed at the end of Section 2.3, the forces measured were values relative to the force when biasing was taken. As a result, the measured forces are *F _{x,m}* =

Measuring a force in an arbitrary direction with respect to the tool shaft of the FS requires at least three FP cavities to provide independent displacement readings. First, the change of FPI cavity length is a linear superposition effect of axial and lateral forces. As indicated by our calibration experiments, the same magnitude of lateral force can induce much larger change in cavity length compared to axial force. In other words, our FS's response to axial and lateral forces are very different. Therefore, it is necessary to differentiate *F _{z}* and

As discussed in Section 2.3, we can directly calculate our FS's response to axial and lateral forces using the mechanical and geometric properties of our tool. However, the mechanical property of flexure might change significantly because Nitinol changes its mechanical property during laser cutting procedures. Moreover, we applied medical adhesive to glue different components together, which would also change the overall mechanical properties of the tool. Besides, due to the extremely small dimension of the tool, the geometry of the fabricated tool might be different from the initial design. As a result, experimental calibration can characterize the performance of our FS much better than a theoretical model or even finite element analysis.

Our calibration results show that different FP cavities have different responses to axial force (*A*_{1}* _{z}* = 4.3nm/mN;

The force measurement sensitivity depends on the sensitivity of phase-resolved displacement measurement as well as ** A**, coefficients relating cavity length change, and force. According to Eq. (13), the variance of force

As shown in Eq. (5) and Eq. (6), we extract displacement from the phase of a complex number that is defined in the domain [-π/2 π/2]. When a time-varying phase reaches and crosses the boundary of this range, the phase calculated from Eqs. (5) and (6) will show an abrupt change of ± π, depending on whether the phase decreases or increases. According to Eq. (7), the value of *d _{i}* is from −100nm to +100nm and therefore we can obtain displacement without ambiguity only in a range of 200nm. Since our FS has 30 ~40 nm/mN sensitivity to lateral force, a 200nm displacement range allows us to measure 5mN ~6mN lateral force at most. Therefore, a severe restriction in the dynamic range of our force measurement is placed by the limited range of displacement measurement. However, we implemented a well-established method called

In this study, we proposed and studied a miniature fiber-optic force sensor based on low coherence Fabry-Pérot interferometry for vitreoretinal surgery. We used a common-path Fourier domain optical coherence tomography to interrogate the change of FP cavity length through phase-sensitive detection. We fabricated a force sensing tool with an outer diameter less than 1 mm. Calibration was performed to characterize the tool. We also conducted experiments to verify that our FS can measure force with both axial and lateral components.

The research reported in this paper was supported in part by NIH BRP grant 1 R01 EB 007969 and in part by Johns Hopkins University internal funds. Hardware, software, and systems developed within the Engineering Research Center for Computer-Integrated Surgical Systems and Technology (CISST ERC) under NSF cooperative agreement EC9731748 provides significant infrastructure in support of this work.

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