Open Access
Add to BibSonomy
Abstract
Biopsy needles with embedded force sensors can eliminate the needle deflection and the needle targeting failure risks during MRI guided biopsy procedures. Fabry-Pérot interferometry (FPI) based sensors are small, compact and immune to electromagnetic and RF interferences, and therefore they are suitable for needle guidance under MRI. In this work, an FPI based fiber optic force sensor design and its integration to an 18-gauge MRI compatible biopsy needle are presented. The custom designed FPI sensor provides a force measurement range up to 13 N with a resolution of 0.1 N through benchtop experiments. The MRI compatibility of the sensor was evaluated using a commercially available prostate phantom under MRI.
© 2018 Optical Society of America under the terms of the OSA Open Access Publishing Agreement
1. Introduction
Magnetic Resonance Imaging (MRI) has a great potential to replace the existing imaging modalities during interventional procedures such as biopsies thanks to its superior soft tissue contrast and ionizing radiation-free nature. Accurate measurement of the force acting on a biopsy needle tip during MRI-guided biopsy procedures may provide important feedback that can increase the safety and accuracy of these procedures [1–4]. A real-time force measurement capability from the needle distal tip during the biopsy operation can allow physicians to appreciate small variations in the mechanical properties of the tissues along the needle insertion trajectory. This valuable information can be used to improve needle targeting and analyze the tissue-needle interaction. In conventional biopsy procedures, the accuracy and safety of the operation depend highly on the operator’s experience. One of the major problems in a biopsy operation is the needle deflection during insertion. The needle deflection can be detected immediately through sudden fluctuations in continuous force measurement if a biopsy needle has an embedded force sensor. It is not possible to use conventional electrical force sensors for MRI guided interventions, because of potential electromagnetic interferences during imaging. Fiber optic force sensors can be used under MRI without causing any danger or any disruption on the MR image. Applied axial force measurement during needle guidance can be performed by Fabry-Pérot Interferometry (FPI) based fiber optic force sensors which can be integrated to the biopsy needle tip.
FPI based fiber optic force sensors work based on light interference formed by superimposition of light beams reflected from two semi reflective mirror surfaces that generally form an air cavity between them (Fig. 1). Reflected light beams form an interference pattern based on a dynamically changing optical path difference. Mirror surfaces are usually formed by cleaved optical fibers which are fixed inside of a capillary tube. Variations in the cavity length can be interpreted based on the phase shift between reflected light beams which occurs due to the change of optical path difference caused by an applied axial force.
Fiber optic force sensors are preferable for having force feedback during needle insertion because their small profile allows better integration into the target medical device and also their RF immunity helps to avoid EM interferences during MRI [5]. The gauge length of the sensor probe which can be defined as the distance between the fixing points of the optical fibers to the encapsulating tube is an important dimension that determines the force measurement range of the sensor. Shang et al. [6] uses fusion welding in two locations for the fixation of optical fibers and glass capillary [Fig. 2(a)], and Singh et al. [7] fixates them using epoxy at the end faces of the glass capillary [Fig. 2(b)]. In this work, FPI sensor housing was formed by creating two micro-holes on the glass capillary by the laser cutting process and then these holes were used as fixation points for optical fibers. This proposed technique has several advantages. First of all, fiber optics can be placed into the glass capillary housing after creating the micro holes so they are free from potential thermal effects of fusion welding on optical fibers. Also, having fixation points closer to the fiber ends provides better control over targeted force range and sensor accuracy. Finally, it allows to align two fibers to maximize the received signal before fixing them to the glass capillary. In certain designs, optical fibers with different diameters also can be fixed into the glass capillary through micro-holes and fiber alignment mismatches can be minimized by this way.
The intensity of successive reflected light beams from FPI cavity drops exponentially depending on the reflectance of the surfaces constructing the cavity. Silica fibers used in the FPI sensor has a reflectance of approximately 4%. According to Fresnel’s equations, the first reflected light beam has an intensity of 4% of the incident light beam intensity. The intensity of the second reflected light beam is 3.69% of that value, and the intensity drops to 0.15% for the third reflected beam [8]. Therefore, the total intensity of light beams reflected from the silica fiber surfaces constructing FPI cavity can be approximated based on the first two of the reflected light beams. The intensity of interfering light beams of FPI can be expressed as follow based on this approximation [Eq. (1)] [9]:
When , the intensity of the interference simplifies to [Eq. (2)]:
where , which is total phase shift in FPI for a round-trip propagation of light beams, which is defined as [Eq. (3)] [10]:
In this expression, is the refractive index of the cavity material, and it equals to 1 in case the cavity material of the FPI sensor is air. The cavity length can be defined as the distance between reflective surfaces. The wavelength of the light beam transmitted to FPI cavity is given as . The initial phase differences between the light beams are expressed with .
Force measurements with FPI sensors strongly depend on the intensity and the wavelength of the light source which might significantly be effected from temperature variations. The rate of charge motions in PN junction, and hence current threshold of a laser diode varies depending on the temperature [11,12]. It is also a known fact that bandwidth and wavelength of the laser source is strongly temperature dependent. Variations in the temperature affect the bandwidth, and also cause shifts in the wavelength of emitted light from the source [13,14]. For accurate force measurements, the temperature of the laser source should be kept constant during the operation.
2. Method
The optical setup of custom design FPI based fiber optic force sensor is given in Fig. 3. The laser source of the system is 8mW laser diode with a wavelength of 635 nm (Thorlabs Pigtail Laser Diode LP635-SF8). The required power for the laser diode is supplied by the laser driver (Thorlabs EK1101 Driver Kit). The laser beam transferred to an optical cage that contains two fiber port collimators (Thorlabs FiberPort Collimator PAF-X-2-B), two photodiodes (Thorlabs PDA36A) and a beam splitter. Transmitted light from the beam splitter is coupled into a single mode optical fiber with a cladding diameter of 125 µm, and transferred to the FPI force sensor located at the biopsy needle tip. Reflected light beams from the FPI force sensor goes back to the optical cage and the beam splitter reflects them towards the photodetector. Finally, the measured light intensity signal by the photodetector is sent to a microprocessor (Texas Instruments MSP430) for signal processing and monitoring.
The temperature of the laser source should be kept constant to have accurate interferometric measurements. A custom designed and dedicated temperature controller was implemented for this purpose using Model Predictive Control (MPC) approach. The temperature control system consists of a custom design Aluminum (Al) housing for the laser source, a thermoelectric cooler (TEC), and a fan for dissipating the heat. Temperature measurements were done by using a digital sensor. (DS18B20, Maxim Integrated). Digital temperature data was transmitted to the electronic board (MSP430F5529 LaunchPad Texas Instruments Inc.), and the recorded data were used as the input for control algorithm created with Matlab MPC toolbox and transferred to the electronic board.
The system was operated with Proportional-Integral-Derivative (PID) and MPC approaches respectively and the performance of each control method was evaluated through benchtop experiments by applying external heat source. The laser housing temperature values were recorded in addition to reference room temperature values.
The custom designed FPI force sensor tip was fabricated using single mode optical fibers, a borosilicate glass capillary and medical grade UV cured adhesives. As seen in Fig. 4, two micro-holes, with a diameter of 100 µm, are created on the borosilicate glass capillary with an inner diameter of 200 µm and an outer diameter of 330 µm by using 60 W CO2 laser engraver (Epilog Legend Series, Epilog Laser). Micro-holes were used as fixation points of the tight-fit optical fibers to the glass capillary. This novel fabrication method provides easier fiber alignment and fixation. Also, it allows to adjust the gauge length of the sensor independent from the total length of the glass capillary. The position of the micro-holes on the glass capillary determines the gauge length of the sensor which plays a critical role in determining both the force range and the sensitivity of the fabricated sensor.
Fig. 4 The two micro-holes created on a borosilicate glass capillary housing by the laser cutting process.
The cavity length formed by the optical fibers fixed into the capillary changes linearly with respect to the amount of axial force applied on the glass capillary as long as the uniformity of the glass capillary of which length is approximately 2 mm is preserved (Fig. 5). The linear relation between the cavity length and the amount of the applied axial force can be given as follows [Eq. (4)] [15–19]:
According to Eq. (4), the cavity length () has a linear relation with the amount of the applied axial force (). Young’s modulus () of the glass capillary is , and Poisson’s ratio is . Inner radius ), outer radius (), and the effective gauge length ) are constant values which are determined during the manufacturing of the FPI sensors based on the specifications of the intended application.
Measuring the applied axial force with a sufficient resolution using a fiber optic sensor requires a sufficient signal to noise ratio (SNR). The alignment and the fixation of optical fibers inside the borosilicate glass capillary are important for this purpose. FPI force sensor components were assembled on an optical table, which provides vibration isolation, using a three-stage manual micro-manipulator and a piezo-electrically driven nano-manipulator stage under an optical microscope. Alignment and fixation of the fibers were completed at the local minima when a sufficiently long linear waveform was acquired. The voltage response based on the intensity of the reflected light from the FPI sensor with respect to the change in cavity length is shown in Fig. 6.
Fig. 6 Sensor response with respect to the cavity length change (left) and a close-up plot between 7 and 10 µm (right). The desired fixation point was marked (solid dot) on the plot.
Custom designed FPI force sensor was integrated into an 18-gauge nitinol biopsy needle. The measurable force range and the sensitivity of the force sensor were evaluated by using a calibrated tension-compression test machine (LF-Plus, Lloyd Instruments Ltd.). The FPI force sensor integrated into the nitinol biopsy needle was placed on the compression testing arm as it is opposed to only axial forces. The measured voltage signal generated by the response of the FPI force sensor against the applied force at the needle tip was recorded simultaneously with the force measurements by the tension-compression test machine. Sensor limitations such as force measurement range and resolution were determined. Final fiber optic force sensor performance was tested by inserting the needle into a beef tissue through benchtop experiments initially. Then, the fiber optic force sensor embedded biopsy needle was tested in a prostate phantom under real-time MRI guidance (3T Magnetom Trio, Siemens GmbH) (Fig. 7).
Fig. 7 The experiment setup of continuous force feedback measurement during biopsy needle insertion into a prostate phantom under MRI (left) and into a beef tissue during benchtop experiment (right).
3. Results
Figure 8 shows the performance of both PID and MPC temperature control systems during sensor operation. The temperature control systems were initiated at time t1, and an excessive heat was introduced to the system at time t2. Figure 8(a) and Fig. 8(b) show the responses of MPC and PID temperature systems respectively. Temperature variations can affect the output power of the laser source or can cause shifts in bandwidth and wavelength of the beam emitted from the laser diode. Therefore, the temperature of the laser diode should be kept constant for accurate force measurements. The MPC approach uses a model to predict future outputs of the system and improves the predicted output in every next step for getting closer to the desired trajectory. It can be seen that MPC outperforms PID in terms of response time and consistency of the response against external effects on the system. The temperature control system operated with the MPC approach has a fast response to reach the target temperature and the target temperature can be maintained for a long time even if there are variations in the ambient temperature as seen in Fig. 8(c).
Fig. 8 (a). The performance of the MPC temperature controller (b). The performance of the PID temperature controller (c). Overall stability of MPC temperature controller.
Figure 9 shows the commercial force sensor’s response (red) and custom design FPI force sensor’s response (blue), against applied increasing axial force with time. The response of the FPI force sensor is consistent with the applied axial force. The measurable force range of the fabricated sensor was determined by the first peak point of the voltage response hits, while the applied axial force is still increasing. The sensor has a force measurement range of 0-13 N. Also, the force sensor provides a quick response to any sudden change in the extent of the applied force.
Fig. 9 Applied force vs. custom FPI based force sensor response. The overall force range of the sensor was marked with dashed line.
The commercial force sensor response and the voltage response of the custom design FPI force sensor to applied axial force were shown in Fig. 10. Close up plots between 13th and 17th seconds were shown below. It can be seen that FPI force sensor can successfully detect the changes in the applied force. Even a small increment of 0.1 N can be detected, which is enough to differentiate different types of tissues during needle insertion.
Fig. 10 Commercial force sensor response (on the left) and custom FPI based force sensor response (on the right) to the applied axial force.
The response of the FPI based force sensor and the commercial force sensor were recorded simultaneously during needle insertion into a beef tissue and a layer of stiff foam (Fig. 11). It can be seen that the response of the two sensors are consistent with each other and FPI based force sensor embedded biopsy needles are capable of detecting penetration and stiffness differentiation.
The fiber optic sensor performance was also tested by performing a biopsy needle insertion experiment into a commercially available prostate phantom under MRI (Fig. 12). It was confirmed that the biopsy needle with an integrated FPI based force sensor is safe to use under MRI and also the fiber optic sensor could detect the transitions between different structures within the phantom. Also, FPI based force sensor embedded needle did not cause any image degradation and it was clearly visible under MRI.
Fig. 12 FPI based force sensor embedded needle insertion into a commercially available prostate phantom under MRI (left), the force response of the sensor (right).
4. Discussion
Performing biopsy procedures under MRI is a promising alternative method thanks to the superior soft tissue contrast of MRI. Real time axial force feedback at the biopsy needle tip during insertion increases the accuracy of the lesion targeting and also enhances the patient safety by determining possible needle deflections during the biopsy procedure.
The force measurement range and the resolution of custom designed FPI sensor is sufficient to perform safe biopsy operations, compared to the needle and soft tissue interaction studies in the literature [20–22]. The sensor resolution can be improved by reducing SNR losses due to misalignments of optical fibers. SNR can also be improved by increasing the reflectance of cleaved optical fiber surfaces through Aluminum (Al) coating by sputtering technique [23,24]. A gradient noise on the sensor response caused by the MRI scanner was detected during MR imaging. This noise can be eliminated by digital filtering methods.
Different tissue types have different mechanical characteristics such as stiffness and elasticity. Therefore, the different tissue types can be detected through their response to the FPI force sensor during needle insertion. A biopsy needle with an FPI based force sensor can be a promising tool for cancer diagnosis by differentiating tumorous tissue from healthy tissue, and even recognizing the difference between benign and malignant tumorous tissues without collecting tissue samples during a biopsy operation.
5. Conclusion
In this work, a low profile and MRI compatible FPI based force sensor was fabricated and integrated into a biopsy needle. The proposed novel sensor fabrication method allows to determine the force measurement range and the sensor sensitivity depending on the clinical application needs, by laser micromachining the micro-holes on the borosilicate glass capillary at desired locations. The ambient temperature of the laser source also could be kept constant successfully using a Model Predictive Control based temperature controller during benchtop tests. MRI compatibility and overall performance of the custom designed biopsy needle was tested through benchtop experiments and using a prostate biopsy phantom under MRI. The test results showed that the FPI based force sensor has a force measurement range of 0-13 N with 0.1 N resolution.
Funding
TUBITAK project (115E271).
Disclosures
The authors declare that there are no conflicts of interest related to this article.
References
1. D. A. Woodrum, K. R. Gorny, B. Greenwood, and L. A. Mynderse, “MRI-Guided Prostate Biopsy of Native and Recurrent Prostate Cancer,” Semin. Intervent. Radiol. 33(3), 196–205 (2016). [CrossRef] [PubMed]
2. I. M. Noebauer-Huhmann, M. A. Weber, R. K. Lalam, S. Trattnig, K. Bohndorf, F. Vanhoenacker, A. Tagliafico, C. van Rijswijk, J. C. Vilanova, P. D. Afonso, M. Breitenseher, I. Beggs, P. Robinson, M. C. de Jonge, C. Krestan, and J. L. Bloem, “Soft tissue tumors in adults: ESSR-approved guidelines for diagnostic imaging,” Semin. Musculoskelet. Radiol. 19(5), 475–482 (2015). [CrossRef] [PubMed]
3. L. Lindenberg, M. Ahlman, B. Turkbey, E. Mena, and P. Choyke, “Evaluation of prostate cancer with PET/MRI,” J. Nucl. Med. 57(3), 111S–116S (2016). [CrossRef] [PubMed]
4. M. Roethke, A. G. Anastasiadis, M. Lichy, M. Werner, P. Wagner, S. Kruck, C. D. Claussen, A. Stenzl, H. P. Schlemmer, and D. Schilling, “MRI-guided prostate biopsy detects clinically significant cancer: analysis of a cohort of 100 patients after previous negative TRUS biopsy,” World J. Urol. 30(2), 213–218 (2012). [CrossRef] [PubMed]
5. H. Su, I. I. Iordachita, J. Tokuda, N. Hata, X. Liu, R. Seifabadi, S. Xu, B. Wood, and G. S. Fischer, “Fiber Optic Force Sensors for MRI-Guided Interventions and Rehabilitation: A Review,” IEEE Sens. J. 17(7), 1952–1963 (2017). [CrossRef] [PubMed]
6. W. Shang, H. Su, G. Li, and G. S. Fischer, “Teleoperation system with hybrid pneumatic-piezoelectric actuation for MRI-guided needle insertion with haptic feedback,” in Intelligent Robots and Systems (IROS), IEEE/RSJ International Conference on, (IEEE, 2013), 4092–4098.
7. J. Singh, J. Potgieter, and W. Xu, “Fibre optic force sensor for flexible bevel tip needles in minimally invasive surgeries,” International Journal of Biomechatronics and Biomedical Robotics 2(2/3), 135–140 (2013). [CrossRef]
8. M. Born and E. Wolf, Principles of Optics (Pergamon Press, 1964).
9. K. A. Murphy, M. F. Gunther, A. M. Vengsarkar, and R. O. Claus, “Quadrature phase-shifted, extrinsic Fabry-Perot optical fiber sensors,” Opt. Lett. 16(4), 273–275 (1991). [CrossRef] [PubMed]
10. S. Yin, P. B. Ruffin, and F. T. S. Yu, eds., Fiber Optic Sensors, Second Edition (CRC Press, 2008).
11. S. Xiang, X. Xiang, and C. Feng, “Effects of temperature on laser diode ignition,” Optik (Stuttg.) 120(2), 85–88 (2009). [CrossRef]
12. M. Ishikawa, H. Shiozawa, K. Itaya, G. Hatakoshi, and Y. Uematsu, “Temperature dependence of the threshold current for InGaAlP visible laser diodes,” IEEE J. Quantum Electron. 27(1), 23–29 (1991). [CrossRef]
13. M. Kondow, T. Kitatani, K. Nakahara, and T. Tanaka, “Temperature dependence of lasing wavelength in a GaInNAs laser diode,” IEEE Photonics Technol. Lett. 12(7), 777–779 (2000). [CrossRef]
14. G. P. Agrawal and N. K. Dutta, Semiconductor Lasers (Springer Science & Business Media, 2013).
15. K. Bremer, E. Lewis, G. Leen, B. Moss, S. Lochmann, I. Mueller, and J. Schrotter, “Fibre optic pressure and temperature sensor for geothermal wells,” Sensors (Basel) 2010, 538–541 (2010).
16. X. Zhou, Q. Yu, W. Peng, B. Moss, S. Lochmann, I. Mueller, and J. Schrotter, “Simultaneous measurement of down-hole pressure and distributed temperature with a single fiber,” Meas. Sci. Technol. 23(8), 85–102 (2012). [CrossRef]
17. Q. Wang, C. Qin, D. Wang, and Y. Zhao, “Investigation on stability of extrinsic Fabry-Perot interferometric pressure sensors for high-temperature/high-pressure underground applications,” Instrum. Sci. Technol. 41(2), 143–153 (2013). [CrossRef]
18. K. Chen, X. Zhou, Y. Yang, and Q. Yu, “A hybrid Raman/EFPI/FBG sensing system for distributed temperature and key-point pressure measurements,” in 2015 International Conference on Optical Instruments and Technology: Optical Sensors and Applications, (2015), pp. 9620.
19. A. Wang, H. Xiao, J. Wang, Z. Wang, W. Zhao, and R. G. May, “Self-calibrated interferometric-intensity-based optical fiber sensors,” J. Lightwave Technol. 19(10), 1445 (2001). [CrossRef]
20. M. C. Yip, S. G. Yuen, and R. D. Howe, “A robust uniaxial force sensor for minimally invasive surgery,” IEEE Trans. Biomed. Eng. 57(5), 1008–1011 (2010). [CrossRef] [PubMed]
21. U. X. Tan, B. Yang, R. Gullapalli, and J. P. Desai, “Triaxial MRI-compatible fiber-optic force sensor,” IEEE Trans. Robot. 27(1), 65–74 (2011). [CrossRef] [PubMed]
22. S. B. Kesner and R. D. Howe, “Design principles for rapid prototyping forces sensors using 3-D printing,” IEEE/ASME Trans. Mechatron. PP(99), 1–5 (2011). [CrossRef] [PubMed]
23. T. Yoshino, K. Kurosawa, K. Itoh, and T. Osa, “Fiber-optic Fabry-Perot interferometer and its sensor applications,” IEEE J. Quantum Electron. 18(10), 1624–1633 (1982). [CrossRef]
24. W. D. Hogg, D. Janzen, T. Valis, and R. M. Measures, “Development of a fiber Fabry- Perot strain gauge,” Fiber Optic Smart Structures and Skins IV 1588, 300–307 (1991). [CrossRef]
