Optical Fiber Sensors Embedded Into Medical Textiles ... - IEEE Xplore

IEEE SENSORS JOURNAL, VOL. 8, NO. 7, JULY 2008

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Optical Fiber Sensors Embedded Into Medical Textiles for Healthcare Monitoring Augustin Grillet, Damien Kinet, Jens Witt, Marcus Schukar, Katerina Krebber, Fabrice Pirotte, and Annick Depré

Abstract—The potential impact of optical fiber sensors embedded into medical textiles for the monitoring of respiratory movements in a magnetic resonance imaging environment is presented. We report on three different designs, all textile based: a macrobending sensor, a Bragg grating sensor, and a time reflectometry sensor. In all three cases, the sensing principle is based on the measure of the elongation of the abdominal circumference during breathing movements. We demonstrate that the three sensors can successfully sense textile elongations between 0% and 3%, while maintaining the stretching properties of the textile substrates for a good comfort of the patients. Index Terms—Fiber Bragg grating, fiber-optic sensor, macrobending sensor, medical textile, medical resonance imaging (MRI), polymer fiber (POF) optical time domain reflectometry (OTDR) sensor.

I. INTRODUCTION EALTHCARE monitoring is a general concern for patients requiring a continuous medical assistance and treatment. In order to increase mobility of such patients, a huge effort is pursued worldwide for the development of wearable monitoring systems able to measure vital physiological parameters such as respiration movements, cardiac activity, pulse oxymetry and temperature of the body [1]. Technical or smart textiles play a growing role in these developments as they are well suited for wearability and can ensure comfort to the user [2]. Most developments up to now were focused on the use of electrical sensors, but recently, a new project supported by the EU under FP6 and called Optical Fiber Sensors Embedded into Technical Textiles for Healthcare (OFSETH) was launched [3]. Its aim is to take advantage of pure optical sensing technologies for extending the capabilities of medical technical textiles for wearable health monitoring, especially in magnetic resonance imaging (MRI) environments [4], [5].

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Manuscript received August 1, 2007; revised December 5, 2007; accepted January 3, 2008. Published July 16, 2008 (projected). This work was supported by the EU Sixth Framework Program under Contract IST-2004-027869. The views expressed here are those of the authors only. The Commission is not liable for any use that may be made of the information contained therein. The associate editor coordinating the review of this paper and approving it for publication was Prof. Jose Lopez-Higuera. A. Grillet was with Multitel, 7000 Mons, Belgium. He is now with Barco n.v., Kennedy Park 35, B-8500 Kortrijk (e-mail: [email protected]). D. Kinet is with Multitel, 7000 Mons, Belgium (e-mail: [email protected]). J. Witt, M. Schukar, and K. Krebber are with the Federal Institute for Material Research and Testing, 12205 Berlin, Germany (e-mail: [email protected]). F. Pirotte is with Centexbel, 4650 Herve, Belgium (e-mail: fabrice.pirotte@ centexbel.be). A. Depré is with Elasta, 8790 Waregem, Belgium (e-mail: [email protected]). Color versions of one or more of the figures in this paper are available online at http://ieeexplore.ieee.org. Digital Object Identifier 10.1109/JSEN.2008.926518

Fig. 1. Example of the integration of an optical fiber into a woven fabric.

Investigation of the capabilities of optical fiber sensors for healthcare monitoring in MRI environment has been largely reported in the past, due to the well-known immunity of fiber optics against electromagnetic radiations. Monitoring of respiratory motions and/or detection of heart beats were demonstrated using pure optical methods, in particular the bending loss technique using either a coil [6], [7] or a loop [8], [9] shape, the interferometry technique [10], and the grating technique using either fiber Bragg gratings (FBG) [11], [12] or long period gratings (LPG) [13]. While most of these sensors have shown great sensing capabilities, their use in a medical environment has been however limited, on one hand because of practical considerations, and especially because of the poor usability of the sensors from medical staff point of view. On the other hand, textile based sensors, such as garments and belts, clearly showed better capabilities regarding this limit, but none of the designs proposed so far and mentioned above turned into real product. This is mainly due to their poor compatibility with cost-effective industrial textile processes, where manual operations such as confection have to be reduced to a minimum. Due to their fibrous nature, optical fibers have however a serious advantage over other kind of sensors when integration into textiles is considered. Indeed, an optical fiber is in some way a filament and can ideally be processed like standard textile yarns, and several techniques, such as wrap and weft knitting, weaving and stitching, which can be used to manufacture textile fabrics with embedded sensing elements, are being explored in this project Fig. 1. In this paper, we review recent developments aimed at producing a textile based respiratory rate sensor, either at the thorax or the abdomen locations. Besides being MRI compatible, the garment should have good sensing capacities, accommodate several sizes, and be comfortable to wear. For practical reasons, a bandage configuration allowing a rapid installation by the anesthetists was selected. This solution is also well in line with our requirements for wearability, as opposed to other concepts based on a sensing bed or mattress, for instance [14], [15]. The bandage should be stretchable so as to accommodate several

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sizes and fit well on the patient skin. As such, an elastic fabric, with at least 50% elasticity should be preferred. The measurement principle will be similar to existing electrical respiratory sensors that detect the elongation of the abdominal or thoracic circumferences through inductance plethysmography technique [16] or use of piezo-resistive yarns [17]. In a first step, no fine precision is required concerning the amplitude of the breathing movements, as only the frequency corresponding to the respiratory rate will be derived from the signal. In OFSETH, three optical sensing techniques have been proposed for the monitoring of elongation in textile fabrics: the bending loss sensor, the FBG sensor and the optical time domain reflectometry (OTDR) sensor, which three architectures are detailed in the next chapters. In the following Section, we report on a simple design based on textile induced variations of macrobending losses in a single-mode optical fiber. The design is capable of up to 100% elongation, and is fully compatible with a narrow fabric production process. In Section III, the same bandage is used with an OTDR sensing system, which has the potential to sense simultaneously movements from various locations on the patient (e.g., abdomen and thorax) with a single monitoring unit. In this example, a polymer fiber (POF) is used instead, for a higher robustness to the textile integration process. Finally in Section IV, we consider a third option based on a silica FBG embedded in an elastic bandage and capable of a better resolution for sensing weak breathing movements.

Fig. 2. Simple arrangement for a macrobending sensor. When stretching the substrate, the curvature in the fiber will open.

II. MACRO-BENDING SENSOR A. Macro-Bends in Optical Fibers Bending effects in optical fibers have been largely studied in the 1970s and 1980s as optical fibers were progressively developed and installed on the field [18]. As is well known, bends cause power propagating in guided modes to be lost by coupling into radiation modes, and consequently, fiber designs as well as cabling processes have been continuously optimized for a higher robustness to bending effects. Meanwhile, their uses for sensing purposes have led to the first generations of optical fiber sensors based on microbending effects that originate from periodic perturbation of the fiber along its axis through a pressure transducer [19]. The use of macrobending sensors was far less used, in particular due to the lower sensitivity and the higher complexity required for designing a transducer able to interact with the environment. However, in the case where no fine precision regarding elongation measurement is requested, macrobending sensors have major advantages. First, their interrogation is very simple and requires low cost and compact components. Second, their integration into textile fabrics may be very straightforward, owing to their simple design. Especially, a macrobending transducer is very simple to achieve using a stretchable substrate, such as an elastic textile fabric, containing an optical fiber bent in a half-a-loop or “U” shape, as shown in Fig. 2. When stretching the substrate, the curvature radius will increase, thus reducing the bend loss in the fiber, and the intensity variations at the output of the optical fiber will therefore reflect the relative variation of the substrate length. It was shown, however, that macrobends in single-mode optical fibers give rise to mode coupling between the fundamental

Fig. 3. Typical loss measurement in a bent optical fiber. Top: varying bending radius. Bottom: varying monitoring wavelength.

mode and whispering gallery (WG) modes reflected at claddingcladding, cladding-coating, coating-coating, or coating-air interfaces, towards the fiber core [20]. Depending on key parameters such as the light wavelength, the curvature radius, the fiber index profile and the core/cladding/coating dimensions, synchronous mode coupling may occur, which will lead to constructive interference and thereby reduce the loss induced by the fiber bend [21], [22]. When varying the bending radius of a specific fiber or the monitoring wavelength, such effect will result in oscillations in the loss curves, thus making the signal analysis far less straightforward, as shown in Fig. 3. Here, the sample substrate was an elastic textile fabric with an initial length of 78 mm. The fiber used was a standard single-mode fiber compatible with ITU-T G.652, and the fiber bend radius at zero strain was optimized so as to keep the total loss compatible with the power budget of the measuring system. Fig. 4 shows the bend loss evolution while stretching several times the sample described just above. The experimental setup

GRILLET et al.: OPTICAL FIBER SENSORS EMBEDDED INTO MEDICAL TEXTILES FOR HEALTHCARE MONITORING

Fig. 4. Typical signal from the macrobend U-shape fiber sensor subject to periodic stretching and relaxation movements. The discontinuities in the signal curve are the main drawback of macrobending sensors.

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Fig. 6. Influence of the total amount of bending on the sensor signal. The input power ( 2.65 dBm) was kept unchanged for all configurations.

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B. Periodic Macrobending Sensor Behavior

Fig. 5. Top: structure of the periodic bending sensor. The white line indicates the stitching points. Bottom: measurement setup (top view); the graph on the right indicates the linear and cyclic movement of the translation table.

used is described in Fig. 5, where a translation table is used together with a cylinder fixed on a breadboard. The linear movement of the translation table results in tangential elongation for the sensing bandage, thereby simulating respiratory movements. In this experiment, the optical source was a FP laser source emitting at 1310 nm. The fabric was stretched from 78 mm to 108 mm, so that if it was part of a respiratory belt worn by a male adult having a 100 cm waistline, it would represent a 3% elongation of the abdominal circumference, which is the typical maximal value for adults. Whereas a frequency information can still be derived from the signal shown in Fig. 4, it is feared that the oscillations in the curve could lead to a possible wrong computation of the respiratory rate, especially in case of weak amplitude breathing movements, where the peak-to-peak amplitude of the sensor signal could be significantly reduced.

In order to reduce the importance of these oscillations, we here consider the potential impact of the total amount of bending in the fiber on one hand, and of the bandwidth of the light source used for the measurement on the other hand. Indeed, previous studies having investigated oscillations effects in bend loss measurements were mostly carried out with a fixed amount of bending and with a narrow band laser [17], [19]. In our work, the bending amount was varied by adding more periods of the U-shape on the textile bandage, as shown in Fig. 5. For the purpose of the investigation, this was realized by hand-stitching technique; however, this design being periodic is also compatible with fully automatized textile processes. The loss in the fiber was then monitored using a tunable laser source as the bandage was stretched using the setup described in Fig. 5. Fig. 6 shows attenuation measurements successively performed at 1310 nm on the bandage being linearly stretched and relaxed, with 10 to 0 curvatures, respectively, each curvature having the same U profile. As anticipated, both the loss budget and the sensitivity of the measurement significantly increase as the amount of bending is increased in the fiber. For the same maximal elongation ( 38%) of the textile substrate, the sensor response was increased by a factor 400, from less than 3 dB with a single-loop design to more than 28 dB with the 10 loops configuration. In addition, the impact in the sensor response of the oscillations originating from mode coupling is significantly reduced, as shown in Fig. 7. For instance at 1310 nm, the amplitude of the oscillations with regards to the total signal amplitude was decreased by a factor 2.5, from 41.7% with a single-loop design, to 16.4% with the 10 loops configuration. In order to further improve the linearity of the sensor, the impact of the bandwidth of the light source used for the measurement was then considered. Because the undulations exhibit strong wavelength dependence, as shown in Fig. 3 and Fig. 7, it is expected that their amplitude could be reduced by using a source with a larger bandwidth, which could lead to an averaging phenomenon. Therefore, the same experiment was run with a different light source (LED type) having a much larger bandwidth, as shown in Fig. 8. The results are shown in Fig. 9, where the sensor response is compared when being interrogated

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Fig. 9. Comparison of the sensor response when being monitored by a narrowband light source or a wideband light source (SLED), for 1 and 10 loops.

Fig. 10. Respiratory rate sensing belt (Belt #1) with the macrobending sensor embedded. Fig. 7. Measured sensor response versus wavelength and substrate elongation for 1 loop (top) and 10 loops (bottom).

C. Textile Integration

Fig. 8. Spectra of the optical sources used for interrogating the macrobending sensor.

with the two light sources, for the 1 loop and 10 loops configurations. The results clearly indicate an optimization in the sensor linearity, since no more oscillations are present especially for the 10 loops configuration, which then turns into a higher reliability for the sensor.

The above described design was then processed in a narrow fabric production using textile yarns and a standard SMF. The optical fiber had an optimized coating for a higher robustness against rub from the metal rolls used in the textile machines, and was fully embedded in the textile fabric during the process. About 10 m of sensing bandage were produced, and the bandage was capable of up to 100% elasticity with no damage on the optical fiber. Two different sensing belts were produced. The first one (Belt #1) was made by the sensing bandage only, whereas the second (Belt #2) had only one part elastic (the sensing area) and the rest nonelastic, similar to the one shown in Fig. 5. The optical fiber was terminated by FC connectors, and Velcro fasteners were stitched on the bandages. The two belts were tested simultaneously by a healthy volunteer having a waistline of 94 cm and performing pure abdominal breathing only as would be normally the case for sedated patients. Thanks to their elastic properties and the Velcro fasteners, the belts could be easily and comfortably fitted to the very waistline of the patient, as shown in Fig. 10. Typical respiration traces measured with a wideband source and a benchtop powermeter are shown in Fig. 11, where the two sensing belts are compared in terms of sensitivity.


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Fig. 12. Sensor performances when used with four different patients. The sensor was monitored with a 1550 nm SLED.

Fig. 11. Respiratory abdominal movements recorded simultaneously by the two belts embedding a macrobending sensor, at 1310 and 1550 nm.

In Belt #2, the whole perimeter elongation is concentrated on the sensing part only, because the sensing part is elastic and the rest of the belt is not. Therefore the variation in the curvature radius of the fiber bends is higher than for Belt #1, wherein the elongation is absorbed by the whole length of the belt. Consequently, Belt #2 design has the best sensitivity, as shown in Fig. 11. In addition, because of the larger mode field diameter at longer wavelengths, using a SLED source at 1550 nm instead of 1310 nm allows increasing the sensitivity further more. Finally, the Belt #2 sensor was used on four volunteers performing abdominal breathing and having different waistlines ranging from 80 cm to 110 cm. For each patient, the initial tension was set so that the comfort properties were excellent, according to the patient’s own feeling. Fig. 12 shows the mean sensor response to a breathing movement for each patient, averaged over more than one minute of respiration. From the results shown in Figs. 11 and 12, we can therefore conclude on the great potential of this simple sensing principle for the monitoring of respiration rate of anesthetized patients. Fig. 12 also illustrates that the influence of different patients’ morphology on the sensor performances should be further explored. III. POF OTDR SENSOR We now consider the interest from the medical staff to take information from both abdominal (for spontaneous ventilation)

and thoracic (for intubated patients) movement. A distributed measurement of the respiratory signal, such as one based on OTDR techniques and thus using only one monitor and one sensor fiber would be advantageous. Using the OTDR technique it is possible to focus on a special part of the fiber and so to differentiate between abdominal and thoracic respiration. A distributed, OTDR measurement makes possible to get only the required sensor information and to neglect loss contributions from nonsensing parts. In addition, a reflectometry sensing system has the advantage of requiring only one fiber connection, which enables a quicker installation of the system on the patient. In this experiment, a polymer fiber was selected instead, because it is feared that the silica fiber would not accommodate an OTDR sensor design with regards to the textile integration process. POF have indeed outstanding elastic properties that make them well suited for integration into technical textiles. Another interesting criterion for selecting POF as medical sensor is its biocompatibility, especially in case of fiber breakage. Therefore, a distributed fiber-optic sensor based on POF and using the OTDR technique is being investigated in OFSETH [23]. Fiber optic sensors based on POF take advantage of the high elasticity and high breakdown strain of POF of up to more than 40% [24], [25]. The feasibility of measuring the respiratory waveform and rate in real time by the POF OTDR technique was demonstrated on a healthy adult during normal breathing [23]. A textile sample with integrated POF, based on macrobending effects in optical fibers described in Section II was tested for the above mentioned purposes (Belt #2 design). It has been shown that the level of the backscattered light changes at locations where any mechanical deformation is applied to the POF [23]–[25]. Macrobending effects in POF induce changes of the backscattering in the corresponding area of the fiber, too, and can be easily detected by the OTDR technique. The textile sample was attached around the abdomen of the adult and the elastic part of the textile was placed in the area experiencing the most elongation due to the breathing movement (see Fig. 13). The sensor signal was acquired by an OTDR device produced by Tempo (OFM20), which operates at 650 nm wavelength, allows a two-point special resolution of 5 cm and has a dynamic range of 20 dB. The device makes possible to measure an OTDR trace in less than 1 s with a sufficient SNR.

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Fig. 15. Picture of the textile fabric used as the FBG substrate, with the fiber containing the FBG illuminated with a red light.

Fig. 13. Monitoring of respiratory abdominal movement of an adult by POF OTDR.

single-mode POF associated with the high attenuation of the POF cladding at standard writing wavelengths has led to a very limited success. Indeed, the fabrication of FBGs in POF was reported by only two groups worldwide [28], [29], thus illustrating both the difficulty and the delay in reaching industrial capacity for this technique. Therefore, other solutions, that could benefit from the proved reliability of silica FBGs and thus with a shorter time to market should be explored. Thanks again to the textile integration, we developed a design able to measure large strain values with an ordinary silica FBG sensor. The FBG was embedded onto an elastic textile fabric, which allows protecting the optical fiber and holding it in place, while still permitting a large flexibility. A. Sensor Design and Experimental Set-Up

Fig. 14. Respiratory abdominal movement recorded by POF OTDR.

The changes of the abdominal circumference due to the breathing movement were recorded simultaneously. The result is shown in Fig. 14 and demonstrates the potential of the POF OTDR technique for the considered monitoring purposes IV. SILICA FBG SENSOR In this last section, we focus on a different sensor design based on the use of a fiber Bragg grating sensitive to strain variations. Although it requires a more complex monitoring system, it is interesting for our application because of its better sensing linearity and higher resolution in case of weak elongation movements. The main difficulty however is to cope with large stretching capabilities required by the application. Silica FBGs have indeed demonstrated large capabilities as strain sensors in many applications including monitoring of civil and aeronautics structures. Optical strain gauges are now standard commercial products allowing to measure deformations in the range of 1% typically. But for higher strain values, silica FBGs traditionally show significant limits due to the poor elasticity of silica glass. On the other hand, polymers such as polymethylmethacrylate (PMMA) have a Young’s modulus about 30 times lower than silica glass, together with a much higher breakdown strain. Therefore, optical sensors—and especially FBGs—based on POF have attracted a huge interest for sensing large deformations [26], [27]. However, the lack of commercial

The FBG was written onto a single-mode silica fiber with a coating transparent to UV radiation at 248 nm and a photosensitive core co-doped with tin [30]. While improving the efficiency of the manufacturing process and its cost, such fiber has the advantage of easing the embedding process due to the absence of splices and recoating. Glass optical fibers are indeed very fragile when compared to other textile yarns, and therefore specific care should be taken when manipulating fibers in the textile manufacturing process. For this reason, only optical fibers with sufficient robustness should be used, thus giving preference to fibers with a UV transparent coating that are not damaged during the photo-writing process. A 35.5 cm long elastic textile fabric, as shown in Fig. 15, was used as a substrate. The fiber was stitched onto the fabric and additional glue was used on both side of the FBG for a better adherence. The distance between the glue points was 11.5 cm, so that when stretching the fabric, only part of the strain is transmitted to the FBG, while the stitching design makes it possible for the rest of the optical fiber not to be impacted by the elongation. The stretching capabilities of the textile including the FBG were explored using a mechanical characterization equipment, which consists of a moveable platform (via a step motor) and a load cell. The textile was fixed to both the platform and the load cell via clamping blocks with an effective stretching length of 24.7 cm. The effective strain on the optical fiber and its integrity were monitored during the stretching experiment, by continuously scanning the Bragg wavelength of the FBG using a Micron Optics si425 reading unit with a wavelength resolution of 0.2 pm. During the tests the room temperature was quite stable so that no temperature correction of the measured Bragg wavelength shift with strain was necessary.


Fig. 16. Bragg wavelength shift as a function of the textile elongation.

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Fig. 17. Signal and force as a function of the time during 10 strain cycles between 0% and 5% strain.

B. FBG Sensor Performances The textile was stretched in steps of 0.4% and the Bragg wavelength shift was monitored as a function of the elongation. , the FBG At 0% elongation and room temperature of 23.5 of 1550.019 nm, a spectral width has a Bragg wavelength to 0.114 nm, and a reflectivity of 99.7%. The textile could be stretched up to 41.2% elongation, with no degradation of the optical fiber. This limit comes from the limited range of our moveable platform and not from the optical fiber breakdown strain. Fig. 16 shows the Bragg wavelength shift with regards to the bandage elongation, where one can see that the shift depends linear on the elongation, except for elongations lower than 8.5%, which is due to the fact that the fiber was not ideally tense between the glue points at the 0% elongation state. In the linear regime, the so-formed optical gauge has a sensitivity of 0.35 nm/% substrate elongation. Such value should easily lead to a 0.1% measurement accuracy for elongation measurements using even low-cost FBG interrogation units. Using the strain coefficient of the Bragg wavelength (standard for Ge doped silica fibers at 1.5 ), the value of 1.2 measured Bragg wavelength shift was converted into strain and is shown in Fig. 16 (right axis). The sensor design translated the high textile strain of more than 40% to a strain of 0.8% that was measured by the FBG, which confirms that the fiber was maintained well below the breakdown strain level, even for a substrate elongation of more than 40%. This result thus confirms that such a bandage could be safely used to monitor textile elongations of a few % with a sub% resolution. In addition, the 40% elongation capability gives some flexibility to set the zero % elongation level, thereby allowing to adjust the bandage to the specific waistline of a patient for instance. In order to show the suitability of the proposed sensor to measure also a cyclic strain such as respiratory movements, 10 strain cycles with a maximum strain value of 5% were applied to the textile and the response of the FBG was recorded. Fig. 17 shows the Bragg wavelength shift and the applied force (in arbitrary units) as a function of the time during the 10 strain cycles. V. CONCLUSION Medical textiles are a potential new niche market for optical fiber sensors. In this paper, we report on optical fiber sensors embedded into textile fabrics for the monitoring of res-

piratory movements in MRI environment. Three different designs—macrobending, FBG and OTDR- allowing to sense textile elongations between 0% and 3% have been investigated and successfully demonstrated. Further study should now focus on the influence of different patients’ morphology on the sensor performances, as well as on the textile integration issues. REFERENCES [1] A. Lymberis and A. Dittmar, “Advanced wearable health systems and applications,” IEEE Eng. Med. Biol. Mag., vol. 26, no. 3, pp. 29–33, May/Jun. 2007. [2] A. Lymberis and S. Olsson, “Intelligent biomedical clothing for personal health and disease management: State of the art and future vision,” Telemed. J. e-Health, vol. 9, no. 4, pp. 379–386, 2003. [3] [Online]. Available: http://www.ofseth.org [4] J. De Jonckheere, M. Jeanne, A. Grillet, S. Weber, P. Chaud, R. Logier, and J.-L. Weber, “OFSETH: Optical Fibre Embedded into technical Textile for Healthcare, an efficient way to monitor patient under magnetic resonance imaging,” in Proc. 29th Ann. Int. Conf. IEEE EMBS, 2007, to be published. [5] M. F. Dempsey and B. Condon, “Thermal injuries associated with MRI,” Clin. Radiol., vol. 56, pp. 457–465, 2001. [6] A. T. Augusti, F.-X. Maletras, and J. Mason, “Improved fiber optic respiratory monitoring using a figure-of-eight coil,” Physiol. Meas., vol. 26, 2005. [7] C. Davis, A. Mazzolini, and D. Murphy, “A new fiber optic sensor for respiratory monitoring,” Austr. Phys. Eng. Sci. Med., vol. 20, no. 4, 1997. [8] A. Babchenko, B. Khanokh, Y. Shomer, and M. Nitzan, “Fiber optic sensor for the measurement of respiratory chest circumference changes,” J. Biomed. Opt., vol. 4, no. 2, pp. 224–229, Apr. 1999. [9] A. Raza and A. T. Augusti, “Optical measurement of respiration rates,” in Proc. 7th Conf. Sensors A, 1995. [10] D. Varshneya, J. L. Maida, and L. A. Jeffers, “Fiber Optic Interferometric Vital Sign Monitor for Use in Magnetic Resonance Imaging, Confined Care and in-Hospital,” U.S. Patent 6816266, Nov. 9, 2004. [11] G. Wehrle, P. Nohama, H. J. Kalinowski, P. I. Torres, and L. G. G. Valente, “A fiber optic Bragg grating strain sensor for monitoring ventilatory movements,” Meas. Sci. Technol., vol. 12, pp. 805–809, 2001. [12] D. Gurkan, D. Starodubov, and X. Yuan, “Monitoring of the heartbeat sounds using an optical fiber Bragg grating sensor,” in Proc. 4th IEEE Conf. Sensors, Irvine, CA, Nov. 2005. [13] T. Allsop et al., “Application of long-period grating sensors to respiratory function monitoring,” Smart Med. Biomed. Sensor Technol. II, vol. 5588, pp. 148–156, 2004, Proc. SPIE. [14] W. B. Spillman et al., “A smart’ bed for non- intrusive monitoring of patient physiological factors,” Meas. Sci. Technol., vol. 15, pp. 1614–1620, 2004. [15] X. Xu, W. B. Spillman, Jr., R. O. Claus, K. E. Meissner, and K. Chen, “Spatially distributed fiber sensor with dual processed outputs,” presented at the 17th Int. Conf. Optical Fibre Sensors, May 23–27, 2005, Paper Tu2-6, unpublished.

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Augustin Grillet graduated in physics from the University of Essex, Essex, U.K., and INSA, Toulouse, France. He started his career in 2001 at Alcatel Submarine Networks, U.K., and in 2002 moved to the Applied Photonics Department, Multitel, Belgium. His main research interests there were related to optical telecommunications, optical amplifiers, and optical fiber sensors, in which fields he coauthored about 15 papers in various conferences and journals. He has been involved in several regional and European R&D projects on these topics. In September 2007, he joined the Corporate Research Division of Barco, Belgium.

Damien Kinet received the License degree in physics (oriented surface studies) from the Facultés Universitaires, Notre de Dame de la Paix à Namur, Belgium, in 1993 and the Special License degree in nuclear sciences from the Université de Liège, Belgium, in 1994. He is now a Member of Research Staff with the Applied Photonic Department, Multitel, Belgium. He studies the design of fiber Bragg grating, the photosensitivity and the effect of gamma radiation in optical glasses, and is involved in the European project OFSETH. He has authored and coauthored several scientific publications and international presentations.


Jens Witt received the Diploma degree in physics from the Technical University of Berlin, Germany, in 2002. He received the qualification “European Laser Engineer” in the postgraduate course program of the EuroLaser Academy from the Technical University of Vienna, Austria. With a dissertation on colorimetric methods for the production of test charts, he received the Ph.D. degree in electrical engineering from the Technical University of Berlin in 2006. Since 2003, he has been with the Federal Institute for Materials Research and Testing (BAM), Berlin. He was involved in the group Visual Methods and Image Reproduction in NonDestructive Testing (NDT). Currently, he is involved in the group Fiber Optic Sensors in research and development of fiber optic sensors within the framework of national and European research projects.

Marcus Schukar studied microsystems technology at the University of Applied Sciences FHTW, Berlin, Germany. He received the graduate Engineer degree in 2006 with a diploma thesis on the monitoring of ionizing radiation on electron storage beams with fiber Bragg grating sensors. Since 2006, he has been an Engineer with the Federal Institute for Materials Research and Testing (BAM), Berlin, where he works on a number of national and European research projects dealing with the development of distributed fiber optic sensors, polymer fiber optic sensors, and FBG sensors for different applications.

Katerina Krebber studied physics and received the Ph.D. degree in electrical engineering in 2001 from the University of Bochum, Germany, with a Ph.D. dissertation on nonlinear effects in optical fibers and distributed fiber-optic sensors. She has 14 years experience in the field of fiber optics and fiber-optic sensors. She was engaged in research and development of fiber-optic communication systems at Siemens AG and in research and development of fiber-optic radiation sensors at the Fraunhofer Institute INT, Germany. Since 2004, she has been with the Federal Institute for Materials Research and Testing (BAM), Berlin. She leads at BAM a number of national and European research projects dealing with the development of distributed fiber-optic sensors, polymer fiberoptic sensors, and FBG sensors for different applications. She is author and coauthor of more than 60 scientific publications including invited talks at international conferences and patents, and serves as a referee for scientific journals.

Fabrice Pirotte has a received the diploma of Textile Engineer in the field of chemical processes. He has been with Centexbel since 1994. He has done research on recycling by using Near Infra Red Analysis technology and on the bleaching process. During the last 5 years, he has participated to the development of intelligent textiles on the European and the national level (Mermoth and OFSETH projects) and of medical textiles. He was also responsible of the knitting platform. Since January 2008, he has been with NG, s.a., Herve, Belgium.

Annick Depré is with Elasta, Waregem, Belgium. She coordinates research and development projects within the company. Elasta is a narrow fabric manufacturer, weaving, knitting, and braiding bands, straps, and cords. Her previous working experience is situated in market development, marketing, and communication in the telecommunication, financial, and textile industries. She has an experienced team of mechanicals and machine operators in the development of research projects.