Wireless Power Recharging for Implantable Bladder Pressure Sensor Peng Cong, Michael A. Suster, Nattapon Chaimanonart, and Darrin J. Young Department of Electrical Engineering and Computer Science Case Western Reserve University Cleveland, Ohio, USA
[email protected] Abstract—This paper presents a wireless power recharging system design for implantable bladder pressure chronic monitoring application. The power recharging system consists of an external 4-turn 15-cm-diameter powering coil and a silicone-encapsulated implantable spiral coil with a dimension of 7 mm x 17 mm x 2.5 mm and 18 turns, which further encloses an ASIC with a programmable charging current and logic control capability, a 3-mm-diameter 12-mm-long rechargeable battery, and two ferrite rods. The ferrite rods are employed to improve the quality factor of the implantable coil. For a constant charging current of 100 µA, an RF power of 2.4 mW needs to be coupled into the implantable microsystem through tuned coil loops. With the two coils aligned coaxially or with a tilting angle up to 30o, an external RF power of 7W or 25W is required, respectively, for a large coupling distance of 20 cm at an optimal frequency of 3 MHz.
I.
INTRODUCTION
Urinary incontinence is a severe medical symptom caused by spinal injury, aging and various medical conditions. Urodynamics diagnose based on urethral catheter insertion for bladder pressure short-term monitoring is inconvenient and unreliable [1-3]. It is, therefore, highly desirable to develop a wireless implantable pressure sensor, which can be implanted under the bladder mucosa layer to chronically monitor the bladder pressure [4-6], followed by wireless data transmission to an external receiver for biomedical analysis and ultimately feedback control through functional electrical stimulation [7, 8]. The proposed wireless implant system will be inserted through a cystoscope, which can accommodate a packaged implant with a dimension of approximately 5 mm x 9 mm x 18 mm [4]. A such miniature size precludes the usage of a primary battery for chronic monitoring, thus calling for a small rechargeable battery. Advances in battery technology provide miniaturized and reliable rechargeable batteries with adequate power density. Among the numerous available technologies, the rechargeable Lithium ion (Liion) batteries exhibit the highest energy densities typically ranging from 200 Wh/l to 450 Wh/l (watt hour per liter) [9]. For instance, rechargeable Li-ion QL-0003I battery (from Quallion LLC) with a small volume of 0.08 cm3 can provide a capacity of 3 mAh with a typical recharging cycle of 3000 times without significant performance degradation. With a daily recharging, the battery can be implanted in patient for approximately ten years. Therefore, an RF recharging
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capability is highly critical in the proposed overall implant system design. Inductive-coupling-based powering techniques have been widely used for biomedical implants [10-13]. In these applications, an external RF power source drives a coil loop, which couples the RF energy into an implanted coil through magnetic coupling to power the microelectronics or to recharge the battery. The coupling coils are typically separated by a short distance on the order of a centimeter. However, in the proposed application depending on individual patient’s size and weight, a large power coupling distance between 10 cm to 20 cm can be required, thus imposing a significant design challenge for achieving an optimal and efficient power coupling. In this paper, a wireless power recharging system architecture is presented with an optimized design of inductive coupling link, followed by coupling characterization and external power requirement calculation at an optimal operating condition. II. WIRELESS POWER RECHARGING SYSTEM The proposed wireless implantable bladder pressure sensor with an RF power recharging capability is presented in Figure 1, where a miniature pressure sensor powered by a rechargeable battery is implanted under the bladder mucosa layer to chronically monitor the bladder pressure.
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Figure 1. Wireless implantable bladder pressure sensing architecture.
IEEE SENSORS 2009 Conference
The bladder pressure information can be wirelessly transmitted to an external receiver for biomedical analysis and ultimately for closed-loop control based on nerve electrical stimulation [7, 8]. Preliminary implant experiments have demonstrated that a cystoscope used to insert an implant can accommodate a packaged device with a dimension of approximately 5 mm x 9 mm x 18 mm [4]. Therefore, an implantable RF coil exhibiting a lateral dimension of 7 mm x 17 mm and a height of 2.5 mm with 18 turns was constructed to enclose an ASIC, a pressure sensor, a rechargeable battery with a 3mm-diameter and a 12-mm-length, and two ferrite rods, as depicted in Figure 2, to implement a wireless implantable bladder pressure monitoring microsystem. The ferrite rods are employed to improve the coil quality factor. The coil dimension is chosen to maximize its inductive coupling factor to an external powering coil.
Figure 2. Wireless implantable bladder pressure monitoring microsystem.
The microsystem can be encapsulated in a biocompatible package with a proper opening port for pressure transduction. Figure 3 shows a photo of a siliconeencapsulated implantable coil with two ferrite rods and a rechargeable battery positioned at center. The remaining space will be occupied by an ASIC, pressure sensor, and other necessary discrete components. Ferrite Rods
Rechargeable Battery
The RF power is coupled to the implanted coil, L2, from an external coil, L1, tuned to the same frequency through mutual inductance. The received RF power is then rectified and filtered to recharge a rechargeable battery through the recharging electronics as shown in the figure. In the prototype design, a 15-cm-diameter 4-turn coil exhibiting an inductance of approximately 5 µH is employed as the external powering loop. The coil dimension is chosen to be suitable for future implementation as a part of a wearable unit. It is difficult to achieve a high supply voltage for the implant by RF powering over 10 cm to 20 cm distance without demanding a large external RF power. A CMOS voltage doubler rectifier is thus chosen to implement the rectifier for a reduced voltage gain requirement compared to full-wave rectifier and half-wave rectifier counterparts. A constant current mode operation followed by a constant voltage mode operation is typically required to fully recharge a battery. However, to increase the battery recharging cycles, a full battery recharge is not desirable. A constant current recharging scheme is, therefore, chosen to charge up the battery to a few hundred of milli-volts below the maximum allowed battery voltage. Furthermore, the constant current mode operation results in an approximately constant impedance loaded to the LC tank in the implantable microsystem, which is desirable for achieving a stable RF powering efficiency. It can be shown that the voltage gain, Av, from Vin to Vout can be expressed as [14]:
Av =
Vout = Vin
ω 2 ML2
(ω L2 ) 2 (ω M ) + R1 R2 + R1 RAC _ load
where M is the mutual inductance between the external coil of L1 and implantable coil of L2, ω is the tuned resonant frequency of two coils, R1 and R2 are the series resistances associated with L1 and L2, respectively, and RAC-load is the load resistance of the CMOS rectifier presented to the tuned implantable coil. Further simplifying the equation can result in the following expression:
Av ≈ kQ1Q2 L2 / L1 17 mm Figure 3. Silicone-encapsulated coil with battery and ferrite rods.
Figure 4 presents the system design architecture of the overall RF recharging electronics.
Figure 4. RF recharging electronic design architecture.
,
2
1 , 1+ β
where k is the coupling factor, Q1 and Q2 are the loaded quality factor of the external coil and unloaded quality factor of the implantable coil, respectively, and β is the ratio between the implantable coil equivalent serial load resistance
(ω L2 ) RAC _ load and R2, 2 as β = (ω L2 ) RAC _ load R2 . of
2
which can be expressed
The power coupling efficiency, defined as the ratio of the power delivered into the CMOS voltage doubler and the power dissipated in the external tuned coil loop, under a weak coupling condition can be further derived as
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ηcoupling =
k 2Q1Q2 β (1 + β ) 2
.
From the above equations, it is evident that improving inductive coupling factor and coils quality factors can directly enhance the voltage gain as well as the power coupling efficiency. Moreover, β plays in a critical role in performance. Figure 5 plots the terms of 1/1 + β and
β / (1 + β ) 2 versus β, indicating (1) for a small β, a large voltage can be obtained, however, with a penalty of a reduced power coupling efficiency, and (2) for a large β both gain and efficiency will be degraded. An optimal condition as a trade-off between voltage gain and efficiency occurs around β = 1, which represents a power impedance matching between RAC_load and an equivalent parallel resistance from R2 at the tuned resonant frequency. In other words, given L2, R2, and RAC_load, an operating frequency can be selected to maximize the power coupling efficiency, provided that the chosen frequency is well below the coil self-resonance. It should be noted that the terms of
β / (1 + β ) 2 and 1/1 + β
enhancement effect due to the ferrite rods; (3) silicone package introduces an negligible effect; and (4) testing in saline solution causes a quality factor drop by 20% compared to the performance data obtained in air. Characterization of external coil reveals an inductance of approximately 5 µH with quality factor of 216, 224, 185, and 165 at 1MHz, 2 MHz, 3 MHz, and 4 MHz, respectively. The inductive coupling factor, k, between the siliconeencapsulated implantable coil and the external coil is then characterized as a function of gap size between the two coils, position and tilting angle of the implantable coil with respect to the external coil. Figure 6 shows the measured inductive coupling factor as a function of coils gap size and position of the implantable coil with respect to the external coil while maintaining a zero-degree tilting angle. As shown in the plot, the k factor varies from 0.3x10-3 to 0.4x10-3 with a 20 cm distance between the two coils, whereas with the gap size of 8 cm, the k factor varies from 1.9x10-3 along the external coil edge to 3.2x10-3 at the external coil center.
exhibit a variation within 10%
for β between 0.5 and 1.5.
Figure 6. Measured inductive coupling factor. Figure 5. Performance characteristic curves versus β.
III.
INDUCTIVE COUPLING CHARACTERIZATION
The silicone-encapsulated implantable coil enclosing a rechargeable battery and two ferrite rods, as shown in Figure 3, is characterized for its inductance value and quality factor in saline solution, emulating an in vivo environment after the implant. Measurement reveals an inductance value of approximately 11 µH below 4 MHz and a maximum quality factor of 37 at 3 MHz. Measurement data also indicates that (1) insertion of battery in the coil degrades the coil quality factor by approximately 50% while maintaining a nearly constant inductance value; (2) insertion of battery with ferrite rods results in a quality factor reduction of about only 10% while increasing the inductance value by 30%, indicating an inductance
It should be noted that the coupling factor is only a function of coils geometry and their relative position, and is independent of materials, for example battery and ferrite rods enclosed in the middle of the implantable coil, and testing medium such as air and saline solution. Characterization with a 30o of tilting angle between the coils reveals that the k factor varies from 0.2x10-3 to 0.3x10-3 with a 20 cm distance between the two coils, whereas with the gap size of 8 cm, the k factor varies from 1x10-3 along the external coil edge to 2.3x10-3 at the external coil center due to the direction of the magnetic flux with respect to the coils plane. To minimize the coupling factor variation due to coil tilting, it is envisioned that a number of external coils will be employed with overlapped regions in the final system design. For a constant charging current of 100 µA, which is adequate for the proposed application, an RF power of
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approximately 2.4 mW needs to be coupled into the implantable microsystem through the tuned coil loops with RAC_load of 7.5 kΩ. Based on the implantable coil characterization results, an optimal frequency of 3 MHz is required for maximizing the power coupling efficiency with a corresponding β value of unity. Figure 7 presents the worst-case calculated external power requirement as function of coils gap size and tilting angle. As illustrated in the plot, an external RF power of 7W is needed for a coils gap size of 20 cm with 0o tilting angle and 28W for the same distance with 30o tilting angle. As mentioned above, a number of overlapped coils will be considered to minimize the coupling factor variation, thus the external power requirement variation due to coils tilting. In addition, RF power level sensing inside the implant with an adaptive feedback and control of external RF power will be incorporated in the design to minimize external power dissipation and to ensure a reliable implant operation [15].
ACKNOWLEDGMENT This work was supported in part by the Rehabilitation Research & Development Service of the Department of Veteran Affairs. REFERENCES [1]
[2]
[3]
[4]
100 o
0 Tilt
[5]
o
Required External Power (W)
30 Tilt 10
[6] 1
[7] 0.1
[8]
0.01
0.001 0
[9] 5
10
15
20
Distance Between External and Implantable Coils (cm)
[10]
Figure 7. Required external power versus distance.
IV.
[11]
CONCLUSION
An optimized wireless power recharging system for implantable bladder pressure chronic monitoring application is presented. The wireless recharging system, consisting of an external 4-turn 15-cm-diameter powering coil and a silicone-encapsulated implantable spiral coil with a dimension of 7 mm x 17 mm x 2.5 mm and 18 turns, which encloses an ASIC, a 3-mm-diameter 12-mm-long rechargeable battery, and two ferrite rods, demonstrates the capability of coupling a required RF power into the implantable microsystem with an adequate external RF power over a large distance up to 20 cm. Ferrite rods are employed as critical components to ensure a high quality for the implantable coil to achieve a maximum power coupling efficiency at an optimal frequency.
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