Magnetically-powered Implantable Doppler Blood Flow Meter

Sai Chun Tang*, David Vilkomerson#, Tom Chilipka#

*Department of Radiology, Harvard Medical School, Brigham and Women’s Hospital, 75 Francis Street, Boston, MA 02115, USA. (sct@bwh.harvard.edu)

#DVX, LLC,

31 Airpark Road, Princeton, NJ 08540

Abstract— Implantable active devices use batteries for power, requiring expensive and unwieldy metal enclosures. Some such devices are re-charged via magnetic fields, but require external coils to be close to and well aligned with internal coils. We show here a different method: powering a device using segmented Helmholtz coils, allowing much greater leeway in internal coil location; and transmitting magnetically wide-bandwidth data. A proof-of-concept experiment demonstrates practicality.

Keywords- Diffraction grating transducer, Doppler blood flow meter, implantable device, segmented coil, wireless power transfer

I. INTRODUCTION

Transcutaneous wireless energy transfer using magnetic coupling has been used for implantable medical devices, such as bone growth stimulators, drug delivery systems [1], retinal prosthesis [2], and neuromuscular stimulators [3], for decades. In these devices, the external energy transmitting planar coil must be placed on the skin surface precisely aligned with the implanted receiving planar coil to optimize the power transfer; the coil separation should be less than 1-2 cm [4] or the received power will be attenuated significantly. Previously, we demonstrated a method [5] that wirelessly powers an implantable device deep in the body using a transmitting coil wrapped around the body coaxially with an internal selenoidal coil. With this coil arrangement, precise coil alignment is not necessary when the receiving coil is close to the transmitting coil plane. By adding another transmitting coil separated from the first by a distance equal to the coil radius, the structure becomes the well-known “Helmholtz coil”, where the magnetic field inside the body in the region bounded by the coils is substantially uniform; the receiving coil can be located virtually anywhere in that region without coil alignment [6]. However, the high impedance of a transmitting coil large enough to fit around the body or large limb requires excitation voltages at the high frequencies desirable for this use that are higher than 1 kV. This requirement for high voltage could be a serious problem in terms of patient’s safety and for manufacturing costs. As described in detail in soon-to-be-published papers, [7],[8], we found that dividing the transmitting coil into multiple segments by capacitors so that resonance is achieved, the voltage on the coil is significantly reduced and the power transfer efficiency improved. Here, we apply this coil segmentation method to

power an implantable blood flow monitoring device using only a low excitation voltage.

At a previous Ultrasonics Symposium [9], we outlined a “smart graft”, i.e. one that monitored by Doppler ultrasound flow through itself and communicated the results so falling flow could be treated before graft failure. This graft was powered by a pacemaker battery. We found that using an implant with a pacemaker battery increased the cost of such a device significantly, not only because of the expense of the pacemaker battery but the need for a hermetically sealed titanium case to contain the battery. The large amount of energy in a pacemaker battery, ~ 30,000 Joules, requires such an enclosure. The cost therefore of the device, and the requirement that the implant in its titanium container be connected via leads to the graft – surgeons hate such leads – made the resulting system not suitable for commercialization.

A magnetically-powered system would eliminate the large battery and its required titanium case and leads, and eliminate the special medical implant radio transceiver required in the earlier device. Such a system would cost much less and be more desirable and commercializable. We will present here such a magnetically-powered system for a Doppler ultrasound blood flowmeter used for a smart graft.

II. WIRELESS ENERGY TRANSFER

A. Energy Transmitting and Receiving Coils

The implantable Doppler blood flow meter is powered wirelessly using energy-transmitting coils wrapped around the patient’s limb where the device is implanted. The coils are placed in the Helmholtz configuration that provides an even magnetic field around the implanted device so precise coil alignment is not necessary. As shown in Fig. 1, there are two co-axial coils, each coil of 3 turns. The coil diameter is 20.6 cm (enough to fit around a very large leg) and the separation between the coils is equal to the coil’s radius, 10.3 cm, i.e. the Helmholtz configuration. Each coil is made of AWG 20 enameled copper wire, with 3 mm separation between the turns.

The operating frequency for power transmission is set to the 6.78 MHz industrial, scientific and medical (ISM) band. This frequency was chosen so that at the magnetic intensity levels needed to power the Doppler flowmeter, the temperature rise in tissue is less than one degree C. The measured coil inductance is

This work was partially supported by NIH/NHLBI HL071359

1622978-1-4799-7049-0/14/$31.00 ©2014 IEEE 2014 IEEE International Ultrasonics Symposium Proceedings

10.1109/ULTSYM.2014.0402

11 H, so the coil impedance is 469 at 6.78 MHz. To power the blood flow meter, the required transmitting current is 1.6 Arms. By Ohm’s Law, a simple coil would require a coil voltage of 750 Vrms, or 1.06 kVpk assuming the wire resistance is ignored. Such a high coil voltage requirement significantly increases the cost of the coil driving circuit. In addition, expensive high-voltage electrical insulation for both the driving circuit and the coil would be required. A previous study [7] demonstrated that dividing the coil into multiple segments using high-frequency capacitors can significantly reduce the coil voltage to a safe level. In another study [8], we showed that coil segmentation can eliminate the dielectric loss caused by the displacement current flowing through the coil insulation and the bobbin material. In the prototype here, the coils are divided into 24 segments, 4 segments per turn, by 24 capacitors. To tune the series coil-capacitor circuit to 6.78 MHz, the required total series capacitance value is 1218pF.

The impedance of the segmented energy transmitting coil was measured using an impedance analyzer HP4194A. Fig. 2 shows the magnitude and phase of the impedance between 6 MHz and 8 MHz. At the resonant frequency, the impedance is minimum and equals to 1.65 , which is much lower than the impedance of an unsegmented coil inductance. Thus, the required voltage for the segmented coil is 2.64 Vrms, much lower than that for its unsegmented counterpart, 750 Vrms.

Figure. 1 The energy transmitting and signal receiving coils.

Figure. 2 Measured impedance of the segmented energy transmitting coil.

The energy receiving coil is wound around a 7-mm artificial graft (i.e. as it will be in a real graft). The coil has 10 turns and is made of an AWG 20 wire. The separation between windings is 2 mm, and the coil length is 2 cm. The coil is air-cored, so is MRI compatible. Its measured inductance is 228nH. A capacitor network is connected to the coil, and the coil-capacitor circuit is tuned to 6.78 MHz to maximize the power transfer capability. Fig. 3 illustrates the receiving resonant network. While all the capacitors, Cr, C1 and C2, determine the resonant frequency, C1 and C2 forms a passive potential divider to step down the coil voltage to the desired level.

III. DIFFRACTION GRATING TRANSDUCER FLOWMETER

The diffraction-grating transducer (“DGT”) utilizes the principle of a diffraction-grating to produce beams at an angle perpendicular to its surface [9]. Using a DGT and a slab transducer (ST) as shown in Fig 4, allows the transducers to be placed in the wall of a graft, where it does not interfere with blood flow, and obtain a Doppler signal from blood moving through the overlap of their beams, shown dark in Fig.4. DGT’s can be made of piezoelectric plastics that are spin-coated or dip-coated onto thin polyimide film [10].

Figure. 3 Schematic of the energy receiver circuit.

Figure. 4 Side view of flowmeter.

Figure. 5 Flowsensor embedded in the graft wall.

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Electrodes define the transducer elements, rather than kerfs, so high frequency operation is possible. As blood scatters ultrasound as a Rayleigh scatterer, i.e. as f4, using 30 MHz ultrasound rather than 5 MHz ultrasound increases the signal levels by 64, or more than 30 dB. This, plus having no attenuating tissue between the transducer and the blood, allows very low drive levels on the DGT to produce good signal-noise performance. This low power requirement eases the demand on the needed magnetically transmitted power.

IV. IMPLANTABLE CIRCUIT

The implantable circuit consists of 3 main parts: (1) power receiver, (2) DGT driver, and (3) signal amplifier and transmitter as shown in Fig. 6. The power receiver converts the magnetic energy to electric energy, rectifies the high-frequency ac voltage to DC, and charges a tiny (3mm×9mm×10mm) lithium polymer battery with a capacity of 8mAh. (The stored energy is less than 200 Joules, less than a hearing aid battery, so the implant requires no pacemaker-battery type enclosure.) When the external energy transmitting coil is enabled, the battery will be charged and the switch, SW1, disconnects the rectifier from the load shortening the charging time as all the rectifier output current is diverted to the battery. The charging time is about 20 seconds. When the transmitting coil is disabled, SW1 will close, the rectified voltage regulated by a low-dropout (LDO) regulator, and the implanted circuit starts to operate. The transmitting DGT is excited by the DGT driver at 30 MHz. The driver contains a matching circuit to boost the transducer voltage to 10 Vpk-pk. The Doppler signal from the transducer receiver is amplified by a low-noise amplifier (LNA) with a gain of 43 dB. The amplified signal is transmitted to the external receiver via the coil driver and a 3-turn signal-transmitting coil. Both the energy receiving and the signal transmitting coils are wound on a 7-mm graft. The total current required by the implant is 60 mA; the battery, after being charged for 20 seconds, drives the implant for 5 seconds.

Figure. 6. A board diagram of the implantable blood flow meter.

Figure. 7. Demodulation circuit.

V. BASE STATION SIGNAL RECEIVER

A. Signal Receiving Coil

The signal containing the Doppler information produced by the implanted device is received by the external signal-receiving coil between the Helmholtz coils, shown in Fig. 1. The coil has 2 turns and a diameter of 20 cm. It is made of a coaxial cable with the shield grounded, to block unwanted signal coupled by electric fields, but open-ended to avoid circulating current that can significantly reduce the magnetic coupling. The coil is divided into 4 resonant segments by capacitors. Without segmentation, the self-resonant frequency of the coil is lower than the signal frequency (30 MHz). When the coil is segmented by series capacitors and tuned to the 30 MHz, the impedance of the coil segment is canceled by the impedance of the adjacent capacitor. Therefore, the resultant segment impedance, which comes from the ESR of the capacitor and the winding conductor, is much lower than the impedance between the center wire and the ground through the dielectric. The 30 MHz signal current will mainly flow through the center wire, and thus the current flowing through the parasitic capacitance between the center wire and the shield is negligible. As a result, coil segmentation for the signal-receiving coil effectively reduces the effect of the parasitic capacitor and increases the self-resonant frequency of the coil.

B. Demodulation circuit

As shown in Fig. 7, the Doppler signal from the receiving coil is amplified, multiplied by the sine and cosine of the transmit frequency, and low-passed filtered to obtain “I” and “Q” signals, “quadrature detection”. These signals are digitized, and the complex number I+iQ is placed in an array and an FFT formed; from the Doppler power spectrum found from the FFT the velocity of the blood is calculated – see [11].

VI. TESTING AND VERIFICATION

A. Implantable Device Prototype

A “proof-of-concept” magnetically powered implantable blood flow meter is shown in Fig. 8. The implanted device was implemented in two separate printed circuit boards (PCBs), occupying a total area of ~ 7 cm2. (As the coils require ~ 2.5 cm of the graft length, the remaining 10-20 cm length of a graft, with an area of 20-40 cm2 area, provides ample space for fabricating the electronics directly on the graft in the eventual clinical version.)

For a demonstration of the proof-of-concept, the wireless signal and energy transfer is simulated from inside a leg. As seen in Fig. 9, the implant is placed inside the leg-sized (20 cm) cylinder filled with liquid developed to have the same conductivity, permittivity, and permeability as human tissue in our operating range. The flow of blood-mimicking fluid is set by the difference in height of the two bottles seen above the graft-sized tube, seen at the lower-left of Fig. 9, in which is placed the flowmeter of Figs. 4 and 5. The Doppler signal is brought inside the “leg” by the long wires visible to be transmitted by the implant circuit to the receiver coil on the outside of the leg.

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coil

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Amp 

Comp A/D 

LPF  LPF A/D 

cos sin 

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Figure. 8 A prototype of an implantable blood flow meter; see Fig. 6.

The energy transmitting coils were energized at 6.78 MHz, with 1.6 Arms current that required 2.64 Vrms. In ~ 20 seconds, the battery was fully charged which operated the implant for 5 seconds.

The output Doppler spectrum shown in Fig. 10 shows a peak at 3327 Hz, which corresponds to a velocity of 22.6 cm/s [10]; the velocity calculated from timed volume flow through the tube was 24 cm/s.

VII. SUMMARY

We have demonstrated a proof-of-concept of a Doppler flowmeter powered by magnetic coupling and transmitting data via magnetic coupling. While the signal-noise ratio of the Doppler signal is less than would be required for clinical use, it demonstrates the possibility of implanted medical devices using safe magnetic coupling and data transfer.

 

Figure. 9. The demonstration “proof-of-concept” – see text.

 

Figure. 10 Doppler spectrum corresponding to 22.6 cm/s.

REFERENCES [1] W. Greatbatch, and C. F. Holmes, "History of implantable devices," IEEE

Engineering in Medicine and Biology Magazine, Vol. 10, No. 3, September 1991, pp.38-41.

[2] G. Wang, W. Liu, M. Sivaprakasam, and G. A. Kendir, “Design and Analysis of an Adaptive Transcutaneous Power Telemetry for Biomedical Implants,” IEEE Transactions on Circuits and Systems – I: Regular Papers, Vol. 52, No. 10, October 2005, pp.2109-2117.

[3] S. Y. Lee and S. C. Lee, “An Implantable Wireless Bidirectional Communication Microstimulator for Neuromuscular Stimulation,” IEEE Transactions on Circuits and Systems – I: Regular Papers , Vol. 52, No. 12, December 2005, pp.2526-2538.

[4] G. B. Joung, and B. H. Cho, “An energy transmission system for an artificial heart using leakage inductance compensation of transcutaneous transformer,” IEEE Trans. Power Electron., Vol. 13. No.6, November 1998, pp.1013-1022.

[5] S. C. Tang, F. A. Jolesz, and G. T. Clement, “A wireless batteryless deep-seated implantable ultrasonic pulser-receiver powered by magnetic coupling,” IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, Vol. 58, No. 6, June 2011, pp.1211-1221.

[6] R Puers, R Carta and J Thoné, “Wireless power and data transmission strategies for next-generation capsule endoscopes,” J. Micromech. Microeng., Vol. 21, No. 5, 2011.

[7] S. C. Tang, “A Low-Operating-Voltage Wireless Intermediate-range Scheme for Energy and Signal Transmission by Magnetic Coupling for Implantable Devices,” Special Issue on Wireless Power Transfer, IEEE Journal of Emerging and Selected Topics in Power Electronics, to be published.

[8] S. C. Tang and N. J. McDannold, “Power Loss Analysis and Comparison of Segmented and Un-segmented Energy Coupling Coils for Wireless Energy Transfer,” Special Issue on Wireless Power Transfer, IEEE Journal of Emerging and Selected Topics in Power Electronics, to be published.

[9] D. Vilkomerson and T. Chilipka, “Implantable Doppler System for Self-Monitoring Vascular Grafts,” Proceedings of 2004 IEEE UFFC International Symposium, IEEE Press, Piscataway 2005.

[10] J. M. Cannata, T. Chilipka, H-C Yang, et al, “A Flexible Implantable Sensor for Post-Operative Monitoring of Blood Flow,” J Ult Med 2012;31:1795-1802.

[11] Evans D and McDicken W, Doppler Ultrasound, 2nd Ed. Chapter 12, John Wiley and Son, Chichester (2000).

Energy receiving coil 

Signal transmitting coil 

Lithium polymer rechargeable battery

Coil driver 

Resonant capacitors 

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Energy transmitting coil 

Tissue‐equivalent fluid 

Energy transmitting coil driver 

Implanted device (energy receiving 

and signal transmitting coils) 

LNA and DGT driver DGT (inside the tube) 

Signal receiving coil 

Receiver amplifier Demodulator

Blood mimicking fluid flowing through the tube 

1625 2014 IEEE International Ultrasonics Symposium Proceedings