124

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. BME-23, NO. 2, MARCH 1976

Controlled External Powering of Miniaturized Chronically Implanted Biotelemetry Devices ROLAND KADEFORS

Abstract-Radio frequency energizing of implanted transmitters presents an interesting possibility to make feasible comparative measurements from sites within the body over very long periods of time. Disturbances in the energy transport may, however, cause the implant carrier frequency to change, which is seen as artifacts or drift at the demodulator end. Here, this problem was solved by feeding back the demodulator output level to control the amplitude of the external energizing oscillator. The closed loop modifies the frequency characteristic of the transmitter-receiver system in a predictable way. The artifacts due to implant movements can be minimized by careful application and by analysis of the field characteristics inside and outside the energizing solenoid, thus positioning the implant adequately.

INTRODUCTION I N biomedical measurements it is sometimes of interest to monitor signals from the internal of the body over long periods of time. Radio frequency (RF) telemetry presents one way of penetrating the skin without risk of irritation or infection. The use of implanted telemetry transmitters is, however, limited on account of the problem of energizing the implant. In previous papers from this group [1], [2], [3] the experiences with RF energy supply of an implanted frequencymodulated (FM) transmitter have been reported. It was shown [3] how different designs of the primary antenna coil affect the energy transport at various positionings of the primary coil relative to the pickup coil. The efficiency of a number of energy converters for powering the FM transmitter circuit was also studied, with the particular problem in mind of how to achieve a stable operating point for the transmitter at small disturbances in the efficiency of transmission of energy. It was concluded that sideways parallel orientation of the coils rather than coaxial orientation was to be preferred in order to ensure stability. In recent years we have gained experience with on-line RF energizing of transmitters for telemetry of EMG to monitor muscle activity, not only in myo-electric control systems [4], [5], [6] but also in long term studies on how muscles are affected by changes in their operating conditions. The present study was concerned with improvement of the stability of the overall system to a level where comparative measurements can be taken with intervals of months, or even longer. This goal was reached by stabilizing the operating point of the system by controlled powering of the implant [7]. In contrast to a previously suggested method [8], the control signal is derived from the transmitter itself. Manuscript received May 13, 1974; revised April 28, 1975. This work was supported by the Swedish Board for Technical Development. The author is with the Department of Applied Electronics, Chalmers University of Technology, Goteborg, Sweden, and the Department of Clinical Neurophysiology, Sahlgren Hospital, Goteborg, Sweden.

METHOD

The Closed-Loop System Assume that the signal to be telemetered has a lower limiting frequency f, $ 0, and an upper limiting frequency f2. Consider a system with radio frequency energizing according to Fig. 1. The FM detector output level u0 has a one-to-one relationship to the carrier frequency ft of the modulated signal within the frequency band considered. Disturbances in the energy transport caused by changes in the coupling factor of the two coils occur as a result of geometrical displacements and give rise to motion artifacts, perceived as drift or sudden jumps in uo. The problem is to stabilize uo against variations in the coupling factor. We shall investigate the unique opportunity to stabilize the operating point of the transmitter by controlling the amount of power transferred from the external energizing unit. Fig. 2 shows a block diagram of a complete system including the feedback loop. The reference value uref is compared with the actual value u0 at the output of the FM detector, and the integrated error signal t

t

uodt =f (Uref - u) dt 0

0

controls the amplitude of the energizing external oscillator. The feedback signal is low-pass filtered in the receiver and the summation circuit. Fig. 3 shows how linearized transfer functions can be associated with the blocks of Fig. 2. The open-loop transfer function yields

GOJ(s

=

1sK(K2K3K4 Klo+ p s(I + sTj) ( +sT2) s(I + sTj) (I +sT2)'

(1)

The system is stable as long as the loop gain Kloop = K1K2K3K4 is moderate; it is easy to show that the condition for stability is T1 + T2

(2)

T, T2 Skin

Energy converter

Energizing

oscillator

Output FM

receiver

/ / / /

FM transmitter Electrodes

/

Fig. 1. The principle of the externally powered system under study.

125

KADEFORS: CONTROLLED POWERING OF BIOTELEMETRY DEVICES DIFFERENTIAL AMPLIFIER

LOW PASS FILTER

INTEGRATOR

OSCILLATOR

Fig. 2. External powering of an FM transmitter for telemetry of bioelectrical signals: block diagram.

*Li0

Uref

Fig. 3. Transfer functions associated with the blocks of Fig. 2.

Transmission of Information It is clear that the closed-loop control system described will interfere with the transfer characteristics of the telemetry system. For example, a d.c. offset at the FM transmitter input will be compensated for by the control circuit. The inputoutput transfer function for the information yields (see Fig. 3)

G5(s)

Ks

K4 +

sT2

K4

KIK2K3

I + sT2 s(l + sT1) K4K5 (I + sTj)s

s(l + sTj ) (I + sT2) + K0oop( In our case T2 is the time constant of the receiving system, which includes the limitations due to restricted RF bandwidth and the characteristics of the demodulating circuit. Here, T2 is chosen much smaller than Tl, the smoothing filter time constant. It is evident that, as co

jG.,(jco)j

I1

2L

1

K4K5

+jiT2I

i.e., the high frequency asymptote of the transfer function has a slope of -6 dB/octave. At low frequencies, the characteristic is determined by a +6 dB/octave asymptote as long as Kloop is small; large loop gain causes a peak to appear which interferes with the low-frequency behavior. THE MOTION ARTIFACT Field Considerations Movements of the body are likely to displace the implant relative to the external energizing unit, and thereby to cause fluctuations of the transmitter center frequencyf,. The feedback loop will compensate for these changes at a rate settled by the transfer function [Eq. (3)], but more efficiently so if the implant is adequately positioned with respect to the external coil. In a previous report [31 the field strength around an external energizing cylindrical antenna coil of dimensions much smaller than the wavelength was calculated. The geome-

Z/

Fig. 4. The external, energizing solenoid with notations used in the field calculations. Case (a) along the z axis; case (b) in the plane of the shaded area in the coil center.

try is shown in Fig. 4. Two cases were studied: (a) along the longitudinal axis of symmetry, and (b) in a symmetry plane,

perpendicular to this axis. The derivations covered the situation outside the coil only. The case of a small implanted transmitter positioned inside the external coil has been found to be of considerable practical interest, and is accounted for as follows. In Case (a), the space angle Q2 under which the coil end surface is seen from the field point is replaced by 4r Q as the field point passes through the end of the coil [9]. Denoting the distance between the field point and the coil center point o, we find the following expression for the magnetic flux density: -

B

o

Bo 22 {

4

1

(

)}

(4)

the plus sign being valid within the coil. Here, Q22 and Q2 are the space angles of the coil end surfaces. The quantity Bo is the internal (constant) flux density of a very long coil with the

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, MARCH 1976

126

same number of ampere turns per meter as the coil under study. Fig. 5 shows how B_ varies according to Eq. (4) as the coil dimensions are changed. Each one of the curves starts at the coil center pertaining to the Lla value studied. In Case (b), the two space angles are Ql and 47r - S21, respectively. The expression for the magnetic flux density in the internal of the coil is

B, L/a 4

-

1

"I=Bo

a2L

0.2_

V(L2 +r2)3/2+l -B

2 (1 +1 L2) 2

+

0.1. -4a

(5)

2

In the outside, BZ = BO - BZ. Fig. 6 shows the curves Bz(r) according to Eq. (5) for various coil dimensions. It is seen that the flux density Bz assumes much higher values in the internal of the coil than outside. Movement artifacts are evoked by sudden displacements of the pickup coil relative to the energizing coil, either by movements of the body tissues in which the implant is embedded or by movements of the external unit on the body surface. Such perturbations can be translatory in character, i.e., the implant is subject to a parallel displacement, or rotational, involving rotation around an axis of inertia. Fig. 7 illustrates the various degrees of freedom on the assumption of a cylindrical coil aligned in parallel with an applied field B. It is convenient to choose cylindrical coordinates ¢, p, up for the representation. As seen in the figure, AO denotes tilting angle. It is evident that for reasons of symmetry, Aqp perturbations can be neglected. Considering the effect of moderate tilting, we assume rotations AO to be small. If the implant is properly aligned, the magnetic flux 4) encircled will be modified according to the cosine law (AG in radians): A4)

_Zt

(AG)2

2

4)

Hence, AO disturbances are small and will be neglected in the present treatise. The quantities Al and Ap to be considered are not small in relation to the dimensions of the pickup coil; in fact, they may be of the same order of magnitude as the dimensions of the external coil. The pickup coil is assumed to be of small dimensions compared with those of the external coil and is thereby exposed to the field strength variations depicted in Figs. 5 and 6. From the point of view of movement artifacts due to translatory displacements Ai and Ap, the conditions for best immunity in the two cases (a) and (b) defined above are

dBz=0 dz

dBz

0.

seen

-3a

Il

-a

-2a

coil endpoint

a

2a

3a

4a

d

Fig. 5. Variations of the magnetic flux density in case (a) at various geometrical proportions of the solenoid, according to Eq. (6). Each curve starts at the coil center. Bz L/a

2

Bo 2

0.5

0.1

B

4

Fig. 6. Variations of the magnetic flux density in case (b) at various geometrical proportions of the solenoid, according to Eq. (7). Each curve starts at the coil center, changes sign at the surface, and extends into the outside.

Fig. 7. Various classes of displacements of the receiving coil discussed in connection with movement artifact generation.

coil dimensions in the center point. Condition (6a) is also met

(6a) for any r in the plane studied. It is found that large derivatives (6b)

dr

As

Q

BO 0.5>

B =B (1-

Bo

from Figs. 5 and 6, both conditions are met for any

occur, particularly along the z axis at the coil endpoints. This area, as pointed out previously [3], is undesirable for energy

transfer from the movement artifact point of view. It is evident that the vicinity of the center point position for the implant is the most suitable one. The internal of the coil is, on

127

KADEFORS: CONTROLLED POWERING OF BIOTELEMETRY DEVICES

the whole, because of the moderate B, derivatives and the large flux densities, more advantageous than the outside. The field calculations also show that displacements along the axes of symmetry cause the flux density to which the implant is exposed to modify depending upon the Lla value, i.e., the coil proportions. The variations within the coil are small at high Lla values; this holds in both cases. The intemal flux density of a very long coil is constant except for close to the endpoints. It is obvious that at low Lla values there is a very high susceptibility to motion artifacts caused by displacements in the r direction. The Movement Step Response Eq. (3) can be used to calculate the shape of the artifact evoked by sudden displacements. We consider a unit step function g(t) defined by

g(t) = 1,

(t > )

g(t)0=,

(t> T2, the stability criterion (2) yields K100p < 1/T2 ; 1.6 104. The diagram also contains an example of the characteristics of an implemented system designed to monitor muscle activity during changes in the operational state of the muscle over periods of time of several months. The flux density plots strongly motivated positioning of the implant inside the coil. In this application it was impractical for experimental reasons with a coil of larger L/a value than unity. Of the cases studied in Fig. 9, an analysis of the equation D = 0 gives a step response of the form given by Eq. (10) for Kloop = 10 and 15; the remaining two cases Kl0.p = 20 and 30

128

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, MARCH 1976

POWER UNIT

IMPLANT

RECEIVER

Fig. 8. Circuit diagram of the entire system. Time constants, etc., are given in the text.

Fig. 9. A set of curves calculated from Eq. (3) showing the magnitude of the input-output transfer function at various Kioo values. Circles: a practical case. T, = 15 ms, T2 = 60 ,s.

Fig. 10. Step responses corresponding to the cases shown in Fig. 9 and calculated from Eqs. (8) and (9).

give oscillatory responses. Fig. 10 shows the calculated step responses in the four cases studied. The initial response is fast, followed by a decaying wave, slower for low Kloop values. The overshoot is hardly visible in the gain range studied. It should be emphasized that in the practical measurement situation, the disturbance is, in general, not well represented by the step function g(t). Artifacts due to voluntary muscle movements are slower, with a risetime rarely exceeding 100 ms. Fig. 10 implies that such artifacts are well canceled by the control loop, which agrees with practical experience. Further optimization of the control loop parameters has not been called for. In Fig. 11, the response in implant telemetry of EMG at sudden activation of the biceps muscle is shown with and without controlled powering. Note the large movement artifact efficiently eliminated in the modified system. DIsCUSSION It can be argued that the system performance as characterized in Eq. (3) would be possible to arrive at through simple high-pass filtering at the receiving end. Such processing, however, leaves important aspects uncovered. Motion artifacts or drift, whenever large enough, will cause fo to exceed the dynamic frequency range of the tuned receiver and detector; the output signal becomes noisy and changes its amplitude as the detector limiter passes into saturation. Automatic frequency control (AFC) is frequently employed to lock a receiver to the center frequency of the transmitter concerned. This technique can also be applied here. It is, however, undesirable to allow the center frequency ft to drift more or less freely. The system described here allows external locking of fo to a frequency suitable for transmission, for example, from the point of view of avoiding interference with other transmitters. Eqs. (4) and (5) and corresponding figures (5 and 6) show the general appearance of the flux density variations in the

Fig. 11. EMG signals derived (left) without, (right) with controlled RF powering. M. biceps brachii, sudden voluntary activation.

two cases studied. In a practical measurement situation, the characteristics may be somewhat modified. For example, close to the solenoid winding, the field will be distorted due to the local effect of the individual turn. Introduction of high magnetic permeability core material changes the characteristics in a more or less pronounced way. With these reservations, the calculations have been found sufficiently accurate for acting as a basis for design and interpretation of experiences with practical systems. The implications are in accordance with favorable results with positioning the transmitter inside the extemal coil reported by Tucker and Scott [10]. Further improvements in the stability and artifact cancellation are undoubtedly possible, for instance, by using a three-

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. BME-23, NO. 2, MARCH 1976

129

dimensional pickup arrangement [8], [II] at some sacrifice ACKNOWLEDGMENT of the miniaturization and simplicity of the implant. Series The author is indebted to Mr. Sterner Hedberg, who has perand shunt power regulation present alternative methods for formed skillful technician work on the practical design of the the system design. hardware system. The advantages are gained at the expense of d.c. signal transREFERENCES mission capability. However, d.c. signals should not be transand I. Petersen, "Telemetry of myoC. Hirsch, E. Kaiser [11 mitted directly by means of RF powered FM transmitters. Acta Orthopaed. Scand., vol. 37, pp. 156-165, 1966. potentials," Subcarrier technique, which is recommended in such cases, is [2] C. Hirsch, R. Kadefors, E. Kaiser and I. Petersen, "An implantperfectly compatible with the method of controlled powering. able micro-circuit for long-term telemetry of myo-potentials," Digest of the 7th ICMBE (Stockholm, August 1967), p. 365. It is noted that the feedback control system in fact makes a [3] R. Kadefors, E. Kaiser and I. Petersen, "Energizing implantable virtue of a deficiency of the implant circuit: its vulnerability transmitters by means of coupled inductance coils," IEEE Trans. to disturbances in the energy supply. Of course, improved Bio-Med. Eng., vol. 16, pp. 177-183, 1969. [4] P. Herberts, R. Kadefors, E. Kaiser and I. Petersen, "Implantation oscillator stability makes the demand for controlled powering of micro-circuits for myo-electric control of prostheses: clinical less acute. The implant circuit of Fig. 8 has a high sensitivity aspects," J. Bone & Joint Surgery (British), vol. 50B, pp. 780to changes in the efficiency in energy supply, which is re791, November 1968. [5] R. Kadefors, The voluntary EMG in prosthetics. (Res. Lab. Med. flected as a high K3 value. An improved stability will lower Electr., Chalmers Univ. Techn., Goteborg, 1970). K3, and thereby the loop gain. The system's properties can [6] R. Kadefors, Control components in rehabilitation engineering. be maintained, if so desired, by increasing K1, K2 or K4 (Res. Lab. Med. Electr., Chalmers Univ. Techn., Goteborg, 1971). adequately. On the other hand, it should be emphasized that [71 R. Kadefors, "Closed-oop techniques offer advantages in RF powering of implanted transmitters," Digest of the 3rd ICMP a more stable oscillator circuit to replace the one investigated (Goteborg, 1972). should be able to operate on a supply voltage around 1 V, with [8] E. Yon, A study of R.F. power supply for implanted transmithigh modulation sensitivity and prospects for extreme minters. (Report No. EDC 3-62-1, Case Western Reserve Univ., Cleveland, Ohio, May 1962). iaturization, otherwise stability will be gained at the expense E. Hallen, Electromagnetic Theory. New York: Wiley, 1962, [91 of sensitivity and physical size of the implant. pp. 155-156. The system described has been found reliable and, in prac- [101 F. R. Tucker and R. N. Scott, "Development of a surgically implanted myo-telemetry control system," J. Bone & Joint Surgery tice, useful. It is believed that the principle could be em(British), vol. SOB, pp. 771-779, November 1968. ployed particularly in acquisition at the skin surface of various [11] W. H. Ko and M. R. Neuman, "Implant biotelemetry and microbioelectric signals from deep sites. electronics," Science, vol. 156, pp. 351-360, 1967.

A High-Power Gastric Photocoagulator for Fiberoptic Endoscopy DAVID C. AUTH, MEMBER, IEEE, VINCENT T. Y. LAM, RAYMOND W. MOHR, STUDENT MEMBER, IEEE, FRED E. SILVERSTEIN, AND CYRUS E. RUBIN

Abstract-A single quartz fiber delivery system capable of remotely depositing in excess of 9 watts of continuous laser radiation has been developed and tested. Of particular interest in this development was the achievement of a catheter that can be inserted into the biopsy channel of a conventional flexible endoscope for noninvasive control

Manuscript received November 5, 1974; revised May 16, 1975. This work was supported in part by the NIH under Grant GM-16436. The work of C. E. Rubin was supported by the NIH, USPHS under Career Research Award CA 03499 and under Research Grant NIAMDD16059. The work of F. E. Silverstein was supported by the University of Washington Alcoholism and Drug Abuse Institute and under a Veterans Administration Research and Education Grant. D. C. Auth and R. W. Mohr are with the Department of Electrical Engineering and the Center for Bioengineering, University of Washington, Seattle, WA 98195. V. T. Y. Lam was with the Department of Electrical Engineering, University of Washington, Seattle, WA 98195. He is now with Pacific Northwest Bell, Seattle, WA. C. E. Rubin and F. E. Silverstein are with the Department of Medicine, University of Washington, Seattle, WA 98195.

of gastric bleeding. In animal testing, the prototype device was found capable of producing a large useful area of superficial tissue coagulum, effectively controlling gastric hemorrhage. Future clinical application is highly encouraging.

INTRODUCTION DECENT advances in laser fiberoptic technology coupled IIX.with significant clinical experimentation [1]-[5] suggested the application of visible laser radiation to the control of gastric hemorrhage. Considerable clinical experience [4] has accumulated in the use of argon lasers in ophthalmological procedures involving photocoagulation. By scaling the power requirements of ophthalmological practice and considering the power used in previous CO2 laser coagulation [1] studies, an optical power level of about 10 watts appears appropriate for useful clinical arrest of bleeding. The substantial improvement in flexible fiberoptic endoscopes for inspection, diag-