144 IEEE TRANSACTIONS ON ELECTRON DEVICES, VOL. ED-29, NO. 1, JANUARY 1982

response to RFP-NIH-NINCDS-80-05,1980. p. 1373. See also, P. Shah, “High performance, high density [12] P. Clark and R. Hallworth, “A multiple electrode array for a MOS processing using polyimid interlevel insulation,” in Proc.

cochlea implant,” J. Laryngology and Otology, vol. XC, no, I , 1976.

IEEE 23rd Electric Components Conf. [I41 H. Schonhorn, “Surface modification of polymers for adhesive

M. Sonn and W. Feist, “A prototype flexible microelectrode bonding,” in Polymer Surfaces, by D. Clark and W. Feast, Eds. array for implant-Prosthesis applications,” Med. Biol. Eng., New York: Wiley, 1978. p. 778, Nov. 1974. [15] K. Chopra, Thin Film Phenomena. New York: McGraw-Hill,

[ 131 Developed by Dupont in 1965. See J. Polymer Sci., pt. A, vol. 3, 1969, p. 311.

Thin Linear Thermometer Arrays for Use in Localized Cancer Hyperthermia

Absfract-A thin linear array of silicon diodes has been developed to measure one-dimensional temperature profiles in tissue during treat- ment of cancer by localized hyperthermia. The array is composed of six discrete diodes on three flexible stainless steel wires, with a maximum cross-sectional dimension of 0.5 mm, so that it can be introduced into a tumor area via a small puncture wound. Temperature data are ex- tracted using an external electronics system under microprocessor con- trol; the present overall accuracy of the system is 0.2OC over the range of 37-45’C. The array has been tested in ultrasound, RF, and micro- wave heating fields. Computer simulation shows this array to be non- perturbing of the thermal field in tissue, so that it can provide accurate temperature data. Development of a batch-fabricated array of twenty diodes on five leads is under way. A monolithic silicon diode array is shown to produce large temperature perturbations because of its high thermal conductivity, while an alternative technology using silicon micromachining and adaptations of existing techniques for lead fabri- cation should produce an array of low thermal conductivity which can obtain accurate measurements. The present and future arrays should also be suitable for data collection in many in-vivo situations other than hyperthermia.

L I. INTRODUCTION

OCALIZED HEATING (hyperthermia) of cancerous tissue has been shown by several investigators to produce shrink-

age of tumors when used as the sole treatment modality and when used in conjunction with either chemotherapy or ionizing radiation [ l] - [ 6 ] . Hyperthermia is accomplished using ultra-

Manuscript received June 3, 1981; revised September 3, 1981. This work is part of the research of the National Resource for Silicon Bio- medical Transducers, sponsored by the NIH Division of Research Re- sources under Grant NIH-RR-01086.

The authors are with the Department of Electrical Engineering, Stan- ford University, Stanford, CA 94305.

sound, RF, or microwaves for deep heating of tissues, often in conjunction with heating or cooling of the skin to produce a peak temperature at a given depth below the surface of the body. One of the basic problems in hyperthermia is determin- ing the amount and extent of temperature rise in tissue. Great variations in tissue structure, vascularization, and blood flow occur from individual to individual, at different body locations within the same individual, and at the same body location in an individual over the course of time; as a result, identical treatment regimens with a given heating apparatus can produce wildly different temperature profiles and unrepeatable results.

Successful use of hyperthermia in a clinical setting will re- quire feedback in the form of temperature measurement, at several points within the tumor and in the tissue surrounding the tumor. Ideally, such temperature measurement should be noninvasive, so that no disturbance of tissue occurs due to the temperature measurement. In practice, some minimal amount of tissue invasion is necessary if temperature is t o be measured at several localized interior points. At a minimum, a single puncture wound is necessary. If several “thermometer points” can be located along the length of such a puncture, a practical optimum will be achieved.

This paper discusses the development of a miniature linear thermometer array composed of silicon diodes on flexible wires which will provide a near-optimal means of temperature measurement during hyperthermia. This array should also prove suitable for use in many other in-vivo situations. For example, temperature profiles through the heart wall and on the surface of the heart in a normal experimental animal are of interest to cardiologists, and should be obtainable using arrays of this type.

0018-9383/82/0100-0144$00.75 0 1982 IEEE

BARTH AND ANGELL: LINEAR THERMOMETER ARRAYS IN LOCALIZED CANCER HYPERTHERMIA 145

11. DIODE MULTIPLEXING TECHNIQUE The current-voltage characteristic of a semiconductor p-n

I = Z , X [exp(V/V,)- 11 (1) where I is the current through the diode, Vis the voltage across the p-n junction, I, is the “saturation current” (defined as the extrapolated value of I at V = G), and V, = kT/q is the ‘.‘ther- mal voltage,” where k is Boltzmann’s constant, T is absolute temperature, and q is the magnitude of the electron charge. Temperature dependence in the I-V characteristic arises from both the I, and V, terms [7] . For a silicon diode under forward bias near room temperature, the temperature coefficient of the diode voltage at constant injected current has a value of - 2 mV/”C, while the temperature coefficient of the diode cur- rent at constant applied voltage is 7 percent of I/”C. The lin- ear thermometer array employs a constant injected current, and uses the voltage temperature coefficient to measure temperature.

Fig. 1 illustrates the configuration in which diodes are con- nected in the array. Current sourcing and sinking, along with voltage measurement, are performed on a given pair of wires, while the remaining wire floats electrically. For a given direc- tion of current on a given pair of wires, only one diode is strongly conducting. For example, if current is sourced to lead 1, and sunk from lead 2, diode A will be conducting with a forward voltage drop of 0.7 V, diodes B, C, and F will be reverse-biased and nonconducting, and diodes D and E will be forward-biased at a voltage of only 0.35 V. Because of the exponential I-V characteristic of the diodes, the current flow- ing through diodes D and E will be on the order of times the current through diode A . As a result, almost all of the temperature coefficient seen on leads 1-2 comes from diodeA.

A temperature profile is obtained from the thermometer array by sequentially biasing pairs of leads and monitoring the voltages across those leads. The multiplexing technique can be extended to any number of wires, wi thN X ( N - 1) diodes on N wires. This technique provides the smallest number of wires per thermometer point, when many thermometer points are present, of any available technique using passive sensors. For example, sensors such as thermistors and thermocouples re- quire a common lead plus one lead for each sensing element. (The incorporation of active multiplexing electronics in a ther- mometer array, as suggested by Jerman [B] , could permit many thermometer points to be placed on only two leads; barring the use of such a technique, the passive technique used here results in the smallest number of leads per sensor.) The small number of leads in the array discussed here permits its cross section to remain as the number of thermometer points increases, providing a size advantage over competing techniques.

111. ARRAY FABRICATION

diode is given by the familiar diode equation

The fabrication procedure for the thermometer array consists of three major steps: diode fabrication, attachment of diodes to wires, and incorporation of the wires into the final lead assembly. These steps are discussed below.

The individual diodes in the array are fabricated using silicon batch fabrication techniques, as shown schematically in Fig. 2.

3

Fig. 1. Schematic of the linear thermometer array.

CONTACT ALUMINdM

(c) ( 4 Fig. 2. Fabrication process for a silicon diode used in the thermometer

array. (a) Chemical micromachining. (b) n-type dopant diffusion. (c) Contact holes opened in silicon-dioxide layer. (d) Aluminum de- posited and etched to form bonding pads.

The starting material is a wafer of p-type (100)-oriented silicon. In the area where a diode will be formed, chemical micro- machining is performed using KOH in water as an orientation- dependent etch and silicon dioxide as an etch mask. The result is a set of grooves on the silicon surface which will later be use- ful to mechanically alignit to the stainless steel wires (Fig. 2(a)); in addition, a bevel is formed around the edge of the diode. Next, an n-type region is selectively diffused into the wafer to form the p-n junction (Fig. 2(b)). Two contact holes are opened in an insulating layer of silicon dioxide on the wafer surface (Fig. 2(c)); these holes provide electrical contact to the diode. Next a layer of aluminum is deposited on the wafer surface and is selectively etched to form three bonding pads (Fig. 2(d)). Two of these bonding pads make electrical contact to the diode, while the third is just a mechanical bonding point for subse- quent wire attachment. Finally, the wafer is sawed apart into individual diodes. In order to provide six diodes on three wires as illustrated in Fig. 1, three types of diodes are necessary; by orienting these three types in both “forward” and “backward” configurations during wire attachment, the six electrical con- nections are obtained.

The wires attached to the diodes are 25-pm diam stainless steel, plated with a thin layer of nickel and a 6-pm layer of gold. These wires are prestrung on a frame about 30 cm long. An exact 100-pm spacing is maintained between the wires by stringing them over two diodes, with their micromachined grooves 100 pm apart, which are placed at the ends of the frame. The diodes to be incorporated into the array are then attached to the strung wires by thermocompression bonding of the gold plating on the wires to the aluminum pads on the diodes. A view of an individual diode attached to wires is shown in Fig. 3. The bevel at the edge of the diode, which was formed during the chemical micromachining operation, and which is insulated by a layer of silicon dioxide, ensures that the

146 IEEE TRANSACTIONS ON ELECTRON DEVICES, VOL. ED-29, NO. 1, JANUARY 1982

Fig. 3. Silicon diode attached to stainless steel wires.

P CII B

F -$ 1 ,

r D i C

Fig. 4. Typical lead assembly. (A) Connector. (B) Three-conductor cable. ( C ) Epoxy block over wire joints. (0) Gold-plated wires with parylene coating. ( E ) Silicon diodes, 5 mm apart. Wire lengths are drawn shorter than actual scale.

wires do not contact bare silicon so that a potential shorting problem is avoided.

After attachment of the diodes, the wires, diodes, and frame are coated with a 30-pm-thick layer of Parylene (a vapor- deposited biocompatible water-resistant polymer). The wires are then cut free of the frame for incorporation into the final lead assembly. The exact structure of the final lead assembly depends on the situation in which the array is used. Such assemblies must be durable enough to withstand use, but should not interact excessively with the hyperthermia treatment field. A typical lead assembly is shown in Fig. 4. The thin stainless steel wires are joined to thicker wires by silver paint over a twisted-wire joint. This joint, and all cables, are electrically insulated by an additional 15-pm layer of Parylene; the joint is then potted in epoxy for mechanical stability. Leaving the array exposed at the leading edge of the assembly as shown in Fig. 4 provides several advantages: minimum cross section, fastest thermal response, mechanical flexibility, and minimum interaction with the hyperthermia treatment field. In this configuration, the array may be inserted into tissue by insert- ing an 18-gauge hypodermic needle from the far side of the tumor area, threading the array into the open end of the needle, and removing the needle while the array stays in place. The array is strong enough (breaking tension 350 g) to be subse- quently pulled out of soft tissue without breaking.

The lead assembly can be strengthened by sheathing the array in a hypodermic needle or a Teflon tube. This needle or tube can be anchored in the epoxy block which covers the twisted wire joints, providing a durable structure suitable for use in muscle and connective tissue. Selection of the proper sheath material depends partly on the treatment method; a steel needle may interact excessively with a microwave field, while Teflon is a strong ultrasound absorber and could produce excess heat- ing of the array in an ultrasound field.

The major problems with the assembly technique described above are fabrication time and yield. Fabrication time per array is on the order of 10 h, while yield is low because the operator must make a total of 18 thermocompression bonds per array with no opportunity for reworking mistakes. In

addition, the temperature profile obtainable with a six-ther- mometer array is relatively coarse. The multiplexing technique discussed above is readily extendible to twenty diodes on five leads, but such an array cannot be fabricated with any useful yield by the techniques discussed above.

The development of batch fabrication techniques for fabrica- tion of a twenty-diode thermometer array is a goal presently being pursued. As an “existence proof,” a monolithic silicon array has been fabricated, incorporating twenty diodes in a silicon sliver with a 550-pm X 375-pm cross section. This rigid sliver sacrifices the flexibility of the six-diode array, but gives a finer temperature profile. In addition, it can be fabri- cated with good yield in a relatively short time per array, How-‘ ever, the thermal conductance of the silicon sliver is large and, a shown below, computer simulation predicts that this sliver can seriously perturb the temperature field in living tissue.

Further development of batch fabrication techniques is pres- ently under way. These techniques involve the use of silicon micromachining to leave single silicon diodes, etched from a monolithic starting wafer, attached to flexible, insulated wires of low thermal conductivity. The technology for batch fabri- cating strong, flexible lead assemblies has been previously developed [9] , and should be readily adaptable to incorpora- tion into a silicon diode array. The arrays fabricated by such technology should result in accurate measurement, a fine tern- perature profile, and high yield.

IV. ANALYSIS OF ELECTRICAL CROSSTALK BETWEEN DIODES

The resolution and accuracy of the multiplexing technique used for the linear thermometer array is ultimately limited by electrical crosstalk between the diode which is addressed by the injected current on a given lead pair, and the diodes else- where in the array which are partially forward biased by that same injected current. An analysis of such crosstalk in the six- diode array, combined with experimental diode characteristics, shows that the magnitude of the crosstalk is well below the 0.loC resolution needed for most biomedical applications.

The I- V characteristic given by (1) is an oversimplification for real diodes. A better approximation, for low to medium current densities, is given by

I = I s m X [exp(V/K)- 11 + I s l X [exp(V/25)- 11 (2)

where I,, = I , is the extrapolated saturation current for the medium-current regime, and Isl is a second saturation current, defined by extrapolation of the low-current characteristic in a manner similar to that by which I,, is defined by extrapola- tion of the medium-current characteristic.

Referring to Fig. 1, again assume that diode A is addressed by injecting a current I,, to lead 1 and sinking that current from lead 2. Current through the back-biased diodes is approxi- mately zero, while that through the half-forward-biased diodes D and E is approximately

I& = Isl X exp [( V/2)/(2 5 > ] . (3)

The overall current is then

I12 = I a t Ide = constant (4)

BARTH AND ANGELL: LINEAR THERMOMETER ARRAYS IN LOCALIZED CANCER HYPERTHERMIA 147

where I , is the current through diode A and is the current through diodes D and E. Crosstalk occurs when a change in temperature at diode D or E changes the current through diode A . The worst case situation occurs when a uniform tempera- ture change Tde occurs in both diodes D and E. In this case

aIa/aTde = - aIde/aTde. ( 5 )

The electrical "crosstalk coefficient" is defined as the false temperature change computed from the current change through diode A as Tde changes, divided by the actual temperature sensitivity of diode A . This coefficient can be expressed as

c x e = (aIdaTde)l(a&laT2. (6)

As noted above, the temperature coefficient of current for a silicon diode can be expressed as a percentage change in the current per change in temperature. For diode A

ar,/aTa = K , x I, (7)

where K, is the temperature coefficient in the medium-current regime, Similarly, in the low-current regime

aIde/aTde = KI X Idea (8)

The crosstalk coefficient thus becomes

C x e = - (KI x Ide)/(Km x Ia). (9)

The low-current temperature characteristics of silicon diodes are not well understood at present [ 101 ; making a worst case assumption that Ki = 2 X K, gives

e = - 2 x &&Ira. (10)

The measured currents through the silicon diodes incorpo- rated into the linear thermometer array are 100 pA at 0.7 V, and 40 nA at 0.35 V. The injected current I,, used for the array is 100 PA. The value of the electrical crosstalk coefficient is thus

e,, = - 8 x ("cl"c). ( 1 1)

This value of C,, implies that there is essentially no inter- action between diodes down to the 0.loC accuracy required for biomedical applications. In principle, Td, could change more than 100°C without causing a noticeable change in the temperature computed for diode A .

The above analysis is readily extendible to arrays of more than six diodes. In such arrays, there will be additional diodes which are forward biased at half of the voltage across the ad- dressed diode. The worst case crosstalk coefficient willincrease with each lead added to the array, but the magnitude of cross- talk will remain below any practically troublesome value in the twenty-diode arrays presently under development.

V. TEMPERATURE INACCURACIES ARISING FROM ATTACHED WIRES

In the present thermometer array, the temperature coefficient of resistivity (TCR) of the attached wires is the limiting factor in accuracy of temperature measurement. The measured linear resistivity of these gold-plated stainless steel wires is rI = 3.6 Qlcm at 27"C, and the measuked TCR is 1.3 X 10-3/"C.

A general expression for the inaccuracy in temperature resulting from changes in the temperature of wires attached to a given diode is

Xd ATw = 2(kTw/STd) I r l J [T(x) - TO1 dx (1 2)

X 0

where k ~ , is the TCR of the wires,, STd is the temperature coefficient of voltage of a diode, I is the current through the wires, T(x ) is the temperature at a given point x along the wires, To is the reference temperature at which rz is evaluated, x. is the base end of the thermometer array, and xd is the location of the diode along the array. The factor of 2 in ( 1 2 ) accounts for current flow to and from the diode. Both k t , and rl are assumed to be constant along the length of the wires, and so are pulled out of the integral expression in x.

Using kTw = 0.0013/"C, ST^ = - 0.002 V/"C, I = A, rI = 3.6 !il/cm, To = 27"C, x,, = 0, x d = 10 cm, and making the worst case assumption that T(x) = 45°C for all x, the spurious temperature change calculated is AT, = - 0.08"C. In more realistic situations than this worst case situation, AT, will be smaller by a factor of two or more.

VI. MODELING OF THERMAL PERTURBATION IN TISSUE

A question to be answered in assessing the utility of any temperature measurement technique is whether the technique perturbs the thermal field, and if so to what degree. Intuitively, it would seem that a small array composed mostly of thin stainless steel wires would have low thermal conductance along its length and would, therefore, be nonperturbing. However, if such an array is encased in a steel needle for added strength, its overall thermal conductance increases and it becomes more perturbing. A quantitative assessment of the degree of pertur- bation requires a three-dimensional computer simulation. Such a simulation has been developed, modeling a thermal conductor in a heated homogeneous piece of material. This simulation is based on a one-dimensional analytical model [ 1 1 1 , and incor- porates the following features:

1) Steady-state temperature is calculated by a relaxation method [ 121.

2) Temperature is calculated in a cylinder of tissue centered around a linear thermal conductor. The radial symmetry of this situation permits calculation of the three-dimensional temperature using a two-dimensional grid.

3) Skin temperature is fixed (i.e., the boundary condition at one end of the cylinder is a constant temperature).

4) Hyperthermia is simulated by a heat source with a char- acteristic exponential decay length in axial direction of the cylinder.

5 ) The homeostatic effects of blood flow are simulated by a thermal resistance from each grid point to a 37°C temperature.

6) The axial length and radius of the cylinder are set large enough to be considered semi-infinite for practical purposes.

Fig. 5 depicts a typical grid point in the array. The solutions generated use a grid of 21 X 11 points, with an over-relaxation parameter used to speed convergence. The convergence crite- rion is set so that the sum of the absolute values of the tem-

148 IEEE TRANSACTIONS ON ELECTRON DEVICES, VOL. ED-29, NO. 1 , JANUARY 1982

Fig. 5 . A point in the grid used for the computer simulation.

perature differences from the previous iteration at each grid point does not exceed 0.001’C. Typically, 50-70 iterations are required to obtain a solution starting from a previously stored “guess” temperature field. The model gives good agree- ment with the analytical model of [ 1 I ] for the case of unper- turbed tissue (i.e., tissue with no thermometer array in place).

Fig. 6 shows two temperature fields generated by this com- puter model. The fitst (Fig. 6(a)) is in homogeneous tissue which is heated by a source with a characteristic exponential decay length of 24 mm (this length corresponds to the decay length in tissue of a microwave source at 9 15 MHz [6], but the source could equally well be an ultrasound generator). The important points to note in this unperturbed temperature field are that the temperature gradient in the radial direction is zero, and that the peak temperature occurs at a point below the sur- face. The position of this peak is determined by thermal con- duction from the tissue to the 20’C “skin,” in combination with the heating field, the thermal conductivity of the tissue, and the blood flow characteristics of the tissue. The parameters used for this simulation are of the same order of magnitude as those used in [I 1 3 . \,

Fig. 6(b) illustrates a highly distorted temperature field, which would be produced by inserting a rigid silicon thermom- eter array inside an 18-gauge stainless steel hypodermic needle into the tissue during hyperthermia. The main source of the distortion is not the needle, but the silicon: silicon has a ther- mal conductivity of 150 mW/mm/’C [14]. With a thermal conductivity approximately 3 that of copper, the silicon needle constitutes an intrusive “heat pipe” which conducts A large heat current to the “skin” and a smaller heat current to the interior of the model. The distortion introduced has two im- portant effects: it lowers the maximum temperature measured from 44.98 to 41.24’C, and it changes the position at which this maximum is measured by 1.5 cm. It should be emphasized that this is a real perturbation of the temperature, as opposed to an inaccuracy in measurement.

Table I summarizes the effects on the temperature field of the flexible six-diode array, hypodermic needles of various sizes, and of the rigid twenty-diode array. The criterion of perturbation used in the table is the temperature change pro- duced at a depth of 1.5 cm, as opposed to the change in maxi- mum measured temperature. It can be seen that the flexible array should be less perturbing than even the smallest available

T(Temperature)

R(Distance from axis)

)

20 c Z(Depth into tissue)

(a)

T(Temperature)

(R,Z,T)=(5cm,l.5cm,44.Q8C) / R(Distance from axis)

20 c Z(Depth into tissue)

(b)

Fig. 6 . Simulated temperature fields in a tissue model during hyper- thermia. (a) Best case; no thermometer array inserted, no perturba- tion. (b) Worst case; silicon sliver inside 18-gauge needle inserted into center of tissue. Note warping of entire temperature field near axis (Y = 0). Parameters: k (thermal conductivity of tissue) = 0.42 mW/ mm/’C; ?&d = 37’~; TSkin = 2 0 ” ~ ; DZ = Dr = 5 mm (grid spacing); Input power = 5 mW/mm2; decay length of input power = 24 mm; Kblood (volume thermal conductance to simulate blood flow in tissue) = 0.01 m ~ / r n m ~ / ” ~ .

stainless steel needles, while the rigid array is too perturbing to be useful in a field with a high temperature gradient.

VII. SIGNAL PROCESSING ELECTRONICS A microprocessor-controlled electronics system has been

designed and built t o process temperature information from the linear thermometer array. The system performs five gen- eral functions: It 1) provides a current of 100 pA sequentially to each diode in the array; 2) measures the resulting voitage drop across each lead pair; 3) calculates the absolute tempera- ture of each diode; 4) displays the calculated temperatures; and 5) performs a two-temperature calibration procedure for each array used. The system will drive and monitor arrays of one to twenty diodes, over a temperature range of 0-99.9”C, with 0.025”C resolution and 0.2’C accuracy over a calibration range of 37-45°C. The user controls the calibration procedure and the output display with a 16-key pad.

BARTH AND ANGELL: LINEAR THERMOMETER ARRAYS IN LOCALIZED CANCER HYPERTHERMIA 149

TABLE I EFFECTS OF NEEDLE THERMAL CONDUCTIVITY ox THE TEMPERATURE

FIELD DURING HYPERTHERMIA' IDESCRIPTION I DIODE I I I I I I Si I

1 ARRAY I 27 GA I 24 GA 1 20 GA I 18 GA I Si 1 NEEDLE! I OF ITHERMOMETER 1 (SIX I SS I SS I SS 1 SS 1 NEEDLE I & I

I 18 GA L 1 I I I I I I I 1 ICR O S S - S E C T I O N A L I I I I I I I I

I 0.25 I 0.12 I 0.21 I 0.45 I 0.81 I 0.20 I 0.81 I 1 I I I I I I I

I I I I I I I I

1 (mw-mml I I I I I I I I

I THERMAL ICONDUCTAt$E I 1.02 I 2.16 I 3.64 I 7.69 I 13.9 I 31.5 I 45.4 1

I TEMPERATURE I I I I I I I I I PERTURBATION I I I I I I I I

I 0.21 I 0.45 I 0.77 I 1.61 I 2.78 I 5.32 I 6.78 I (AT PEAK IPOSITION OF I I I I I I I I I UNPERTURBED I I I I I I I I 1 FIELD (QC) I

aPararneters used in the simulation are the same as in Fig. 6 .

Calibration of an array is performed by immersing it in two temperature-controlled baths (typically at 37 and 45°C). The system records the diode voltages at these two temperatures and performs a straight-line fit t o calculate temperature versus voltage. Calibration data for two arrays can be stored in battery- backed memory, and are retained when system power is turned off.

The use of completed arrays in the calibration procedure ensures that the measured voltages and temperature coefficients across each lead pair include parasitic effects such as wire resis- tance and spreading resistances internal to a given diode. The temperature coefficients of internal diode resistances do not degrade measurement accuracy, since these resistances track in temperature with their associated p-n junction in a given diode. Since some short position of the attached wires near an individ- ual diode will also track in temperature with that diode, the TCR effects of the wires are compensated a small extent.

VIII. PERFORMANCE Tests of the linear thermometer array in an ultrasound field

show that it gives accurate temperature measurement with no spurious signal pickup in-vitro. Acute in-vivo testing with ultra- sound in dogs provides measurement as accurate as a nearby implanted thermocouple. The diode array, therefore, appears to be a suitable means of obtaining temperature profiles when ultrasound is used as the heating method.

Tests of the array in-vitro with RF and microwave fields re- veal three problems, all arising from the presence of metallic wires in the array. The wires act as antennas, transmitting high- frequency signals to the electronics system and to the diodes. The pickup of high-frequency signals by the electronics system creates fluctuations at the system output; this problem can be lessened by proper filtering within the system. The high-fre- quency signals transmitted to the diodes create two more serious problems: rectification by the diodes creates a spurious dc voltage component, while power dissipation in the diodes leads to excess heating at the thermometer points (survival of the array in electromagnetic fields of high power density has not been a problem). The rectification and power dissipation problems can be reduced by braiding the wires in the array to reduce the antenna cross section. However, it will still be

necessary to use a pulsed-mode RF or microwave field in ob- taining data from the array. The field will have to be turned off for a fraction of a second to eliminate the spurious voltage due to rectification and allow an accurate temperature reading to be obtained. If, in addition, excess heating remains a prob- lem, it will be necessary to keep the power off for 2-10 s to allow the diodes to come to thermal equilibrium with the sur- rounding tissue before a measurement can be made. Such times are much less than the time constants involved in the return of locally heated tissue to the surrounding body tem- perature, so that the use of the diode array should remain a workable technique in spite of this problem.

IX. DISCUSSION The diode thermometer array must be compared to other

available techniques in assessing its potential usefulness. Ther- mocouples and thermistors can also be used for temperature measurement during hyperthermia, but as noted above they are less suitable for arrays of small cross section. In addition, they must be connected to Conductive leads, and so share the problem of electromagnetic pickup and excessive heating in RF and microwave fields. A competing technique which does not use conductive leads is the fiber optic/gallium arsenide sensor developed by Christensen [ 151, [16]. That sensor does not suffer from electromagnetic pickup problems; however, it requires 2N fiber optic leads for N sensors, so that the cross section of an array of such sensors grows rapidly as the number of sensors increases. Thus the diode thermometer array dis- cussed here may prow to be the optimal device for measuring temperature during localized hyperthermia.

X. CONCLUSIONS Successful clinical use of hyperthermia for cancer treatment

will require the measurement of temperature profiles in the heated tissue. The thermometer array discussed here is capable of measuring such profiles, with good accuracy and minimum mechnical disturbance of the tissue into which it is inserted. This array provides a useful experimental tool for mapping temperature distributions in animals and in physical tissue models. For clinical use, a cheaper array which produces a finer temperature profile is desirable. Present efforts are focused

150 IEEE TRANSACTIONS ON ELECTRON DEVICES, VOL. ED-29, NO. 1 , JANUARY 1982

on the development of a cheap batch-fabricated array of twenty diodes on five flexible leads.

ACKNOWLEDGMENT Grateful thanks must be given to several people for help in

this work. Much help has been given by researchers involved in hyperthermia at the Stanford University Medical Center: P. Kernehan and T. Nelsen of the Department of General Sur- gery, and M. Bagshaw, P. Fessenden, G. Hahn, J. Marmor, D. Pounds, and S. Prionas of the Department of Radiology. In the Stanford University Integrated Circuits Laboratory, much work has been performed by N. Downs, J. Harrison, B. Madsen, F. Shapiro, B. Swaminathan, and S . Wagenseller.

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[3] J. B. Kal and G . M. Hahn, “Kinetic response of rpurine sarcoma cells to radiation and hyperthermia in-vivo and in-vitro,” Cancer Res., vol. 36, pp. 1923-1929, 1976.

[4] J. B. Marmor, N. Hahn, and G. M. Hahn, “Tumor cure and cell survival after localized radiofrequency heating,” Cancer Res., vol.

[5] P. P. Lele, “Hyperthermia,’: presented at the 9th Northeast Bio- engineering Conf., Rutgers Univ., New Brunswick, NJ, Mar. 19- 20, 1981.

[6] F. Sterzer, et al., “RF therapy tor malignancy,” IEEE Spectrum, vol. 17, pp. 32-37, Dec. 1280.

[7] D. F. Hilbiber, “A new semiconductor voltage standard,” in 1964 International Solid State- Circuits Conf , Dig. Tech. Papers, pp. 32, 33, Feb. 19,1964.

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[8] 3. H. Jerman, private communication. [9] S . A. Shamma-Donoghue e t al., “Thin-film multielectrode arrays

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