56 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 45, NO. 1, JANUARY 1998

A Versatile Microprocessor-Based MultichannelStimulator for Skeletal Muscle Cardiac Assist

Erik A. Cheever,Member, IEEE, Dirk R. Thompson, Brian L. Cmolik,William P. Santamore, and David T. George,*Member, IEEE

Abstract— A versatile, microprocessor-based stimulator forskeletal muscle cardiac assist (SMCA) has been designed, con-structed, and used in several studies. The stimulator uses multi-ple bipolar electrodes to deliver arbitrarily specified electricalstimulus sequences to three nerve branches of the latissimusdorsi muscle. The electrodes are electrically isolated to effectregional stimulation of the muscle. The width, amplitude, andinterpulse interval of each pulse in the stimulus sequence areindependently variable, and the three channels are independentlyprogrammable, allowing a wide variety of stimulus patterns.Battery powered units have been used in studies for up to oneyear. In this paper, the stimulator and sample applications inSMCA are described.

Index Terms—Cardiac assist, muscle, stimulator.

I. INTRODUCTION

I N this paper, a microprocessor-based multielectrode stim-ulator for use in skeletal muscle cardiac assist (SMCA)

is described. SMCA is a new treatment for chronic heartfailure in which the contractile power of the latissimus dorsimuscle (LDM) is used to increase the output of a failingheart—providing additional therapeutic options to medicaltherapy, cardiac transplantation, and the total artificial heart.The three major modalities of SMCA under study are aorto-myoplasty (muscle wrapped around the aorta), skeletal muscleventricle (muscle formed into an independent ventricle), andcardiomyoplasty (muscle wrapped around heart epicardium).SMCA has inherent advantages due to the use of autologoustissue: there is no danger of tissue rejection, there is noproblem with donor shortages, and power requirements are lowbecause skeletal muscle requires only minimal external energyto evoke contraction. Cardiomyoplasty is now undergoingFood and Drug Administration (FDA)-approved Phase IIIclinical trials.

Manuscript received December 20, 1996; revised July 7, 1997. This workwas supported by a grant from the Whitaker Foundation. The study ofperistaltic aortomyoplasty was supported by a grant from the Northeast OhioAffiliate of the American Heart Association.Asterisk indicates correspondingauthor.

E. A. Cheever is with the Department of Engineering, Swarthmore College,Swarthmore, PA 19081 USA.

D. R. Thompson and B. L. Cmolik are with the Department of Surgery,Division of Cardiothoracic Surgery, Case Western Reserve University,Cleveland, OH 44106 USA.

W. P. Santamore is with the Department of Surgery, University ofLouisville, Louisville, KY 40292 USA.

*D. T. George is with the Department of Surgery, Division ofCardiothoracic Surgery, Case Western Reserve University, Cleveland, OH44106 USA (e-mail: dtg2@po.cwru.edu).

Publisher Item Identifier S 0018-9294(98)00245-6.

SMCA, however, has difficulties due to the physiologyof skeletal muscles. Normally, the LDM contracts relativelyinfrequently, and tires quickly with continued exertion. Toovercome lack of fatigue resistance, the muscle is conditionedover several months to increase resistance to fatigue—whichunfortunately also decreases muscle strength and speed [1].This problem may be due in part to limitations with thestimulator and electrodes now in use.

The standard stimulus now used for SMCA is a fixed patternof impulses—a pulse train. The pulse train is delivered througha single pair of intramuscular electrodes woven into the LDMin the region of the thoracodorsal nerve [2]. The train consistsof six monophasic, rectangular pulses, each 210s in duration,with an interpulse interval of 33 ms (30 Hz). This patternis repeated with each heartbeat, or on alternate beats [3].Muscles stimulated with pulse trains of nonuniformly spacedelectrical impulses, however, can achieve greater performanceand reduced fatigue compared with muscles stimulated withpatterns of uniformly spaced pulses [4]–[6]. This suggeststhat a muscle stimulator capable of delivering more complexstimulus patterns could elicit greater performance from theskeletal muscle; thereby increasing the amount of cardiacassist.

Another potential improvement involves the use of multipleelectrode pairs instead of a single pair. With multiple electrodepairs, one could take advantage of the neural anatomy of theLDM to selectively stimulate regions of the muscle, achievingbetter control of both the overall “workload” and spatialcontraction pattern of the muscle. Primary innervation of LDMis through the thoracodorsal nerve, which bifurcates twiceto innervate three distinct segments (Fig. 1)—the transverse,oblique, and lateral [7]–[9]. Solaet al. [9] demonstratedvia electromyography, that human LDM segments functionindependently of one another while executing specific mo-tions. Artificially stimulating LDM segments selectively mightfacilitate improvements in the muscle conditioning protocol.Selective stimulation of LDM segments would allow portionsof muscle to be cooperatively rested—lessening the duty cycleof any one segment. Effecting chronic muscle contraction ata lower duty cycle might allow the muscle to retain the speedand strength characteristics of untrained muscle as compared tomuscle trained at rates much greater than the normal duty cycle[10], [11]. Selective stimulation in combination with complexstimulus patterns may compensate for tradeoffs inherent inless-flexible training protocols.

0018–9294/98$10.00 1998 IEEE

CHEEVER et al.: MICROPROCESSOR-BASED MULTICHANNEL STIMULATOR FOR SMCA 57

Fig. 1. Schematic diagram of LDM and neural anatomy. The main branchof the thoracodorsal nerve bifurcates twice into three separate nerve branches.Also shown is the arc over which electromyographic activity was recorded toconfirm selective stimulation in the dog.

In this paper, we describe a stimulator that allows arbitrarypulse patterns to be sent to three isolated, bipolar electrodes,which allows selective stimulation of LDM. The amplitude,width, and interpulse intervals of each pulse are independentlyprogrammable. The device was designed for low power, soit can be powered by batteries and worn in a backpack forlong-term studies.

II. DESIGN

The design of the programmable multichannel stimulator(PMCS) for SMCA incorporates two important features. First,the voltage pattern delivered to the bipolar stimulus electrodescan be programmed with complete flexibility—that is, anyarbitrary voltage waveform as a function of time can begenerated (within certain broad time and voltage constraintsoutlined here). When restricting stimulator output to the com-mon practice of activating skeletal muscle with monophasicor biphasic stimulus pulses, the train length, pulse amplitudes,pulse widths, and interpulse intervals are easily programmed.Second, the PMCS has three separate, electrically isolatedoutput channels that can be used to stimulate different regionsof the LDM without cross-talk. This second feature wasimplemented in consideration of the anatomy of the LDMdescribed. Circuitry has been developed that can independentlystimulate each of the three nerve branches with independentstimulus patterns.

The system is built around a microcontroller (MotorolaMC68HC811) which sends the necessary control signals tothe analog electronics to generate the desired pulse sequence.A Microsoft Windows-based personal computer (PC) is usedto rapidly develop new sequences, which are subsequentlydownloaded to the microcontroller via a serial (RS-232) in-terface. Three versions of the stimulator have been produced:a portable, battery-powered unit with a single output channelconstructed using a printed circuit board for use in muscletraining studies not involving selective stimulation; a secondportable, battery-powered unit with three output channels, alsoon a printed circuit board, for use in studies concerning selec-tive stimulation or using the rapid-pacing heart failure modelwith simultaneous chronic muscle conditioning; and finally, abench-top three-channel unit for use in acute studies where a

high degree of flexibility is needed, but power consumption orsize is not of concern. The single-lead unit is accommodated ina 6 12 2.5 cm case, and, because it has only one channel,does not contain circuitry for electrical isolation. The case forthe three-lead unit is roughly 912 2.5 cm. The portableunits are small enough to be worn in a harness on the back ofa dog and have a separate board which contains the circuitryfor the RS-232 interface that is connected temporarily via aribbon cable when communicating with the PC. The bench-top unit utilizes a Motorola 6811 EVB and two wire-wrappedcards. Otherwise, the three units are substantially similar inhardware and software. Descriptions of the microcontroller,associated analog electronics, and the PC software follow. Thedifferences among the units are highlighted.

A. Microcontroller

The microcontroller (c) communicates with the PC andkeeps track of the timing of the three independent pulsepatterns. The c chosen has a built-in serial port, as well asparallel ports, 256 bytes of random-access memory (RAM),2 kbytes of electrically eraseable programmable read-onlymemory (EEPROM), an eight-channel analog-to-digital (A/D)converter, several timers, and pulse accumulators, and becauseit is a complimentary metal oxide semiconductor (CMOS)processor, draws very little current. The internal timers areused to keep track of the timing of the pulses. Several of thedigital input–output lines are connected to digital-to-analog(D/A) converters within the analog electronics to control thethree separate stimulator channels. The serial port can beconnected to the PC with appropriate voltage level shiftingcircuitry (MAX233, Maxim). When the c is connected to thePC, it can process such commands as downloading a pulsesequence from the PC, generating the pulse sequence once,or repetitively generating the sequence at a rate determinedby the user. If the latter command is given, thec can bedisconnected from the PC and level shifting circuitry to runindependently.

The operation of the c as it produces pulses proceedsas follows. The voltage waveform to be generated by thestimulator is first entered into a program written in MicrosoftVisual C for Windows. The waveform is entered as a list oftimes together with the corresponding decimal representation(0–7) of the binary number to be written to each D/A converter(Table I). The binary number shall be referred to as a voltageamplitude. (In the examples that follow, it is assumed thatPMCS output is adjusted such that decimal numbers 0–7correspond to 0–7 V.) A single time together with voltageamplitudes is termed an “event” and corresponds to a changein the output voltage of one or more channels. The samplewaveform given in Table I is a four-event waveform in whichthe first event, occurring 100s from time zero, sets channels1, 2, and 3 to 1, 3, and 4 V, respectively. Event number2, occurring at 210 s, sets all channels to 0 V. At 10 000

s in the waveform, the third event, channel 1 is set to 1V and, finally, the fourth event at 10 500s is to set allchannels to 0 V. Although absolute times are entered intothe Windows program, before downloading the waveform into

58 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 45, NO. 1, JANUARY 1998

TABLE ISAMPLE WAVEFORM FILE

microprocessor memory, the absolute times are first convertedinto a series of time intervals (relative times). The fourcorresponding time intervals from Table I would be 100, 110,9790, and 500 s.

Each event in the waveform requires three bytes of memory.The most significant 12 bits of the three bytes is the timeinterval. The exact range and accuracy of the time interval isdetermined by the c counting rate. For the portable units,the time interval has a maximum of 131 ms (33 ms for thebench-top unit). The maximum time interval is determined bythe time resolution of the stimulator, which is 32s (8 s forbench-top). Intervals larger than 131 ms (33 ms for bench-top)are achieved in software by dividing into smaller intervals. A50-ms time interval on the bench-top model would, thus, becreated from two 25-ms intervals.

The least-significant 12 bits control the three channels, usingfour bits per channel on two of the parallel ports of themicrocontroller. Of these four bits, three are connected toa D/A converter to set the amplitude of the voltage output( , and in Fig. 2), and the fourth bit (the Enableline) controls a CMOS switch (described later) that determineswhether or not the channel is isolated from the muscle. Whilea pulse sequence is being generated, thec uses an on-boardtimer to wait for a time specified by the 12-bit time interval,then outputs the channel amplitudes to the analog section of thecircuit. It repeats this procedure until it arrives at the end of thesequence. If the pulse sequence is to be repeated, the internaltimers are used to manage the pulse train repetition rate anddetermine when the pulse sequence is replayed. Whenever the

c is idle because it is waiting for the timer, it goes into alow-power “wait” mode.

For the portable units, pulse sequence data is stored inEEPROM so that the stimulator can resume generating pulsesafter a battery change, or following a momentary power loss,without needing to be reprogrammed. The current programoccupies roughly 0.5 kbyte of EEPROM leaving 1.5 kbyte fordata or additional program code. For the bench-top version,the pulse sequence is more conveniently stored in RAM,because it is much quicker to write data into RAM andthere is more RAM than EEPROM (8 kbytes of RAM versus2 kbyte of available EEPROM). In practice, the portablestimulators would be programmed once per week or less often,so download time is not important and is between 5 s and 120s (depending upon the waveform) compared to less than 1 sfor the bench-top unit. The benefit of having the stimulatorcontinue to run following a momentary power disturbance,

Fig. 2. Schematic diagram of analog section of electronics. All op amps areLM324, and analog switches are AD7510. See text for details.

which may occur if the batteries to the externalized stimulatorare bumped or need changing, far outweighs the inconveniencecaused by the longer download time.

The bench-top version has an added feature to facilitatemonitoring the stimulus output. Hardware was added to pro-duce a “stretched” copy of each of the stimulus pulses 5–10ms (adjustable) in width. In this way, a data acquisitionsystem can utilize a sample period of 1–2 ms (appropriatefor physiological signals such as force or pressure), and beensured of capturing the pulse sequence. Pulses only 210sin duration would require an unnecessarily high sampling rate.These stretched pulses also make the stimulus pattern easier toview on a digital oscilloscope (a 210-s pulse repeated every33 ms can be quite difficult to view).

B. Analog Electronics

Fig. 2 shows the analog electronics connected to the positiveelectrode of one of the bipolar output channels. When theanalog switches (at the right) are open, the electrode iselectrically isolated. The 100-kresistor between the electrodeand ground was added solely to protect the analog switchesby acting as a shunt for any static electricity. The resistorvalue was chosen to be large enough so that it would have noaffect on the operation of the circuit or the functioning of themuscle, and none was observed. Whenever the switches areopen, outputs , , and are all zero, so that the outputof the op-amp is also zero and the 100-kfeedback resistorat the op-amp output prevents saturation.

When the switches are closed, , , and act as a 3-bitbinary word going into an R-2R resistor D/A converter. In thisconfiguration, the 100-k resistor at the op-amp output haslittle effect, because the much lower resistance of the analogswitches shunts it. Thus, the op-amp acts as a noninvertingamplifier whose gain is controlled by the variable resistor,and amplifies the output of the D/A converter. Two switchesare used in this configuration because the impedance of theanalog switches is roughly 10–20% of the impedance ofthe electrodes and tissue, causing a significant voltage dropacross the uppermost switch due to the current going to theelectrode. There will be little drop across the lower switch,

CHEEVER et al.: MICROPROCESSOR-BASED MULTICHANNEL STIMULATOR FOR SMCA 59

however, since only very small currents enter the feedbacknetwork. This circuit ensures that the voltage at the electrodewill be held constant even if there are substantial changes inelectrode current. A similar scheme is employed at the negativeelectrode to implement an active ground. When the switchesare open, the negative electrode is electrically isolated; whenthe switches are closed, the electrode is maintained at groundpotential regardless of the current being supplied.

The circuit is repeated for the other two channels. Whena pulse sequence is being generated at a given electrode, thechannels for that electrode are activated (that is, the switchesare closed). When individual electrodes or pairs of electrodesare activated, opening the analog switches isolate those notactivated. For example, if channel 1 is to be activated whilechannels 2 and 3 remain inactivated, the switches for channel1 are closed and the switches for channels 2 and 3 are opened.

The analog electronics are powered by a dc-to-dc voltageconverter (MAX633, Maxim) generating 15 V from a 5-V(battery) source. This simplifies the analog circuitry and allowsa wide range in the outputs of the op-amps used to stimulatethe muscle. Without the dc-to-dc converter, the electronicswould have to operate between5 V and ground, limiting themaximum obtainable output voltage and making it impossibleto maintain the active ground (which requires a negative outputvoltage from the associated op-amp). Currently, the analogcircuitry accounts for 75–80% of the power consumption ofthe portable stimulators.

The capabilities of the microcontroller and associated cir-cuitry are generally similar for the portable and bench-top units(Table II). When capabilities differ, however, the value forthe portable units is presented, followed by the value for thebenchtop unit in braces “{benchtop value}.” The maximumvoltage output is variable up to 12 V. The output voltagecan be set to any of eight values between zero and themaximum voltage, inclusive. Adjusting the potentiometer inthe feedback loop (Fig. 2) can effect finer adjustment, or achange in the voltage range. The maximum current of 35-mA per channel is the maximum specified current for theLM324 op-amp. When the stimulator is programmed to deliverrectangular monophasic or capacitor-blocked biphasic pulses,the pulse width, interpulse interval, and R-wave delays areall individually programmable with a resolution of 32 {8s}.The minimum time allowed is 32 {30 s}. The number ofpulses and length of the pulse train is determined by availablememory. As stated previously, the pulse sequence data isstored as a series of 3-byte events, in which each eventcorresponds to a change in one or more of the stimulatoroutputs. The only exception is for long time intervals, wherethe outputs do not change—these longer intervals must bebroken down into a series of events each no longer than 131{33} ms. A rectangular pulse requires two events, one for theupstroke and one for the downstroke. The stimulators haveapproximately 1.5 {8} kbyte of free memory and can store upto approximately 250 {1300} pulses (fewer for long interpulseintervals, which must be broken down into multiple events).In practice, these broad timing parameters have not posed anylimitations on the use of the system. The battery-poweredstimulators run for five days on a 2450-mAh 5-V battery

TABLE IISUMMARY OF THE HARDWARE CAPABILITIES OF THE

PROGRAMMABLE MULTICHANNEL STIMULATOR*

*When different, specifications for the portable unit are listed with those for the bench-topunit in braces {}.

(four AA alkaline batteries). Battery life does not depend uponstimulus intensity because the duty cycle for stimulus outputis less than 1%.

C. Personal Computer

To communicate with the microcontroller, the PC runs aprogram with an easy-to-use graphical interface. The programhas a series of windows that allow the user to specify var-ious pulse sequences (described in Section III), download asequence to the microprocessor, and specify when and howfrequently the PMCS applies the pulse sequence to the muscle.A typical screen for pulse specification might include the pulseamplitudes, pulse widths, interpulse intervals, as well as thenumber of pulses (Fig. 3). The sequence can then be displayedin a separate window. As new pulse sequences are developed,they can be saved individually, or as a list of sequences, forsubsequent retrieval. The program was written with MicrosoftVisual C for Windows computers. The format of the filewritten to c memory does not depend upon the sequence—the

c always processes a list of events. The primary task ofthe Windows-based software, therefore, is to provide a user-friendly interface through which the event file is created,saved, and written to microcontroller memory.

III. OPERATION

A. Stimulus Waveforms

The stimulator software running on the PC allows the userto specify six different stimulus patterns and program thePMCS to deliver patterns in six different operating modes(Table III). The general stimulus pattern is any voltage patternas a function of time, subject to the restrictions discussedpreviously. To create a general waveform, one enters a listof times (in s) and bit levels (0–7) for the three channels[Fig. 3(a)]. The general waveform is the most tedious to createbut there is a high level of flexibility, because the resultingwaveform can be quite complex. The software allows theprogrammer to enter one line at a time, and each line canbe easily edited through an “Edit” subwindow. Entering a bitlevel of 1 denotes the channel voltage is to remain unchangedfrom the previous line.

60 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 45, NO. 1, JANUARY 1998

(a)

(b)

(c)

Fig. 3. Screen captures of pattern creation windows and resulting output waveforms. Three different patterns are shown: (a) general, (b) variable-frequencytrain, and (c) peristaltic.

More commonly, muscles are stimulated with a sequence ofmonophasic square-pulse or capacitor-blocked biphasic pulses,hence, the expression “pulse train.” The remaining examplesof easily created stimulus patterns are pulse trains (Table III).Theconstant-frequency trainis a series of pulses with constant

interpulse interval. In the constant-frequency train window,one enters either the interpulse interval (s) or the intratrainfrequency (Hz), the number of pulses in the train or trainduration ( s), and the pulse width (s) Finally, the voltagelevel (bit level) for each channel is entered. The software

CHEEVER et al.: MICROPROCESSOR-BASED MULTICHANNEL STIMULATOR FOR SMCA 61

TABLE IIISUMMARY OF THE (a) STIMULUS PATTERNS AND (b) OPERATING MODES

(a)

(b)

automatically calculates interpulse interval when the intratrainfrequency is entered (or vice versa) and calculates the numberof pulses in the train when the train duration is entered (or viceversa). The voltage level for each pulse is constant within thetrain but may differ among channels.

Thevariable-frequencytrain is a train of pulses with nonuni-form interpulse intervals. In the variable-frequency train win-dow, a list of pulse widths (s) and interpulse intervals (s)are entered one line (pulse) at a time [Fig. 3(b)]. The samestimulus pattern is sent to the three channels, however, thereis an option to set a variable delay among the three channels sothat the three channels can optionally be “phase-shifted” fromone another. The bit level for each channel is also entered. Forthe variable-frequency train, the voltage level for each pulsewithin a train is constant.

A peristaltic pattern is one in which an identical patternis delivered to each lead, however, the outputs are shifted intime from one another. A peristaltic pattern is created in thesame manner as the constant-frequency or variable-frequencytrain, with the addition of separately entering of time delays( s) between the first and second, and then second and third

channels [Fig. 3(c)]. The voltage output on each channel maybe different for peristaltic patterns.

In rotary stimulation, the individual impulses on the threechannels are phase shifted slightly from one another. The threechannels generate stimulus trains simultaneously, however, theindividual pulses among the three trains are interleaved. Forconstant-frequency rotary stimulation, one enters the desiredphase shift among trains as a percentage of the constantinterpulse interval. For a variable-frequency train, the intervalis not constant, so a true rotary pattern is not possible.Nevertheless, entering a nonuniform time delay among thethree channels may form a rotary-like pattern (Fig. 3).

An additional feature in the software is morphing. Morphingis a feature whereby the first and last train in a series ofdesired trains are entered; the morphing algorithm recognizesthe variables that differ between the two trains, and creates aseries of trains with intermediate characteristics. For example,if one morphs two constant-frequency trains with identicaltrain duration, but differing interpulse interval, the morphingalgorithm will create a series of trains with interpulse intervalsbetween the two original trains. Morphing is especially useful

62 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 45, NO. 1, JANUARY 1998

in those situations where a series of stimulus patterns are to betested—for example, in studies of skeletal muscle in which theforce-frequency relation is measured. Also, in studying the ef-fects of a muscle wrapped around the heart (cardiomyoplasty)or aorta (aortomyoplasty) one typically varies the R-wavedelay and train duration to find the train that maximizescardiovascular effect. This can be done quickly and easilywith the morphing routine.

A final stimulus pattern available with this stimulator wasimplemented to meet our need of inducing heart failure viarapid ventricular pacing, while simultaneously chronicallystimulating the LDM to induce fatigue resistance. A multichan-nel stimulator can accomplish both tasks. In rapid-pacing heartfailure model, the heart is paced through an intraventricularpacing electrode to beat at 200–250 beats per minute (bpm).Over the course of several weeks, the heart becomes dilated,left-ventricular ejection fraction falls, and clinical signs ofheart failure become evident [12]–[16]. The level of heartfailure can be manipulated by adjusting the rapid pacing rate.To study the effects of SMCA, one must first convert thenormally fatigueable latissimus dorsi muscle into a fatigue-resistant phenotype. One technique of converting skeletalmuscle into fatigue resistance is to stimulate chronically witha chronic, low-frequency (2-Hz) stimulus [17]–[19]. To studySMCA in the rapid-pacing heart failure model, requires thestimulator to have the ability to stimulate both heart andmuscle at independent rates.

The two-channel asynchronousstimulus pattern was im-plemented to achieve these goals. In this mode, constant-frequency stimuli are generated on two of the three channels.One enters the output amplitude, pulse width, and stimulusrate for each channel. If the ratio of the two frequencies is arational number, for example 2 Hz on the channel connectedto the muscle electrodes and 4 Hz (240 bpm) on the channelconnected to the heart electrodes, there is no difficulty. Thestimulator repeats at a rate of 2 Hz and generates two pulses onthe muscle channel for every four pulses on the heart channel.If the ratio of frequencies is irrational, say 2 Hz on the musclechannel and 3.83 Hz (230 bpm) on the heart channel, thissimple scheme fails. To correct this problem, a single, verylong waveform is generated with the heart-pacing channelrunning at 3.83 Hz, and the muscle channel at 2 Hz. Thesequence is made as long as possible under the constraintsof memory available. The sequence length and repetition rateare chosen such that heart-pacing channel runs constantly at3.83 Hz. The muscle-pacing channel will generally run at 2Hz, but may miss a pulse at the very end of each long pulsesequence. Because this method requires a long waveform witha large number of events, this waveform takes the longest todownload from the PC to the portable stimulator—up to 120s. All other waveforms download in just a few seconds.

B. Stimulus Output Modes

Once the stimulus waveform has been created and down-loaded to PMCS memory, there are six “modes” in which thestimulus pattern can be outputted (Table III). Theconstant-train-rate mode outputs the stimulus waveform to the muscle

at a user-specified rate and for a user-specified number of repe-titions or indefinitely. (If indefinitely, a separate command tellsthe PMCS to quit.) Inexternal triggermode, the stimulatoroutputs a pulse pattern in response to an external trigger. Thistrigger is generally applied by the data acquisition hardwareto synchronize muscle contraction to data acquisition. In thevariable trigger-to-stimulusmode, the stimulator will outputthe waveform on integer increments of a transistor-to-transistorlevel (TTL) trigger input. When using this mode, an externalR-wave detector is connected to the TTL trigger input so thatthe muscle can be made to contract on integer increments ofthe normal cardiac rhythm. The minimum ratio is one (themuscle is stimulated on every heartbeat), up to 255 (muscleis stimulated on every 255th heartbeat). In practical usage, amaximum R-wave-to-muscle-contraction ratio is in the rangeof 10–20 which, over time, is then decreased to a ratio ofone or two following conversion of the muscle into a fatigue-resistant phenotype. Insleep/wakemode, the user can specifythe duration of time that the stimulator output is enabledand disabled. The user enters the number of hours that thestimulator is on and off and the rate at which the waveformis output. This mode is useful, for example, in studies inwhich the muscle is stimulated for less than 24 h/day. Thestimulator developed here could be used to accomplish sucha study without requiring any attention from the user—thestimulator need not be manually turned on or off each day, butrather does so automatically. Finally, insequential-download-and-runmode, upon receipt of a TTL trigger, the PMCS willoutput the waveform currently in its memory, download thenext sequence from a list of waveforms, and waits for the nextTTL trigger to repeat the process. (When this mode is initiallyactivated, the PMCS downloads and delivers the first sequence,downloads the next sequence in the currently open file.) Thisis useful in acute studies because, when combined with themorphing feature, it allows one to create a series of stimuluswaveforms to be tested and then run the test automatically.

IV. RESULTS

A. In Situ Measurements

In six acute studies, the PMCS was used to selectivelystimulate the canine LDM [20]. The three branches of thethoracodorsal nerve were isolated and instrumented with bipo-lar nerve cuff electrodes. The PMCS successfully activatedindividual nerve branches of the canine LDM with no cross-talk to the other branches as confirmed by measurementof the regional EMG (electromyography). Electromyographicactivity was measured at 10increments along an “arc”originating from the lateral border of the LDM (Fig. 1). Theroot-mean-square (rms) of the electromyographic response toa single stimulus pulse at each site was normalized to themaximum rms value calculated at each respective site. Wheneach electrode was activated, the corresponding muscle regionshowed EMG activity, while the remaining muscle segmentsshowed little or no activity (Fig. 4). Using a mechanical testingsetup designed to assess canine latissimus dorsi contractilefunction noninvasively [21], gradation of muscle force during

CHEEVER et al.: MICROPROCESSOR-BASED MULTICHANNEL STIMULATOR FOR SMCA 63

Fig. 4. Regional activity measured as a percentage of maximum rms whilethe three branches of the thoracodorsal nerve were individually activated.

by selective stimulation was recorded (Fig. 5). In this sam-ple study, stimulation of the transverse, oblique, and lateralsegments produced 34, 30, and 44% of muscle peak force,respectively. When stimulated in tandem, the combinationsof transverse-oblique, transverse-lateral, and oblique-lateralproduced 65, 76, and 76% of peak muscle force, respectively.

B. Application to the Cardiovascular System:Peristaltic Aortomyoplasty

An example of applying selective stimulation of the latis-simus dorsi muscle during SMCA is in aortomyoplasty, atechnique in which the muscle is wrapped around the tho-racic aorta, is shown. For this study, the LDM was dividedinto two strips—one from the lateral segment and secondfrom the transverse-oblique segment. The muscle was in-strumented with two pair of stimulus electrodes to activatethe two separate segments. Stimulating the transverse-obliquesegments as one combined segment was done because theoblique segment was small. The lateral portion of the musclewas wrapped around the distal descending aorta while thetransverse-oblique segment was wrapped around the proximaldescending thoracic aorta. Pressure transducers were placedinto the left ventricle and the ascending aorta (between theheart and wrapped segment). The aorta was instrumentedwith two Doppler-ultrasonic flow probes; one above and onebelow the wrapped region to calculate total blood volumeejected from the wrapped segment during muscle activation,as well as the relative blood volume ejected proximally anddistally. Cardiac R-wave detection from the electrocardiogram(ECG) signal was provided by a physiologic recorder, whichthen triggered the PMCS. The two muscle segments werestimulated synchronously, as would be done with nonselectivestimulation, and then with various time-delays between thetwo segments in an attempt to effect peristalsis—preferentiallydirecting blood proximally toward the heart, or distally downthe descending aorta. Fig. 6 shows the resulting pressure, flow,and ECG signals during aortomyoplasty. The stimulus pulsesare seen as noise in the ECG. Stimulation was initiated ap-proximately at the dicrotic notch and continued until just priorto the R-wave of the ensuing cardiac cycle. Aortic pressureduring the stimulated beats was increased compared with thenonstimulated beats. The proximal and distal flow signalswere similar during nonstimulated beats and diverged duringstimulated beats indicating blood volume was ejected from the

Fig. 5. Force as measured at the dog’s forepaw while the three branchesof the thoracodorsal nerve were individually and collectively activated. Datawere first normalized to the maximum force measured during activation ofthe main nerve cuff.

wrapped segment. The difference between the two flow signalsshows total flow rate out of, or into, the wrapped region.The difference signal was near zero during unstimulated beats,was positive during muscle stimulation as blood was ejectedfrom the wrapped region, and was negative during the musclerelaxation period and blood refilled the wrapped region. Fromthe integrals of the individual flow tracings and the integralof the difference signal, the relative direction and total ejectedvolume were calculated. During synchronous stimulation, 35%of the blood was ejected proximally (Fig. 7). When the distalmuscle segment was made to contract first, followed bydelayed contraction of the proximal segment (cephalad delay),there was greater blood flow proximally. Conversely, whenstimulation of the distal segment was delayed, there was apreferential blood flow distally. Ejecting blood proximallyresulted in greater increases in mean diastolic aortic pressureand endocardial-viability ratio (ratio of the left-ventricularsystolic pressure-time integral to the aortic diastolic pressure-time integral: a clinical index of the supply demand ratio forthe heart).

V. DISCUSSION

Researchers have been working on selective muscle stim-ulation, primarily for functional neuromuscular stimulation[22]–[26], but also for SMCA [27]–[29]. De Lucaet al.[27] used a multielectrode nerve cuff placed around thethoracodorsal nerve of goats to stimulate selectively the LDM.Their technique used a nerve cuff with up to 12 electrodes fixedto the cuff wrapped around the nerve. By appropriate tuning ofthe stimulus current among electrode combinations, selectivestimulation of individual fascicles within a multifasciculatednerve fiber was often, but not always, achievable. They do notdescribe the stimulator used in their studies.

Thomaet al. extended their efforts in multichannel systemsfor the stimulation of the lower extremities in paraplegia andin the diaphragm [30] to SMCA [29], [31]. They attempt toactivate individual fascicles within a multifasciculated nerveby suturing multiple monopolar electrodes directly to theepineurium, much like DeLucaet al. [27], but instead of prepo-sitioning the electrodes on a cuff, the electrodes are placedaround the nerve, individually—a delicate surgical procedure.Passing current across electrode pairs then achieves selectivestimulation. They describe stimulating “isolated portions ofthe left or right distal muscle or the right proximal part of themuscle,” but present no data [29]. In contrast, the stimulator

64 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 45, NO. 1, JANUARY 1998

(a)

(b)

(c)

Fig. 6. Sample raw data from peristaltic aortomyoplasty. (a) Left-ventricular pressure (LVP), aortic pressure (AoP), and ECG. (b) Aortic blood flowdistal (Dflow) and proximal (Pflow) to the wrapped aortic segment. (c) The difference between the proximal and distal flows (Wflow) giving the netflow rate from the wrapped segment.

described here effects selective stimulation through multiplebipolar electrodes positioned about each nerve branch, and hasalso been used with intramuscular electrodes woven into themuscle in the region of the nerve branches. Finally, selectivitywas documented by regional electromyography (EMG) andgradation of muscle force.

Lanmuller and colleagues recently reported on an im-plantable, ECG-triggered, multichannel nerve stimulator foruse in SMCA [31]. Their device has eight monopolar outputchannels, presumably attached to epineurial electrodes asdescribed in their earlier work [29]. The device describedhere, though not implantable, is considerably more flexible.The stimulator of Lanmuller and colleagues is limited to a peakstimulus frequency of 66 Hz for single-channel output (one setof four epineurial electrodes) and to 33 Hz for dual-channeloutput (two sets of four epineurial electrodes). Conversely,the stimulus output of the device described here does notdepend upon the number of stimulus channels. The stimulatorof Lanmuller and colleagues has an impulse duration limitedto 200–1000 s and maximum current output limited to 4mA. This may constrain the stimulator to use with nerveelectrodes (epineurial or cuff). They do not describe the

ability to output arbitrary waveforms, nor variable-frequencypulse trains. Lanmuller and colleagues control the ratio ofmuscle contractions to detected cardiac beats by adjustingthe duration of a “disabletime” during which R-wave is notsensed or counted by the microcontroller. This techniqueresults in heart-rate dependent changes in the muscle-to-cardiac contraction ratio. The stimulator described here cancount a trigger signal and, more conveniently, stimulates themuscle on integer increments of the trigger count rangingfrom 1–255 (Table III).

Laboratories working on SMCA and the effects of chronicelectrical stimulation have also developed single-channel out-put stimulators for research purposes [32]–[35]. A commer-cially available stimulator is presently in use in FDA-approvedclinical trials of cardiomyoplasty [36]. These stimulators donot appear to be able to output an arbitrary, user-specifiedstimulus waveform or are constrained to constant-frequencystimulus patterns [33], however, preliminary reports indicaterecent efforts on a device capable of more complex stimuluspatterns [34], [35].

Pattisonet al. [32] constructed an external, battery-operatedstimulator from discrete components. Their device sensed the

CHEEVER et al.: MICROPROCESSOR-BASED MULTICHANNEL STIMULATOR FOR SMCA 65

Fig. 7. Sample derived data from peristaltic aortomyoplasty. Cephalic delayindicates the distal muscle segment contracted first followed by the proximalsegment to direct aortic blood flow preferentially cephalad. Synchronousindicates the lateral and transverse-oblique segments contracted simultane-ously. Caudal delay indicates the proximal segment contracted first to directaortic blood flow preferentially caudad. Peristaltic muscle motion was able todirect blood flow preferentially toward the proximal aorta (cephalad direction)or toward the distal aorta (caudad direction). Increased blood flow to theascending aorta increased mean aortic pressure and endocardial-viability ratio.

R-wave and could output constant-frequency pulse patterns ofduration 60–400 ms, intraburst frequency of 14–140 Hz, andpulse-width of 100–900 s. A constant-voltage output wasadjustable over 1–7 V. They do not indicate the battery-lifeof their stimulator. The muscle-to-cardiac ratios are limited bya binary counter to 1:1, 1:2, 1:4, or 1:8. The stimulator ofPattison and colleagues is less flexible than the unit presentedhere, primarily because it is not microprocessor based.

Callewaertet al. [33] developed an implantable stimulusunit for the purposes of studying the adaptive response ofmuscle to chronic stimulation. Analogous to the stimulatorpresented in this work, their unit does not sense the R-wave. The stimulator of Callewaert and colleagues was anapplication-specific integrated circuit of hybrid technologydesigned for small size and low power consumption. At con-struction, the device was preprogrammed with eight differentconstant-frequency stimulus waveforms, which could not bemodified following construction. The single-channel unit, witha constant-voltage output, was programmable to output eithersingle stimulus pulses or bursts of pulses at intervals of 0.1,0.2, 0.4, or 0.8 s. Through a transcutaneous optical link burstparameters could be set at either 40 or 80 Hz, and durations ofeither 100 or 200 ms. Peak stimulus current is 2 mA and for a1-mA stimulus current, the unit is anticipated to function forone year on a single battery. The low stimulus current limitsuse to electrodes positioned in close proximity to the nerve(cuff or epineurial). Advantages are a long lifetime in a small,fully implantable unit.

More recent work by this group indicates progress onmanufacturing a more advanced version of their earlier design[34], [35]. Though not yet extensively described, the newerimplantable microcontroller-based neuromuscular stimulatorwas developed with the ability to output greater number ofpossible stimulus patterns than was accomplished in theirearlier work. The new stimulator was again developed witha preprogrammed array of 12 [35] or 16 [34] stimulus output

patterns which could be chosen following implantation viaoptical link. An improvement in using a microcontroller, wasthat the stimulus pattern could be changed after constructionof the stimulator. The maximum stimulus current is notindicated, but presumably is similar to the 2 mA of their earlierstimulator. The range of possible stimulus patterns availableis not mentioned, however, the use of a microcontroller wouldsuggest the potential for a high degree of flexibility.

The cardiomyoplasty stimulator presently in use in FDA-approved Phase III clinical trials is manufactured by Medtron-ics, Inc. (Minneapolis, MN) [36]. An implantable unit capableof R-wave detection and single muscle channel output, thedevice outputs constant-frequency trains. It is also a demandVVI pacemaker capable of pacing the heart should rates dropbelow a programmed value. Muscle-to-cardiac contractionratio can be chosen from one of eight ratios in the rangeof 1:1–1:8. Overstimulation of the muscle is prevented byinhibiting muscle output when the R-R interval decreasesbelow a programmable value. The device has programmablepulse amplitude (0.1–9 V), pulse width (15–504s), interpulseinterval (8–242 ms), and number of pulses in the constant-frequency burst (1–14). At standard stimulus values, estimateddevice lifetime is six years. As a commercially available de-vice, the Medtronics unit is certainly highly flexible, adaptive,and has the advantage of a demand VVI pacemaker. Its musclestimulus output appears to be limited to constant-frequencystimulus patterns.

An interesting application for the stimulator described hereis as a simultaneous cardiac pacer and muscle stimulator. Tounderstand the therapeutic effects of heart failure treatments,it is useful to cause heart failure in experimental animals. Therapid-pacing heart failure model is well documented [12]–[16].To achieve rapid ventricular pacing with commercially avail-able pacemakers requires overriding the pacemaker safetyby gluing a magnet to the stimulator casing [12]. With thestimulator described here, one output channel could be usedfor rapid ventricular pacing via pacing leads implanted intothe heart, while another channel could chronically stimulatethe skeletal muscle. In this way, a single PMCS can conditionthe muscle while simultaneously inducing heart failure. Thetwo-channel asynchronous operating mode (Table III) wasimplemented specifically with this task in mind—the heartcan be paced at a rate independent of the skeletal musclestimulation rate. In this way, the heart-pacing rate could beadjusted as necessary to control the severity of heart failurewithout influencing the muscle-conditioning process.

To date, the PMCS has been used in acute and chronicstudies of in-situ LDM stimulation [20] and in acute andchronic studies of aortomyoplasty [37]. The stimulator has alsoseen use in our laboratory for simultaneous induction of heartfailure by rapid pacing and muscle conditioning. These studieshave not required stimulator-based R-wave detection. In takinghemodynamic measurements during studies of aortomyoplasty,an ECG trigger was provided externally by a separate phys-iological amplifier. However, for experimental protocols inwhich muscle contraction is timed with cardiac contraction,on-board, epicardial R-wave detection can be implemented.There are a variety of methods in the literature for detecting

66 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 45, NO. 1, JANUARY 1998

the R-wave [38, 39]. Any of these methods easily would beapplied to the PMCS and there is plenty of unused memoryavailable 1.5 kbyte) in the present implementation. If ahardware detector is used, the microcontroller can spend thevast majority of time in a very low-power “sleep” mode withthe R-wave detector connected to one of the interrupt lines.Alternatively, it would also be possible to use an adaptiveR-wave detector in software [40] using the on-board A/Dconverter.

ACKNOWLEDGMENT

The authors would like to thank J. J. Michele for hisexcellent surgical support and J. R. Buckett for his criticalreview of the manuscript. They would also like to thank theprogrammers who worked on this project, J. R. Birge, A. M.Marsh, and J. Tracy

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Erik A. Cheever (S’81–M’87) received the B.S.degree in engineering from Swarthmore College,Swarthmore, PA in 1982, and the M.S.E. and Ph.D.degrees in bioengineering from the University ofPennsylvania, Philadelphia, PA in 1984 and 1989,respectively.

He is currently an Associate Professor in the De-partment of Engineering at Swarthmore College. Hiscurrent research interests include design of devicesfor SMCA, and pulse oximetry and bilirubinometry.

Dirk R. Thompson was born in Rochester, MN, in1968. He earned the B.S. degree from the Universityof Minnesota, with a major in biology and minorsin chemistry, mathematics, and physics. He attendedgraduate school at Drexel University, Philadelphia,PA, in the Program in Biomedical Engineering, anddid his Master’s thesis research at the PhiladelphiaHeart Institute, Philadelphia, PA. His Master’s thesisinvolved selective neural stimulation of latissimusdorsi muscle; and since then, he has applied math-ematical optimization routines to the parameters of

skeletal muscle stimulation.He is presently a Researcher in the Cardiothoracic Surgical Research

Laboratory, Case Western Reserve University, Cleveland, OH. His researchinterests center around SMCA. He has developed algorithms to acquire andanalyze hemodynamic data for studies of aortomyoplasty.

Brian L. Cmolik received the B.S. degree inchemistry from Xavier University, Cincinnati,OH, in 1979. He received the M.D. degree fromthe Northeastern Ohio Universities College ofMedicine, Rootstown, in 1985.

From 1985–1993, he did his residency in generaland thoracic surgery at the Medical Collegeof Pennsylvania, Allegheny General Hospital,Pittsburgh, PA. He is presently an AssistantProfessor of Surgery at Case Western ReserveUnivesity, Cleveland, OH, and Staff Surgeon at the

University Hospitals of Cleveland, Cleveland, OH. His research interestsinclude aortomyoplasty and cardiomyoplasty.

William P. Santamore received the B.S. degreein electrical engineering from Manhattan College,Bronx, NY, in 1969. He received the M.S. degreein biomedical engineering from Drexel University,Philadelphia, PA, in 1972 and the Ph.D. degree inphysiology from Temple University, Philadelphia,PA, in 1975.

He has held faculty appointments at Temple Uni-versity, University of Pennsylvania, Philadelphia,and Wake Forest University, Winston–Salem, NC.He is currently Professor of Surgery at the Uni-

versity of Louisville, Louisville, KY, where he directs research for theDivision of Thoracic and Cardiovascular Surgery. His research interests arein cardiomyoplasty, congenital heart disease, and minimally invasive bypasssurgery.

David T. George (M’89) received the B.S. degreein chemical engineering from the State Universityof New York at Buffalo, Buffalo, NY, in 1983.He received the M.S.E. and Ph.D. degrees in bio-engineering from the University of Pennsylvania,Philadelpha, PA, in 1985 and 1990, respectively.

From 1990–1994, he was a Research Scientist atthe Philadelphia Heart Institute of the PresbyterianMedical Center, Philadelphia, PA. He is presentlyan Assistant Professor of Surgery and Biomedi-cal Engineering, Case Western Reserve Univesity,

Cleveland, OH. His research interests include optimizing skeletal musclefunction for use in cardiac assist, aortomyoplasty, cardiomyoplasty, analysisof ventricular function, and cardiac mechanics.