while the present invention is not directed solely to a cochlear implant , it is appropriate to briefly describe the construction of one type of known cochlear implant system with reference to fig1 . known cochlear implants typically consist of two main components , an external component including a speech processor 29 , and an internal component including an implanted receiver and stimulator unit 22 . the external component includes a microphone 27 . the speech processor 29 is , in this illustration , constructed and 15 arranged so that it can fit behind the outer ear 11 . alternative versions may be worn elsewhere on the recipient &# 39 ; s body . attached to the speech processor 29 is a transmitter coil 24 that transmits electrical signals to the implanted unit 22 via a radio frequency ( rf ) link . the implanted component includes a receiver coil 23 for receiving power and 20 data from the transmitter coil 24 . a cable 21 extends from the implanted receiver and stimulator unit 22 to the cochlea 12 and terminates in an electrode array 20 . the signals thus received are applied by the array 20 to the basilar membrane 8 and the nerve cells within the cochlea 12 thereby stimulating the auditory nerve 9 . the operation of such a device is described , for example , in u . s . pat . no . 4 , 532 , 930 . as depicted diagrammatically in fig1 , the cochlear implant electrode array 20 has traditionally been inserted into the initial portion of the scala tympani of the cochlea 12 up to about a full turn within the cochlea . a sound processor ( not shown ) of the external component 29 includes an amplifier and a speech processor that uses a coding strategy to extract speech from the sounds detected by the microphone 27 . in the depicted embodiment , the speech processor of the cochlear implant can perform an audio spectral analysis of the acoustic signals and output channel amplitude levels . the sound processor can also sort the outputs in order of magnitude , or flag the spectral maxima as used in the speak strategy developed by cochlear ltd . other coding strategies could be employed . fig2 is a circuit diagram illustrating implementation of an embodiment of the present invention in a cochlear implant system of the type shown in fig1 . the speech processor of the external component 29 drives the transmitter coil 24 , which transmits power and data to receiver coil 23 , for the implanted stimulator unit 22 . in accordance with the present invention , a pickup coil 30 is provided for detecting the strength of a magnetic field proximal to the transmitter 24 . the pickup coil 30 is positioned in a plane substantially perpendicular to a primary axis of the magnetic field produced by the transmitter coil 24 and receiver coil 23 . the pickup coil comprises an open circuited single turn , positioned concentrically wi turns of the transmitter coil 24 . a voltage is induced on the pickup coil which is indicative of a magnetic field strength proximal to the transmitter coil 24 . the output of the pickup coil 30 is passed through a peak detector comprising diode d and capacitor c . in the present embodiment , the rf link of the implant system operates at a signal frequency of smhz . the transmitter coil 24 and receiver coil 23 are stagger - tuned to achieve the bandwidth needed for a 100 % amplitude modulated rf signal . the transmitter resonance circuit 24 is usually tuned below the signal frequency , while the implant receiver circuit 23 is tuned slightly above the signal frequency . as a result , the effective impedance seen by the rp drivers of the speech processor of the external component 29 , at the signal frequency , is inductive . this inductive impedance increases when the coupling between the coils 23 , 24 is increased , by reducing the distance between the coils 23 , 24 . as a result , the current through the transmitter coil 24 , and the magnetic field in the vicinity of the coil 24 , falls when the distance between the coils is reduced . this phenomena can also be explained in terms of the interaction between the magnetic fields surrounding the transmitter coil 24 and receiver coil 23 . the magnetic field generated by the receiver coil 23 is a secondary field that opposes the primary field of the transmitter coil 24 . the interaction between the two opposite fields reduces the effective field near the transmitter coil 24 . this effect is increased as the distance between the coils 23 , 24 is reduced . the invention is based on measuring the strength of the magnetic field in the vicinity of the transmitter coil 24 . as this field increases monotonically with the distance between the transmitter coil 24 and receiver coil 23 , the measured field strength can be calibrated to estimate the distance between the coils 23 , 24 , and also to indicate if that distance exceeds a preset value , for example if the coil has fallen off the user &# 39 ; s head . however , other factors , such as battery current , may not change monotonically with varying transceiver separation , and thus do not enable a one - to - one mapping of battery current to transceiver separation , making it impossible to determine transceiver separation by monitoring or measuring such a factor . further , a measurement of magnetic field strength can be performed with very little power consumption , and with very little loading effect on the transmissions between the external and implanted transceivers , thus providing the advantages of simple low current implementation . the circuit shown in fig2 was simulated using orcad pspice version 9 . 2 . the simulation model included circuit models for the c 124 m implant produced by cochlear ltd , esprit 3g speech processor produced by cochlear ltd and a single turn pickup coil . a simplified spice model was used for both the implant and the speech processor . the bsprit 3g model included the major variable that affects and / or sets the battery current , output rf current , stimulation phase width , and intra - frame gap , as well as the rf - data mark - space ratio . the implant model , on the other hand , included all the power consuming components such as the antenna resistance , transformer losses , diode , ic consumption and stimulation current . the coupling coefficient , k , between the transmitter and receiver coils was expressed as a fimction of the distance d between the coils : this value of k was empirically obtained from the particular antennae used in the circuit depicted in fig2 . the peak detector decay time constant was set to 10 ms . this time constant was chosen much longer than the stimulation period of the speak strategy , set to 2000 pps in the spice model . the circuit was simulated using stimulation rates from 2000 pps to 13900 pps , stimulation current ranging from 0 to 1 . 8 ma and link range from 1 to 20 mm . the circuit parameters shown in the following table were used to study the effect of the istance between the coils . parameter value vbatt 3 . 0 v stim rate 13 . 9 khz phase width 25 us stim current 1 ma transmitter coil tuned freq 4 . 8 mhz receiver coil tuned freq 5 . 25 mhz pickup inductance 60 nh coupling coefficient of pickup coil 0 . 7 the simulation results are shown in fig3 and fig4 . fig3 depicts the peak detector output voltage versus link range ( transmitter 24 / receiver 23 separation ). this output voltage depends on the strength of the magnetic field , normal to the pickup coil 30 . in this example , the pickup coil 30 is a single track printed on a pcb upon which the transmitter coil 24 is also printed . the coupling coefficient between the transmitter 24 and pickup coil 30 is assumed to be 0 . 7 . higher coupling can be achieved in practice by the careful placement of the pickup coil 30 relative to the transmitter 24 . higher output signals can also be obtained if a two - turn ( or more ) pickup coil is used . fig3 reveals that the magnetic field of the transmitter 24 increases with the distance between the transmitter coil 24 and receiver coil 23 ( link range ). when that distance exceeds 20 mm , the output voltage reaches about 780 mv ( not shown in the figure ). fig3 also reveals that the increase in magnetic field is monotonic as the link range increases from 1 mm to 10 mm . fig4 depicts the battery current which reaches a peak value of 18 . 9 ma at 4 mm then gradually drops to 18 ma at 10 mm , and to 17 ma at 20 mm ( not shown in the figure ). thus , the battery current does not vary monotonically with increasing link range between the transmitter 24 and receiver 23 . fig3 and 4 clearly show that the battery current cannot be used to estimate the link range , as a given value of battery current can not be equated to a single value of transceiver separation . on the other hand , there is a one to one correlation between the output voltage of the peak detector c , d and the distance between the transmitter coil 24 and receiver coil 23 . it is to be noted that the battery current is proportional to the total system power . on the other hand , the strength of the magnetic field in the vicinity of the transmitter coil 24 is proportional to the stored reactive energy . the relationship between the active and reactive energy components depends on the phase angle of the coil current relative to the driving voltage . it is this phase angle which changes with the coupling coefficient between the transmitter coil 24 and receiver coil 23 . the peak detector output depends slightly on the implant power , as explained below with respect to fig5 to 10 . the effect of the implant power on the output of the peak detector becomes negligible at maximum link range . as the distance between the coils 23 , 24 is gradually increased from minimum to maximum link range , a number of effects occur . firstly , the power delivered to the implant 22 is reduced . secondly , transmitter losses increase due to increased rf current . these changes determine the behaviour of the battery current , whereas the current through the transmitter coils 23 , 24 , and hence the magnetic field strength , increases monotonically towards an asymptotic value . the peak magnetic field , normal to the pickup coil 30 , depends on the sum of the electric fields produced by the transmitter coil 24 and receiver coil 23 . the peak magnetic field depends slightly on the stimulation parameters , namely the stimulation current and the stimulation rate . the influence of the stimulation parameters is relatively small because the stimulation power represents a small part of the total system power which includes the implant 22 and transmitter coil 24 losses , as follows : on the other hand , the ratio of the stored to dissipated energy is the effective quality factor of the loaded transmitter coil , as follows : the magnetic field is proportional to the reactive power , which is much higher than the stimulation power . therefore , the stimulation parameters can only have a second order effect on the peak amplitude of the magnetic field . this is in agreement with the simulation results . the stimulation rate , however , has a stronger effect due to the fact that the peak detector used in fig2 is not ideal and has a finite decay time constant . the significance of the above discussion is to highlight the fact that , at long link range , the peak detector output is not sensitive to the stimulation current , but is affected by the stimulation rate . this effect must be taken into account when the peak detector output is used to estimate the distance between the coils . one application of the present invention is in estimating a skin flap thickness of a recipient of a cochlear implant system of the type shown in fig1 , that is , the thickness of skin between the implanted receiver coil 23 and the external transmitter coil 24 . to date , estimating the skin flap thickness has been done in a clinic where the speech processor is powered from the programning system . in this case , specific stimulation parameters are used in order to achieve consistent and repeatable sldn flap thickness estimates . however , the circuit of fig2 can be used to estimate the skin flap thickness . a first method by which the skin flap thickness may be estimated by using the circuit of fig2 , involves using the recipient &# 39 ; s own speech processor to create the rf magnetic field . this method requires providing a signal path from the peak detector output to the progranmming system . in this case , the transmitter coil is excited with maximum frame rate at a regulated supply voltage supplied by the programming system . this eliminates the dependency of the peak detector output on the stimulation rate and supply voltage . a look up table stored in the programming system can be used to map the measured voltage to skin flap thickness . a second method by which the skin flap thickness may be estimated by using the circuit of fig2 , involves using a stand - alone device with built - in oscillator and voltage measurement circuit . in this second method , the stand - alone device is essentially a skin flap thickness meter . the meter contains a 5 mhz crystal oscillator with low output impedance drivers to drive a tuned transmitter coil with continuous 5 mhz square voltage . the transmitter coil contains a pickup coil and a peak detector similar to that shown in fig2 . the dc output of the peak detector is measured using a built - in analog to digital converter ( adc ). the output of the adc is converted to skin flap thickness , which is then displayed by the meter . another application of the present invention is in detecting displacement of the external transmitter 24 from the user &# 39 ; s head , for example where the transmitter coil 24 falls off an infant &# 39 ; s head . such coil - off detection is based on detecting a link range greater than a set threshold value , which would typically be set to around 10 - 12 mm . such a circuit solution has to be implemented on the transmitter coil and / or the speech processor . for reliable detection , the circuit has a low sensitivity to battery voltage , stimulation current , stimulation rate , ambient temperature and implant tuning . the circuit also operates without requiring precision measurement of the output voltage of the peak detector . the circuit solution is simple , uses a small number of components and has low current consumption . one manner in which many or all of the above requirements may be met is by comparing the peak detector signal with another reference signal , which has all of the major characteristics of the peak detector signal except its dependency on the coil separation . the reference signal should be generated from a peak detector similar to that shown in fig2 in order to have the same decay time constant , voltage offsets and temperature characteristics as the measured signal , and should be proportional to the battery voltage to track the changes of the measured signal with the battery voltage . further , the reference signal should vary with the stimulation rate in a manner similar to that of the measured signal , and should have low sensitivity to the implant power , especially at relatively large link ranges . a simple manner in which the reference signal can be obtained comprises rectifying and peak - detecting the output of the rf drivers of the speech processor , as shown in fig5 . in fig5 , the output of the speech processor , in this instance an esprit 3g speech processor of the type produced by cochlear ltd , is flull - wave rectified by d 1 and d 2 . the dc voltage across c 2 tracks the amplitude of the esprit 3g rf output voltage . this dc voltage can be made to vary with the stimulation rate in a manner which is similar to that of the voltage across r 1 . this is determined by the time constant : when this time constant is made very small , the voltage across c 2 will strongly depend on the stimulation rate , and vice versa . the voltage divider ratio r 3 /( r 2 + r 3 ) is designed such that the voltage across r 3 is substantially equal to the peak voltage across r 1 at the designated threshold for the maximum link range . the voltage across r 3 is applied to a diode - capacitor ( d 4 , c 3 ) peak detector similar to that used with the pickup coil 30 . this is to match the time variation and the temperature characteristics of the measured signal and the reference signal . a voltage comparator is used to compare the measured and reference signals . the output of the comparator can be used to trigger an audible alarm to alert the carers if the transmitter coil is removed . the way the circuit operates is based on matching the amplitudes of the measured and reference signals at the maximum link range . below that range , the measured signal is smaller and the output of the comparator is disasserted . however , if the separation between the transmitter coil 24 and receiver coil 23 exceeds the maximum link range , the measured signal exceeds the reference signal and triggers the comparator . the recommended component values for typical circuit conditions of the esprit 3g are given below . d 1 to d 4 : low cut - in voltage high - speed diodes r 1 = r 4 = 1mω r 2 = 220 kω r 3 ≅ 100 kω c 1 = c 3 = 10 nf c 2 = 100 pf pickup coil : printed single turn on the transmitter coil pcb . a single turn from an electrostatic shield can be used . where the transmitter coil is implemented on a printed circuit board , the circuit of fig5 can be fully integrated on the pcb of the transmitter coil . the comparator can be replaced with a low voltage - low power low speed operational amplifier . the dc power for the comparator / amplifier can be provided from the rf drivers &# 39 ; signal using a voltage doubler circuit to provide the amplifier with positive and negative dc supply rails . a power cost will be in overloading the rf drivers with the comparator dc power , which can be as low as 50 ua at 3v . however , this is an insignificant cost compared with the total rf power consumed by the system . the advantage of integrating the circuit on the transmitter coil is that it reduces the number of the coil cable connectors , and substantially guarantees the matching between the circuit components especially with respect to changes with temperature . fig6 is a circuit diagram of a circuit used for verification of coil - off detection , for use with an esprit 3g speech processor . the circuit of fig6 was used to investigate and verify the concept and to study the sensitivity to different circuit and stimulation parameters . the prototype was measured in a laboratory with both speak and 14 . 2 khz stimulation , at both quiet and loud sound environments , and at different battery voltages . the circuit is designed for minimum loading on the rf drivers of the esprit 3g . it uses a small number of components which can be all mounted on the transmitter coil printed circuit board . the transmitter coil has 3 open tracks on each side used for electrostatic shielding . one shield track ( nearly a full turn ) is used as the pickup coil . r 1 , r 2 and c 2 form a potential divider and a low pass filter for the rf signal on rfout 1 . the filter parameters are chosen such that the peak voltage across c 2 varies with the pulse width of the rf signal . this allows the output v 1 to track the rf power level at different battery voltages . in fig6 r 1 is a variable resistor to facilitate accurate adjustment for the best detection thresholds . in a non - testing circuit , it is expected that r 1 will be replaced with a fixed resistor . the voltage across c 10 and the voltage across the pickup coil l 2 are peak detected using identical envelope detectors . the dc output v 1 is the coil - off detection voltage threshold . v 2 is the coil - off signal . the voltage v 2 increases as the separation between the transmitter coil and the implanted receiver coil increases . at or above the coil - off detection distance , v 2 exceeds v 1 . the measurement method was as follows . the esprit 3g was loaded with 2 patient maps . the first was a 14 . 2 kpps map while the second was a speak 2 kpps map . a “ quiet ” sound condition was simulated by removing the microphone and replacing it with a 1 kω resistor . a “ loud ” sound condition was simulated by placing a loud radio close to the microphone . the voltages v 1 and v 2 were measured under the conditions shown in table 1 below . each of the following tests was carried out at room temperature . a total of 40 tests ( table 1 ) were carried out . during each test the distance was varied from 0 to 14 mm in 2 mm steps , after which the distance was set to more than 10 cm ( simulating very large distance ). these 40 tests cover the different circuit parameters , in order to demonstrate the sensitivity of the coil - detection method to these parameters . at each distance , the test was repeated 4 times ; at stimulation rates of 2000 pps and 14400 pps , and in both “ quiet ” and “ loud ” sound environments . the measurements were also repeated at different implant tuning frequencies of 5 . 1 mhz , 5 . 25 mhz and 5 . 4 mhz , and at different supply voltages of 2 . 7v , 3 . 0v and 3 . 3v . to check the sensitivity to the transmitter coil tuning the test was repeated for implant tuning of 5 . 25 mhz and power supply voltage of 3v . the transmitter coil was tuned to its minimum limit and then to its maximum limit of 4 . 725 mhz and 4 . 775 mhz respectively . the test results are set out towards the end of the present specification . the distances at which the measured signal ( v 2 ) exceeds the threshold voltage ( v 1 ) are highlighted in the results tables . because the measurements were done at increments of 2 mm , the highlighted points could be equal to or exceed the correct detection point by up to 2 mm . fig7 illustrates the reference and measured voltages , v 1 and v 2 respectively , at 14 . 2 kpps and 2 kpps in quiet and loud sound environments . the battery voltage was set to 3 . 3 v . the implanted coil was tuned to its nominal frequency of 5 . 25 mhz . fig7 shows that the reference voltage is automatically adjusted to a threshold distance of between 12 and 13 mm . above this threshold , an alarm will be triggered to indicate a coil - off condition . fig8 depicts the reference and measured voltages , v 1 and v 2 respectively , at 14 . 2 kpps in quiet and loud sound environments , and at supply voltages of 3 . 3 , 3 . 0 and 2 . 7v respectively . the implanted coil was tuned to its nominal frequency of 5 . 25 mhz . these results indicate the detection distance has low sensitivity to the supply voltage , as the point of intersection of the v 1 and v 2 curves varies by only small amounts . fig9 reveals that the coil - off detection distance is reasonably sensitive to the tuning frequency of the implanted coil . when the implant is tuned to 5 . 4 mhz , the detection threshold distance drops to 8 . 5 mm . the detection distance increases as the tuning frequency of the implanted coil is reduced to 5 . 1 mhz . at this frequency , the circuit will detect coil removal if the distance exceeds about 14 mm . the effect of the transmitter coil tuning is shown in fig1 . the results , at 3v supply voltage and 14 . 2 khz stimulation rate , indicate that varying the transmitter coil tuning from 4 . 725 mhz to 4 . 775 mhz has substantially no effect on the distance threshold . high rate stimulation 3 10 . 8 15 9 . 5 27 6 . 3 4 11 . 4 16 9 . 5 28 6 . 2 7 11 . 8 19 9 . 6 31 6 . 2 8 11 . 8 20 9 . 6 32 6 . 2 11 14 . 9 23 11 . 9 35 8 . 5 12 15 24 12 . 3 36 8 . 5 the above table shows the coil - off detection threshold distance at all combinations of supply voltage and tuning frequencies . similar to the high rate stimulation , the lowest detection distance occurred at low battery voltage and high implant tuning frequency ( tests 25 , 26 , 29 and 30 ). 1 6 . 5 13 6 . 7 25 5 . 3 2 8 . 4 14 7 . 9 26 5 . 6 5 7 . 0 17 6 . 9 29 5 . 5 6 9 . 3 18 8 . 8 30 6 . 0 9 13 . 7 21 12 33 9 . 1 10 13 . 9 22 12 . 7 34 9 . 2 the measurement results discussed above show the usefulness of the coil - off detection circuit embodiment of the present invention . the method discussed has low sensitivity to most of the circuit parameters and variables , except for the implant tuning if at the upper end of the tuning range . this problem can be easily solved by adding a small dc offset to the reference voltage v 1 . by adjusting the value of that offset a detection distance in the range 8 mm to 15 mm can be achieved for all circuit conditions . while an embodiment of the invention has been discussed in which a threshold detection of a coil - off condition is performed , it is to be appreciated that alternative embodiments of the present invention may be used to estimate an actual distance between implanted and external coils . for example , a look - up table may be experimentally derived from a voltage to distance calibration measurement , such as the voltage measurements revealed in fig7 to 10 , such a look - up table may then be used in converting measured magnetic field strengths to estimated transceiver separation values . alternatively , a best - fit algorithm may be derived from the measured voltage / distance values , for use in converting measured magnetic field strengths to estimated transceiver separation values . appendix test results test 1 test 2 test 3 test 4 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 254 29 306 46 412 93 463 136 2 258 42 323 65 435 108 502 157 4 248 90 341 107 453 146 532 194 6 227 196 343 177 473 205 575 267 8 228 311 283 270 423 292 519 330 10 230 387 302 359 414 388 494 436 12 227 422 288 415 425 463 482 506 14 225 446 286 462 446 537 478 566 & gt ; 100 222 565 270 609 440 789 457 811 test 5 test 6 test 7 test 8 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 323 56 405 93 634 262 635 263 2 329 72 405 105 664 274 664 275 4 318 129 419 155 706 316 707 318 6 298 234 415 236 746 391 746 391 8 297 367 400 339 732 477 733 479 10 300 470 408 442 673 586 673 586 12 298 509 409 512 662 671 662 671 14 296 536 412 566 654 741 654 740 & gt ; 100 292 654 338 746 627 1005 632 1010 test 9 test 10 test 11 test 12 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 544 137 633 190 773 340 773 340 2 530 131 638 192 796 335 797 336 4 553 169 654 223 835 356 833 357 6 569 249 685 294 874 421 874 422 8 585 359 645 382 865 499 865 499 10 599 473 630 487 811 607 811 607 12 605 555 631 569 802 697 802 697 14 612 621 631 635 797 773 797 772 & gt ; 100 612 850 623 862 774 1034 774 1033 test 13 test 14 test 15 test 16 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 254 35 349 74 414 105 473 152 2 259 57 290 70 430 138 485 180 4 266 112 306 124 436 194 475 227 6 224 197 263 200 424 269 476 312 8 226 280 278 283 417 358 472 409 10 232 340 272 345 410 431 464 486 12 225 338 264 389 443 524 463 537 14 223 418 276 448 439 581 462 608 & gt ; 100 220 565 267 612 436 786 456 813 test 17 test 18 test 19 test 20 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 328 67 415 114 641 276 637 274 2 334 100 428 145 666 326 661 316 4 334 176 438 208 679 409 674 385 6 297 257 459 310 681 487 676 478 8 298 351 490 372 663 579 660 574 10 304 422 420 462 646 664 645 662 12 299 467 405 518 640 730 647 737 14 296 506 427 592 632 793 640 801 & gt ; 100 292 659 432 793 624 1009 635 1022 test 21 test 22 test 23 test 24 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 585 177 646 216 775 348 774 345 2 593 206 660 248 797 373 797 377 4 605 264 653 297 812 437 810 436 6 601 346 662 383 815 522 815 524 8 588 485 647 462 798 616 805 610 10 593 529 642 558 785 712 790 711 12 594 593 640 618 778 780 786 774 14 594 648 636 677 773 838 780 838 & gt ; 100 594 856 631 874 762 1040 771 1038 test 25 test 26 test 27 test 28 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 256 50 293 68 417 134 463 173 2 256 93 317 120 416 200 458 242 4 257 170 285 192 408 296 456 350 6 224 259 260 280 410 396 447 439 8 229 327 265 351 412 488 452 531 10 229 383 263 406 438 578 461 610 12 224 426 263 457 439 629 452 646 14 225 466 258 492 439 671 455 692 & gt ; 100 222 582 255 615 440 809 458 833 test 29 test 30 test 31 test 32 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 325 88 388 124 637 315 637 316 2 326 141 388 185 647 409 648 409 4 329 226 406 290 641 517 641 517 6 293 327 380 379 636 624 636 625 8 298 407 382 470 631 721 631 721 10 299 461 395 534 629 785 629 786 12 295 509 388 581 627 842 627 842 14 294 553 392 634 628 893 628 893 & gt ; 100 292 678 378 758 627 1038 627 1039 test 33 test 34 test 35 test 36 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 578 204 636 245 777 383 777 386 2 548 248 644 326 790 474 791 474 4 570 366 620 406 786 580 786 579 6 572 462 622 499 783 675 783 675 8 594 556 628 585 781 760 781 760 10 606 636 625 654 778 836 778 836 12 608 690 624 705 777 889 777 888 14 607 734 628 754 776 933 776 933 & gt ; 100 610 878 623 891 774 1070 774 1070 test 37 test 38 test 39 test 40 2 . 0 khz quiet 2 . 0 khz loud 14 . 2 khz quiet 14 . 2 khz loud distance mm v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) v1 ( mv ) v2 ( mv ) 0 325 71 405 115 636 286 637 287 2 330 98 417 146 659 326 659 327 4 341 166 428 208 671 398 671 398 6 296 259 423 295 673 488 673 489 8 294 342 405 378 658 578 658 578 10 301 403 407 447 639 659 640 660 12 297 442 412 500 632 716 632 716 14 293 475 408 549 626 768 627 768 & gt ; 100 290 600 403 695 619 943 619 943 it will be appreciated by persons skilled in the art that numerous variations and / or modifications may be made to the invention as shown in the specific embodiments without departing from the spirit or scope of the invention as broadly described . the present embodiments are , therefore , to be considered in all respects as illustrative and not restrictive .
a method and apparatus for determining a position of an external transceiver relative to an implanted transceiver comprising means for measuring the strength of a magnetic field proximal to thc external transceiver and means for determining a position of the external transceiver relative to the implanted transceiver from said measured magnetic field strength . furthermore there is disclosed is a method and apparatus for determining a skin flap thickness of a recipient of a prosthesis including a transcutaneous link between the external transceiver and the implanted transceiver . a skin - flap thickness meter is also provided .