Patent Publication Number: US-8971559-B2

Title: Switching structures for hearing aid

Description:
RELATED APPLICATIONS 
     This application is a continuation of U.S. application Ser. No. 12/107,643, filed Apr. 22, 2008, which is a divisional of U.S. application Ser. No. 10/244,295, filed Sep. 16, 2002, both of which are incorporated by reference herein in their entirety. 
     This application is generally related to U.S. application Ser. No. 09/659,214 filed Sep. 11, 2000 (now U.S. Pat. No. 6,760,457), which is hereby incorporated by reference. 
     This application is generally related to U.S. application Ser. No. 10/243,412 filed Sep. 12, 2002, which is hereby incorporated by reference. 
    
    
     FIELD OF THE INVENTION 
     This invention relates generally to hearing aids, and more particularly to switching structures and systems for a hearing aid. 
     BACKGROUND 
     Hearing aids can provide adjustable operational modes or characteristics that improve the performance of the hearing aid for a specific person or in a specific environment. Some of the operational characteristics are volume control, tone control, and selective signal input. One way to control these characteristics is by a manually engagable switch on the hearing aid. The hearing aid may include both a non-directional microphone and a directional microphone in a single hearing aid. Thus, when a person is talking to someone in a crowded room the hearing aid can be switched to the directional microphone in an attempt to directionally focus the reception of the hearing aid and prevent amplification of unwanted sounds from the surrounding environment. However, a conventional switch on the hearing aid is a switch that must be operated by hand. It can be a drawback to require manual or mechanical operation of a switch to change the input or operational characteristics of a hearing aid. Moreover, manually engaging a switch in a hearing aid that is mounted within the ear canal is difficult, and may be impossible, for people with impaired finger dexterity. 
     In some known hearing aids, magnetically activated switches are controlled through the use of magnetic actuators. For examples, see U.S. Pat. Nos. 5,553,152 and 5,659,621. The magnetic actuator is held adjacent the hearing aid and the magnetic switch changes the volume. However, such a hearing aid requires that a person have the magnetic actuator available when it desired to change the volume. Consequently, a person must carry an additional piece of equipment to control his\her hearing aid. Moreover, there are instances where a person may not have the magnetic actuator immediately present, for example, when in the yard or around the house. 
     Once the actuator is located and placed adjacent the hearing aid, this type of circuitry for changing the volume must cycle through the volume to arrive at the desired setting. Such an action takes time and adequate time may not be available to cycle through the settings to arrive at the required setting, for example, there may be insufficient time to arrive at the required volume when answering a telephone. 
     Some hearing aids have an input which receives the electromagnetic voice signal directly from the voice coil of a telephone instead of receiving the acoustic signal emanating from the telephone speaker. Accordingly, signal conversion steps, namely, from electromagnetic to acoustic and acoustic back to electromagnetic, are removed and a higher quality voice signal reproduction may be transmitted to the person wearing the hearing aid. It may be desirable to quickly switch the hearing aid from a microphone (acoustic) input to a coil (electromagnetic field) input when answering and talking on a telephone. However, quickly manually switching the input of the hearing aid from a microphone to a voice coil, by a manual mechanical switch or by a magnetic actuator, may be difficult for some hearing aid wearers. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       A more complete understanding of the invention and its various features, objects and advantages may be obtained from a consideration of the following detailed description, the appended claims, and the attached drawings in which: 
         FIG. 1  illustrates the hearing aid of the present invention adjacent a magnetic field source; 
         FIG. 2  is a schematic view of the  FIG. 1  hearing aid; 
         FIG. 3  shows a diagram of the switching circuit of  FIG. 2 ; 
         FIG. 4  is a schematic view of a hearing aid according to an embodiment of the present invention; 
         FIG. 5  is a schematic view of a hearing aid according to an embodiment of the present invention; 
         FIG. 6  is a schematic view of a hearing aid according to an embodiment of the present invention; 
         FIG. 7  is a schematic view of a hearing aid according to an embodiment of the present invention; 
         FIG. 8  is a schematic view of a hearing aid according to an embodiment of the present invention; 
         FIG. 9  is a schematic view of a hearing aid according to an embodiment of the present invention; 
         FIG. 10  is a schematic view of an embodiment of the present invention; 
         FIG. 11  is a circuit diagram of a power source of an embodiment of the present invention; 
         FIG. 12  is a circuit diagram of an embodiment of the present invention; 
         FIG. 13  is a circuit diagram of an embodiment of the present invention; 
         FIG. 14  is a schematic view of a hearing aid cleaning and charging system according to an embodiment of the present invention; and 
         FIG. 15  is a view of hearing aid switch of the present invention and a comparator/indicator circuit. 
         FIG. 16  is a diagram of a switching circuit according to an embodiment of the present invention. 
         FIG. 17  is a diagram of a switching circuit according to an embodiment of the present invention. 
         FIG. 18  is a diagram of a switching circuit according to an embodiment of the present invention. 
         FIG. 19  is a diagram of a switching circuit according to an embodiment of the present invention. 
         FIG. 20  is a diagram of a switching circuit according to an embodiment of the present invention. 
         FIG. 21  is a diagram of a switching circuit according to an embodiment of the present invention. 
         FIG. 22  is a diagram of a switching circuit according to an embodiment of the present invention. 
     
    
    
     DETAILED DESCRIPTION 
     In the following detailed description, reference is made to the accompanying drawings which form a part hereof and in which are shown by way of illustration specific embodiments in which the invention can be practiced. These embodiments are described in sufficient detail to enable those skilled in the art to practice and use the invention, and it is to be understood that other embodiments may be utilized and that electrical, logical, and structural changes may be made without departing from the spirit and scope of the present invention. The following detailed description is, therefore, not to be taken in a limiting sense and the scope of the present invention is defined by the appended claims and their equivalents. 
     Hearing aids provide different hearing assistance functions including, but not limited to, directional and non-directional inputs, multi-source inputs, filtering and multiple output settings. Hearing aids are also provide user specific and/or left or right ear specific functions such as frequency response, volume, varying inputs and signal processing. Accordingly, a hearing aid is programmable with respect to these functions or switch between functions based on the operating environment and the user&#39;s hearing assistance needs. A hearing aid is described that includes magnetically operated switches and programming structures. 
     One embodiment of the present invention provides a hearing aid that includes an input system, an output system, a signal processing circuit electrically connecting the input system to the output system, a magnetically actuatable switch between the input system and the signal processing circuit, and a filter connected to and controlled by the magnetically-actuatable switch. The switch allows the filter to filter a signal from the input system to the signal processing circuit or prevents the filter from filtering the signal. In an embodiment, the switch is a solid state switch. In an embodiment, the solid state switch is a giant magneto resistive (GMR) switch. In an embodiment, the solid state switch is an anisotropic magneto resistive (AMR) switch. In an embodiment, the solid state switch is a magnetic field effect transistor. 
     In an embodiment of the present invention, a magnetically actuatable switch is positioned between the output system and the signal processing circuit. This switch controls operation of a device before the output system or at the output system. In an embodiment, the switch selectively connects an output filter that filters the signal received by the output system. In an embodiment, the hearing aid includes a plurality of filters that are selectable based on the magnetic field sensed by the magnet switch or a magnetic field sensor. 
     An embodiment of the present invention provides a hearing aid that includes an input system, an output system, a programmable, signal processing circuit electrically connecting the input system to the output system, a magnetic field sensor, and a selection circuit connected to the magnetic sensor and at least one of the input system, output system and the signal processing system. The selection circuit is adapted to control the at least one of the input system, output system and the signal processing system based on a signal produced by the magnetic field sensor. The selection circuit is adapted to receive an electrical signal from the magnetic sensor and supply a programming signal to the signal processing circuit. In an embodiment, the magnetic field sensor is a full bridge circuit. In an embodiment, the magnetic field sensor is adapted to receive a pulsed power supply. In an embodiment, the selection circuit is connected to the input system and sends a control signal to the input system based on a signal received from the magnetic field sensor. In an embodiment, the input system includes a first input and a second input, and the input system activates one of the first input and the second input based on the control signal. The first input includes a microphone. The second input includes a magnetic field sensing device. The hearing aid of the present invention further includes a threshold circuit that blocks signals below a threshold value. 
     An embodiment of the present invention provides a hearing aid that includes a programming system that is adapted to sense a magnetic field and based on the magnetic field produce a programming signal. The programming signal, in an embodiment, includes a control sequence or code that allows the hearing aid to be programmed. The programming signal further includes a digital programming signal based on the magnetic field sensed by a magnetic field sensor. 
     An embodiment of the present invention includes a wireless on/off switch. The wireless on/off switch includes a magnetically operable switch. In an embodiment, the magnetically operable switch is a solid state switch. The on/off switch turns off the non-essential power to the hearing aid circuits to preserve battery power. In an embodiment, a system is provided that stores the hearing aid and provides a signal to turn off the hearing aid. 
     An embodiment of the invention includes a wireless switch that activates a power induction circuit in the hearing aid. The power induction circuit is adapted to receive a recharging signal from a power source and recharge the hearing aid power source. In an embodiment, the wireless switch that activates the power induction circuit also turns off the non-essential power consuming circuits of the hearing aid. 
     An embodiment of the invention includes a system that has a magnetic field source. In an embodiment, the magnetic field source being adapted to program the hearing aid. In an embodiment, the magnetic field source is adapted to wirelessly turn off and turn on the hearing aid. The system includes a storage receptacle for the hearing aid. In an embodiment, the magnetic field source provides a power induction signal that is adapted to recharge the hearing aid power source. 
       FIG. 1  illustrates an in-the-ear hearing aid  10  that is positioned completely in the ear canal  12 . A telephone handset  14  is positioned adjacent the ear  16  and, more particularly, the speaker  18  of the handset is adjacent the pinna  19  of ear  16 . Speaker  18  includes an electromagnetic transducer  21  which includes a permanent magnet  22  and a voice coil  23  fixed to a speaker cone (not shown). Briefly, the voice coil  23  receives the time-varying component of the electrical voice signal and moves relative to the stationary magnet  22 . The speaker cone moves with coil  23  and creates an audio pressure wave (“acoustic signal”). It has been found that when a person wearing a hearing aid uses a telephone it is more efficient for the hearing aid  10  to pick up the voice signal from the magnetic field gradient produced by the voice coil  23  and not the acoustic signal produced by the speaker cone. 
     Hearing aid  10  has two inputs, a microphone  31  and a voice coil pickup  32  ( FIG. 2 ). The microphone  31  receives acoustic signals, converts them into electrical signals and transmits same to a signal processing circuit  34 . The signal processing circuit  34  provides various signal processing functions which can include noise reduction, amplification, and tone control. The signal processing circuit  34  outputs an electrical signal to an output speaker  36  which transmits audio into the wearer&#39;s ear. The voice coil pickup  32  is an electromagnetic transducer, which senses the magnetic field gradient produced by movement of the telephone voice coil  23  and in turn produces a corresponding electrical signal which is transmitted to the signal processing circuit  34 . Accordingly, use of the voice coil pickup  32  eliminates two of the signal conversions normally necessary when a conventional hearing aid is used with a telephone, namely, the telephone handset  14  producing an acoustic signal and the hearing aid microphone  31  converting the acoustic signal to an electrical signal. It is believed that the elimination of these signal conversions improves the sound quality that a user will hear from the hearing aid. 
     A switching circuit  40  is provided to switch the hearing aid input from the microphone  31 , the default state, to the voice coil pickup  32 , the magnetic field sensing state. It is desired to automatically switch the states of the hearing aid  10  when the telephone handset  14  is adjacent the hearing aid wearer&#39;s ear. Thereby, the need for the wearer to manually switch the input state of the hearing aid when answering a telephone call and after the call is ends. Finding and changing the state of the switch on a miniaturized hearing aid can be difficult especially when the wearer is under the time constraints of a ringing telephone or if the hearing aid is an in the ear type hearing aid. 
     The switching circuit  40  of the described embodiment changes state when in the presence of the telephone handset magnet  22 , which produces a constant magnetic field that switches the hearing aid input from the microphone  31  to the voice coil pickup  32 . As shown in  FIG. 3 , the switching circuit  40  includes a microphone activating first switch  51 , here shown as a transistor that has its collector connected to the microphone ground, base connected to a hearing aid voltage source through a resistor  58 , and emitter connected to ground. Thus, the default state of hearing aid  10  is switch  58  being on and the microphone circuit being complete. A second switch  52  is also shown as a transistor that has its collector connected to the hearing aid voltage source through a resistor  59 , base connected to the hearing aid voltage source through resistor  58 , and emitter connected to ground. A voice coil activating third switch  53  is also shown as a transistor that has its collector connected to the voice pick up ground, base connected to the collector of switch  52  and through resistor  59  to the hearing aid voltage source, and emitter connected to ground. A magnetically activated fourth switch  55  has one contact connected to the base of first switch  51  and through resistor  58  to the hearing aid voltage source, and the other contact is connected to ground. Contacts of switch  55  are normally open. 
     In this default open state of switch  55 , switches  51  and  52  are conducting. Therefore, switch  51  completes the circuit connecting microphone  31  to the signal processing circuit  34 . Switch  52  connects resistor  59  to ground and draws the voltage away from the base of switch  53  so that switch  53  is open and not conducting. Accordingly, hearing aid  10  is operating with microphone  31  active and the voice coil pickup  32  inactive. 
     Switch  55  is closed in the presence of a magnetic field, particularly in the presence of the magnetic field produced by telephone handset magnet  22 . In one embodiment of the invention, switch  55  is a reed switch, for example a microminiature reed switch, type HSR-003 manufactured by Hermetic Switch, Inc. of Chickasha, Okla. In a further embodiment of the invention, the switch  55  is a solid state, wirelessly operable switch. In an embodiment, wirelessly refers to a magnetic signal. An embodiment of a magnetic signal operable switch is a MAGFET. The MAGFET is non-conducting in a magnetic field that is not strong enough to turn on the device and is conducting in a magnetic field of sufficient strength to turn on the MAGFET. In a further embodiment, switch  55  is a micro-electro-mechanical system (MEMS) switch. In a further embodiment, the switch  55  is a magneto resistive device that has a large resistance in the absence of a magnetic field and has a very small resistance in the presence of a magnetic field. When the telephone handset magnet  22  is close enough to the hearing aid wearer&#39;s ear, the magnetic field produced by magnet  22  changes the state of switch (e.g., closes) switch  55 . Consequently, the base of switch  51  and the base of switch  52  are now grounded. Switches  51  and  52  stop conducting and microphone ground is no longer grounded. That is, the microphone circuit is open. Now switch  52  no longer draws the current away from the base of switch  53  and same is energized by the hearing aid voltage source through resistor  59 . Switch  53  is now conducting. Switch  53  connects the voice pickup coil ground to ground and completes the circuit including the voice coil pickup  32  and signal processing circuit  34 . Accordingly, the switching circuit  40  activates either the microphone (default) input  31  or the voice coil (magnetic field selected) input  32  but not both inputs simultaneously. 
     In operation, switch  55  automatically closes and conducts when it is in the presence of the magnetic field produced by telephone handset magnet  22 . This eliminates the need for the hearing aid wearer to find the switch, manually change switch state, and then answer the telephone. The wearer can conveniently, merely pickup the telephone handset and place it by his\her ear whereby hearing aid  10  automatically switches from receiving microphone (acoustic) input to receiving pickup coil (electromagnetic) input. That is, a static electro-magnetic field causes the hearing aid to switch from an audio input to a time-varying electro-magnetic field input. Additionally, hearing aid  10  automatically switches back to microphone input after the telephone handset  14  is removed from the ear. This is not only advantageous when the telephone conversation is complete but also when the wearer needs to talk with someone present (microphone input) and then return to talk with the person on the phone (voice coil input). 
     The above described embodiment of the switching circuit  40  describes a circuit that grounds an input and open circuits the other inputs. It will be recognized that the switching circuit  40 , in an embodiment, connects the power source to an input and disconnects the power source to the other inputs. For example, the collectors of the transistors  51  and  53  are connected to the power source. The switch  55  remains connected to ground. The emitter of transistor  51  is connected to the power input of the microphone  31 . The emitter of the transistor  53  is connected to the power input of the voice coil  32 . Thus, switching the switch  55  causes the power source to be interrupted to the microphone and supplied to the voice coil pickup  32 . In an embodiment, switching circuit  40  electrically connects the signal from one input to the processing circuit  34  and opens (disconnects) the other inputs from the processing circuit  34 . 
     While the disclosed embodiment references an in-the-ear hearing aid, it will be recognized that the inventive features of the present invention are adaptable to other styles of hearing aids including over-the-ear, behind-the-ear, eye glass mount, implants, body worn aids, etc. Due to the miniaturization of hearing aids, the present invention is advantageous to many miniaturized hearing aids. 
       FIG. 4  shows hearing aid  70 . The hearing aid  70  includes a switching circuit  40 , a signal processing circuit  34  and an output speaker  36  as described herein. The switching circuit  40  includes a magnetic field responsive, solid state circuit. The switching circuit  40  selects between a first input  71  and a second input  72 . In an embodiment, the first input  71  is an omnidirectional microphone, which detects acoustical signals in a broad pattern. In an embodiment, the second input  72  is a directional microphone, which detects acoustical signals in a narrow pattern. The omnidirectional, first input  71  is the default state of the hearing aid  70 . When the switching circuit  40  senses the magnetic field, the switch changes state from its default to a magnetic field sensed state. The magnetic field sensed state causes the hearing aid  70  to switch from its default mode and the directional, second input  72  is activated. In an embodiment, the activation of the second input  72  is mutually exclusive of activation of the first input  71 . 
     In use with a telephone handset, e.g.,  14  shown in  FIG. 1 , hearing aid  70  changes from its default state with omnidirectional input  71  active to its directional state with directional input  72  active. Thus, hearing aid  70  receives its input acoustically from the telephone handset. In an embodiment, the directional input  72  is tuned to receive signals from a telephone handset. 
     In an embodiment, switching circuit  40  includes a micro-electro-mechanical system (MEMS) switch. The MEMS switch includes a cantilevered arm that in a first position completes an electrical connection and in a second position opens the electrical connection. When used in the circuit as shown in  FIG. 3 , the MEMS switch is used as switch  55  and has a normally open position. When in the presence of a magnetic field, the cantilevered arm shorts the power supply to ground. This initiates a change in the operating state of the hearing aid input. 
       FIG. 5  shows an embodiment of a hearing aid  80  according to the teachings of the present invention. Hearing aid  80  includes at least one input  81  connected to a signal processing circuit  34 , which is connected to an output speaker  36 . In an embodiment, hearing aid  80  includes two or more inputs  81  (one shown). The input  81  includes a signal receiver  83  that includes two nodes  84 ,  85 . Node  84  is connected to the signal processing circuit  34  and to one terminal of a capacitor  86 . In an embodiment, node  84  is the negative terminal of the input  81 . In an embodiment, node  84  is the ground terminal of the input  81 . Node  85  is connected to one pole of a magnetically operable switch  87 . In an embodiment, the switch  87  is a mechanical switch, such as a reed switch. In an embodiment, the switch  87  is a solid-state, magnetically actuated switch circuit. In an embodiment, the switch  87  is a micro-electro-mechanical system (MEMS). In an embodiment, the solid state switch  87  is a MAGFET. In an embodiment, the solid state switch  87  is a giant magneto-resistivity (GMR) sensor. In an embodiment, the switch  87  is normally open. The other pole of switch  87  is connected to the second terminal of capacitor  86  and to the signal processing circuit  34 . Switch  87  automatically closes when in the presence of a magnetic field. When the switch  87  is closed, input  81  provides a signal that is filtered by capacitor  86 . The filtered signal is provided to the signal processing circuit  34 . The capacitor  86  acts as a filter for the signal sent by the input  81  to the signal processing circuit  34 . Thus, switch  87  automatically activates input  81  and filter  86  when in the presence of a magnetic (wireless) field or signal. When the magnetic field is removed, then the switch automatically opens and electrically opens the input  81  and filter  86  from the signal processing circuit  34 . 
       FIG. 6  shows a further hearing aid  90 . Hearing aid  90  includes at least one input  81  having nodes  84 ,  85  connected to signal processing circuit  34 , which is connected to output speaker  36 . Node  85  is connected to first pole of switch  87 . Node  84  is connected to a first terminal of filter  86 . The second pole of switch  87  is connected to the second terminal of filter  86 . In an embodiment, the switch  87  is normally open. Accordingly, in the default state of hearing aid  90 , the signal sensed by input  81  is sent directly to the signal processing circuit  34 . In the switch active state of hearing aid  90 , the switch  87  is closed and the signal sent from the input  81  is filtered by filter  86  prior to the signal being received by the signal processing circuit  34 . The  FIG. 6  embodiment provides automatic signal filtering when the switch  87 , and hence the hearing aid  90 , is in the presence of a magnetic field. 
       FIG. 7  shows a further hearing aid  100  that includes input  81 , signal processing circuit  34  and output system  36 . The input  81  is connected to a plurality of filtering circuits  101   1 ,  101   2 ,  101   3 . Thus, signal generated by the input  81  is applied to each of the filters  101 . Each of the filtering circuits  101  provides a different filter effect. For example, the first filter is a low-pass filter. The second filter is a high-pass filter. The third filter is a low-pass filter. In an embodiment, at least one of filtering circuits  101   1 ,  101   2 ,  101   3  includes an active filter. Each of the filters  101  are connected to a switching circuit  102 . In an embodiment, the switching circuit  102  is a magnetically actuatable switch as described herein. The switching circuit  102  determines which of the filters  101  provides a filtered signal to the signal processing circuit  34 . The processing circuit  34  sends a signal to the output system  36  for broadcasting into the ear of the hearing aid wearer. The switching circuit  102  in the absence of a magnetic field electrically connects the first filter  101   1  to the signal processing circuit  34  and electrically opens the second filter  101   2  and third filter  101   3 . The switching circuit  102  in the presence of a magnetic field opens the first filter  101   1  and electrically connects at least one of the second filter  101   2  and third filter  101   3  to the signal processing circuit  34 . In an embodiment, the second and third filters provide a band-pass filter with both being activated by the switching circuit  102 . While the embodiment of  FIG. 7  shows the switching circuit  102  positioned between the filters and the hearing aid signal processing circuit  34 , the switching circuit  102  is positioned between the input  81  and the filtering circuits  101   1 ,  101   2 ,  101   3  in an embodiment of the present invention. In this embodiment, the switching circuit  102  only supplies the input signal from input  81  to the selected filtering circuit(s)  101   1 ,  101   2 ,  101   3 . 
       FIG. 8  shows an embodiment of the present invention including a hearing aid  110  having a magnetic field sensor  115 . The magnetic field sensor  115  is connected to a selection circuit  118 . The selection circuit  118  controls operation of at least one of a programming circuit  120 , a signal processing circuit  122 , output processing circuit  124  and an input circuit  126 . The sensor  115  senses a magnetic field or signal and outputs a signal to the selection circuit  118 , which controls at least one of circuits  120 ,  122 ,  124  and  126  based on the signal produced by the magnetic field sensor  115 . The signal output by sensor  115  includes an amplitude level that may control which of the circuits that is selected by the selection circuit  118 . That is, a magnetic field having a first strength as sensed by sensor  115  controls the input  126 . A magnetic field having a second strength as sensed by sensor  115  controls the programming circuit  120 . The magnetic field as sensed by sensor  115  then varies from the second strength to produce a digital programming signal. In an embodiment, the signal output by sensor  115  includes digital data that is interpreted by the selection circuit to select at least one of the subsequent circuits. The selection circuit  118  further provides a signal to the at least one of the subsequent circuits. The signal controls operation of the at least one circuit. 
     In an embodiment, the signal from the selection circuit  118  controls operation of a programming circuit  120 . Programming circuit  120  provides hearing aid programmable settings to the signal processing circuit  122 . In an embodiment, the magnetic sensor  115  and the selection circuit  118  produce a digital programming signal that is received by the programming circuit  120 . Hearing aid  110  is programmed to an individual&#39;s specific hearing assistance needs by providing programmable settings or parameters to the hearing aid. Programmable settings or parameters in hearing aids include, but are not limited to, at least one of stored program selection, frequency response, volume, gain, filtering, limiting, and attenuation. The programming circuit  120  programs the programmable parameters for the signal processing circuit  122  of the hearing aid  110  in response to the programming signal received from the magnetic sensor  115  and sent to the programming circuit  120  through selection circuit  118 . 
     In an embodiment, the signal from selection circuit  118  directly controls operation of the signal processing circuit  122 . The signal received by the processing circuit  122  controls at least one of the programmable parameters. Thus, while the signal is sent by the magnetic sensor  115  and the selection circuit  118 , the programmable parameter of the signal processing circuit  122  is altered from its programmed setting based on the signal sensed by the magnetic field sensor  115  and sent to the signal processing circuit  122  by the selection circuit  118 . It will be appreciated that the programmed setting is a factory default setting or a setting programmed for an individual. In an embodiment, the alteration of the hearing aid settings occurs only while the magnetic sensor  115  senses the magnetic field. The hearing aid  110  returns to its programmed settings after the magnetic sensor  115  no longer senses the magnetic field. 
     In an embodiment, the signal from selection circuit  118  directly controls operation of the output processing circuit  124 . The output processing circuit  124  receives the processed signal, which represents a conditioned audio signal to be broadcast into a hearing aid wearer&#39;s ear, from the signal processing circuit  122  and outputs a signal to the output  128 . The output  128  includes a speaker that broadcasts an audio signal into the user&#39;s ear. Output processing circuit  124  includes filters for limiting the frequency range of the signal broadcast from the output  128 . The output processing circuit  124  further includes an amplifier for amplifying the signal between the signal processing circuit  122  and the output. Amplifying the signal at the output allows signal processing to be performed at a lower power. The selection circuit  118  sends a control signal to the output processing circuit  124  to control the operation of at least one of the amplifying or the filtering of the output processing circuit  124 . In an embodiment, the output processing circuit  124  returns to its programmed state after the magnetic sensor  115  no longer senses a magnetic field. 
     In an embodiment, the signal from the selection circuit  118  controls operation of the input circuit  126  to control which input is used. For example, the input circuit  126  includes a plurality of inputs, e.g., an audio microphone and a magnetic field input or includes two audio inputs. In an embodiment, the input circuit  126  includes an omnidirectional microphone and a directional microphone. The signal from the selection circuit  118  controls which of these inputs of the input circuit  126  is selected. The selected input sends a sensed input signal, which represents an audio signal to be presented to the hearing aid wearer, to the signal processing circuit  122 . In a further example, the input circuit  126  includes a filter circuit that is activated and/or selected by the signal produced by the selection circuit  118 . 
       FIG. 9  shows an embodiment of the magnetic sensor  115 . Sensor  115  includes a full bridge  140  that has first node connected to power supply (Vs) and a second node connected ground. The bridge  140  includes third and fourth nodes whereat the sensed signal is output to further hearing aid circuitry. A first variable resistor R 1  is connected between the voltage source and the third node. A second variable resistor R 2  is connected between ground and the fourth node. The first and second variable resistors R 1  and R 2  are both variable based on a wireless signal. In an embodiment, the wireless signal includes a magnetic field signal. A first fixed value resistor R 3  is connected between the voltage source and the fourth node. A second fixed value resistor R 4  is connected between ground and the third node. The bridge  140  senses an electromagnetic field produced by a source  142  and produces a signal that is fed to an amplifier  143 . Both the first and second variable resistors R 1  and R 2  vary in response to the magnetic field produced by magnetic field source  142 . Amplifier  143  amplifies the sensed signal. A low pass filter  144  filters high frequency components from the signal output by the amplifier  143 . A threshold adjust circuit  145 , which is controlled by threshold control circuit  146 , adjusts the level of the signal prior to supplying it to the selection circuit  118 . In an embodiment, the threshold adjust circuit  145  holds the level of the signal below a maximum level. The maximum level is set by the threshold adjust circuit  146 . 
       FIG. 10  shows a further embodiment of magnetic sensor  115 , which includes a half bridge  150 . The half bridge  150  includes two fixed resistors R 5 , R 6  connected in series between a voltage source and the output node. Bridge  150  further includes two variable resistors R 7 , R 8  connected in series between ground and the output node. The two variable resistors R 7 , R 8  sense the electromagnetic field produced by the magnetic field source  142  to produce a corresponding signal at the output node. The amplifier  143 , filter  144 , threshold adjust circuit  145  and selection circuit  118  are similar to the circuits described herein. 
     The magnetic sensor  115 , in either the full bridge  140  or half bridge  150 , includes a wireless signal responsive, solid state device. The solid state sensor  115 , in an embodiment, includes a giant magnetoresistivity (GMR) device, which relies on the changing resistance of materials in the presence of a magnetic field. One such GMR sensor is marketed by NVE Corp. of Eden Prairie, Minn. under part no. AA002-02. In one embodiment of a GMR device, a plurality of layers are formed on a substrate or wafer to form an integrated circuit device. Integrated circuit devices are desirable in hearing aids due to their small size and low power consumption. A first layer has a fixed direction of magnetization. A second layer has a variable direction of magnetization that depends on the magnetic field in which it is immersed. A non-magnetic, conductive layer separates the first and second magnetic layers. When the direction of magnetization of the first and second layers are the same, the resistance across the GMR device layer is low. When the direction of magnetization of the second layer is at an angle with respect to the first layer, then the resistance across in the layers increases. Typically, the maximum resistance is achieved when the direction of magnetization are at an angle of about 180 degrees. Such GMR devices are manufactured using VLSI fabrication techniques. This results in magnetic field sensors having a small size, which is also desirable in hearing aids. In an embodiment, a GMR sensor of the present invention has an area of about 130 mil by 17 mil. It will be appreciated that smaller GMR sensors are desirable for use in hearing aids if they have the required sensitivity and bandwidth. Further, some hearing aids are manufactured on a ceramic substrate that will form a base layer on which a GMR sensor is fabricated. GMR sensors have a low sensitivity and thus must be in a strong magnetic field to sense changes in the magnetic field. Further, magnetic field strength depends on the cube of the distance from the source. Accordingly, when the GMR sensor is used to program a hearing aid, the magnetic field source  142  must be close to the GMR sensor. As a example, a programming coil of the source  142  is positioned about 0.5 cm from the GMR sensor to provide a strong magnetic field to be sensed by the magnetic field sensor  115 . 
     When the GMR sensor is used in the hearing aid circuits described herein, the GMR sensor acts as a switch when it senses a magnetic field having at least a minimum strength. The GMR sensor is adapted to provide various switching functions. The GMR sensor acts as a telecoil switch when it is placed in the DC magnetic field of a telephone handset in a first function. The GMR sensor acts as a filter-selecting switch that electrically activates or electrically removes a filter from the signal processing circuits of a hearing aid in an embodiment. The GMR sensor acts to switch the hearing aid input in an embodiment. For example, the hearing aid switches between acoustic input and magnetic field input. As a further example, the hearing aid switches between omni-directional input and directional input. In an embodiment, the GMR sensor acts to automatically turn the power off when a magnetic field of sufficient strength changes the state, i.e., increases the resistance, of the GMR sensor. 
     The GMR sensor is adapted to be used in a hearing aid to provide a programming signal. The GMR sensor has a bandwidth of at least 1 MHz. Accordingly, the GMR sensor has a high data rate that is used to program the hearing aid during manufacture. The programming signal is a digital signal produced by the state of the GMR sensor when an alternating or changing magnetic field is applied to the GMR sensor. For example, the magnetic field alternates about a threshold field strength. The GMR sensor changes its resistance based on the magnetic field. The hearing aid circuit senses the change in resistance and produces a digital (high or low) signal based on the GMR sensor resistance. In a further embodiment, the GMR sensor is a switch that activates a programming circuit in the hearing aid. The programming circuit in an embodiment receives audio signals that program the hearing aid. In an embodiment, the audio programming signal is broadcast through a telephone network to the hearing aid. Thus, the hearing aid is remotely programmed over a telephone network using audio signals by non-manually switching the hearing aid to a programming mode. In an embodiment, the hearing aid receives a variable magnetic signal that programs the hearing aid. In an embodiment, the telephone handset produces the magnetic signal. The continuous magnetic signal causes the hearing aid to switch on the programming circuit. The magnetic field will remain above a programming threshold. The magnetic field varies above the programming threshold to produce the programming signal that is sensed by the magnetic sensor and programs the hearing aid. In a further embodiment, a hearing aid programmer is the source of the programming signal. 
     The solid state sensor  115 , in an embodiment, is an anisotropic magneto resistivity (AMR) device. An AMR device includes a material that changes its electrical conductivity based on the magnetic field sensed by the device. An example of an AMR device includes a layer of ferrite magnetic material. An example of an AMR device includes a crystalline material layer. In an embodiment, the crystalline layer is an orthorhombic compound. The orthorhombic compound includes RCu2 where R=a rare earth element). Other types of anisotropic materials include anisotropic strontium and anisotropic barium. The AMR device is adapted to act as a hearing aid switch as described herein. That is, the AMR device changes its conductivity based on a sensed magnetic field to switch on or off elements or circuits in the hearing aid. The AMR device, in an embodiment, is adapted to act as a hearing aid programming device as described herein. The AMR device senses the change in the state of the magnetic field to produce a digital programming signal in the hearing aid. 
     The solid state sensor  115 , in an embodiment, is a spin dependent tunneling (SDT) device. Spin dependent tunneling (SDT) structures include an extremely thin insulating layer separating two magnetic layers. The conduction is due to quantum tunneling through the insulator. The size of the tunneling current between the two magnetic layers is modulated by the magnetization directions in the magnetic layers. The conduction path must be perpendicular to the plane of a GMR material layer since there is such a large difference between the conductivity of the tunneling path and that of any path in the plane. Extremely small SDT devices with high resistance are fabricated using photolithography allowing very dense packing of magnetic sensors in small areas. The saturation fields depend upon the composition of the magnetic layers and the method of achieving parallel and antiparallel alignment. Values of a saturation field range from 0.1 to 10 kA/m (1 to 100 Oe) offering the possibility of extremely sensitive magnetic sensors with very high resistance suitable for use with battery powered devices such as hearing aids. The SDT device is adapted to be used as a hearing aid switch as described herein. The SDT device is further adapted to provide a hearing aid programming signals as described herein. 
     Hearing aids are powered by batteries. In an embodiment, the battery provides about 1.25 Volts. A magnetic sensor, e.g., bridges  140  or  150 , sets the resistors at 5K ohms, with the variable resistors R 1 , R 2  or R 7 , R 8  varying from the 5K ohm dependent on the magnetic field. In this embodiment, the magnetic sensor  140  or  150  would continuously draw about 250 μA. It is desirable to limit the power draw from the battery to prolong the battery life. One construction for limiting the power drawn by the sensor  140  or  150  is to pulse the supply voltage Vs.  FIG. 11  shows a pulsed power circuit  180  that receives the 1.25 Volt supply from the hearing aid battery  181 . Pulsed power circuit  180  includes a timer circuit that is biased (using resistors and capacitors) to produce a 40 Hz pulsed signal that has a pulse width of about 2.8 μsec. and a period of about 25.6 μsec for a duty cycle of about 0.109. Such, a pulsed power supply uses only about a tenth of the current that a continuous power supply would require. Thus, with a GMR sensor that continuously draws 250 μA, would only draw about 25 μA to with a pulsed power supply. In the specific embodiment, the current drain on the battery would be about 27 μA to (0.109*250 μA). Accordingly, the power savings of a pulsed power supply versus a continuous power supply is about 89.1%. 
       FIG. 12  shows an embodiment of a GMR sensor circuit  190  that operates as both a hearing aid state changing switch and as a programming circuit. Circuit  190  includes a sensing stage  192 , followed by a high frequency signal stage  193 , which is followed by a bi-state sensing and switch stage  201 . The hearing aid state changing switch is adaptable to provide any of bi-states of the hearing aid, for example, changing inputs, changing filters, turning the hearing aid on or off, etc. The GMR sensor circuit  190  includes a full bridge  192  that receives a source voltage, for example, Vs or the output from the pulse circuit  180 . Vs is, in an embodiment, the battery power. The bridge  192  outputs a signal to both the signal stage  193  and the switch stage  201 . The positive and negative output nodes of the full bridge  192  are respectively connected to the non-inverting and inverting terminals of an amplifier  194  through capacitors  195 ,  196 . The amplifier is part of the signal stage  193 . In an embodiment, the output  197  of the amplifier  194  is a digital signal that is used to program the hearing aid. The hearing aid programming circuit, e.g., programming circuit  120 , receives the digital signal  197  from the amplifier  194 . The signal  197 , in an embodiment, is the audio signal that is inductively sensed by bridge  192  and is used as an input to the hearing aid signal processing circuit. 
     The switching stage  201  includes filters to remove the high frequency component of the signal from the induction sensor. The positive and negative output nodes of the full bridge  192  are each connected to a filter  198 ,  199 . Each filter  198 ,  199  includes a large resistor (1 M ohm) and a large capacitor (1 μf). The filters  198 ,  199  act to block false triggering of the on/off switch component  200  of the circuit  190 . The signals that pass filters  198 ,  199  are fed through a series of amplifiers to determine whether an electromagnetic field is present to switch the state of the hearing aid. An output  205  is the on/off signal from the on/off switch component  200 . The on/off signal is used to select one of two states of the hearing aid. The state of the hearing aid, in an embodiment, is between an audio or electromagnetic field input. In another embodiment, the state of the hearing aid is either an omni-directional input or directional input. In an embodiment, the state of the hearing aid is a filter acting on a signal in the hearing aid or not. In an embodiment, the signal  205  is sent to a level detection circuit  206 . Level detection circuit  206  outputs a digital (high or low) signal  207  based on the level of signal  205 . In this embodiment, signal  207  is the signal used for switching the state of the hearing aid. 
       FIG. 13  shows a saturated core circuit  1300  for a hearing aid. The saturated core circuit  1300  senses a magnetic field and operates a switch or provides a digital programming signal. A pulse circuit  1305  connects the saturated core circuit to the power supply Vs. Pulse circuit  1305  reduces the power consumption of the saturated core circuit  1300  to preserve battery life in the hearing aid. The pulse circuit  1305  in the illustrated embodiment outputs a 1 MHz signal, which is fed to a saturatable core, magnetic field sensing device  1307 . In an embodiment, the device includes a magnetic field sensitive core wrapped by a fine wire. The core in an example is a 3.0×0.3 mm core. In an embodiment, the core is smaller than 3.0×0.3 mm. The smaller the core, the faster it responds to magnet fields and will saturate faster with a less intense magnetic field. An example of a saturated core is a telecoil marketed by Tibbetts Industries, Inc. of Camden, Me. However, the present invention is not limited to the Tibbetts Industries telecoil. In a preferred embodiment of the invention, the saturatable core device  1307  is significantly smaller than a telecoil so that the device will saturate faster in the presence of the magnetic field. The device  1307  changes in A.C. impedance based on the magnetic field surrounding the core. The core has a first impedance in the presence of a strong magnetic field and a second impedance when outside the presence of a magnetic field. A resistor  1308  connects the device  1307  to ground. In an embodiment, the resistor  1308  has a value of 100 KOhms. The node intermediate the device  1307  and resistor  1308  is a sensed signal output that is based on the change in impedance of the device  1307 . Accordingly, the saturable core device  1307  and resistor  1308  act as a half bridge or voltage divider. The electrical signal produced by the magnetic field sensing device  1307  and resistor  1308  is sent through a diode D 1  to rectify the signal. A filter  1309  filters the rectified signal and supplies the filtered signal to an input of a comparator  1310 . The comparator  1310  compares the signal produced by the filter and magnetic field sensor to a reference signal to produce output signal  1312 . In an embodiment, the signal output through the core device  1307  varies +/−40 mV depending on the magnetic field in which the saturable core device  1307  is placed. In an embodiment, it is preferred that the magnetic field is of sufficient strength to move the saturable core device into saturation. While device  1307  is shown as a passive device, in an embodiment of the present invention, device  1307  is a powered device. In an embodiment, the saturatable device  1307  acts a non-manual switch that activates or removes circuits from the hearing aid circuit. For example, the saturatable device  1307  acts to change the input of the hearing aid in an embodiment. In a further embodiment, the saturated core circuit  1300  activates or removes a filter from the hearing aid circuit based on the state of the output  1312 . In a further embodiment, the saturatable core device  1307  is adapted to be a telecoil switch. In a further embodiment, the saturatable core device  1307  is adapted to act as a automatic, non-manual power on/off switch. In a further embodiment, the saturatable core  1307  is a programming signal receiver. 
       FIG. 14  shows a system  1401  including a hearing aid  1405  and a hearing aid storage receptacle  1410 . Receptacle  1410  is cup-like with an open top  1411 , an encircling sidewall  1412  upstanding from a base  1413 . The receptacle  1410  is adapted to receive the hearing aid  1405  and store it adjacent a magnetic field source  1415 . The receptacle base  1413  houses the magnetic field source  1415 . Thus, when the hearing aid  1405  is in the receptacle (shown in solid line in  FIG. 14 ), the hearing aid is in the magnetic field. In an embodiment, the magnetic field experienced by the hearing aid in the receptacle is the near field. When the hearing aid  1405  is out of receptacle (broken line showing in  FIG. 14 ), the hearing aid is out of the magnetic field, i.e., the magnetic field does not have sufficient strength as sensed by the magnetic field sensor of hearing aid  1405  to trigger a state changing signal in the hearing aid  1405 . In an embodiment, the hearing aid  1405  includes a magnetically-actuated switch  1406 . The magnetically-actuated switch  1406  is a normally on (conducting) switch that connects the power supply to the hearing aid circuit. When the hearing aid  1405  is in the receptacle, the magnetically-actuated switch changes to a non-conducting state and the power supply is electrically disconnected from the hearing aid circuit. Thus, hearing aid  1405  is placed in a stand-by mode. The stand-by mode reduces power consumption by the hearing aid. This extends hearing aid battery life. Moreover, this embodiment eliminates the need for the hearing aid wearer to manually turn off the hearing aid after removing it. The wearer merely places the hearing aid  1405  in the storage receptacle  1410  and the hearing aid  1405  turns off or is placed in a stand-by mode. Non-essential power draining circuits are turned off. Non-essential circuits include those that are used for signal processing that are not needed when the hearing aid wearer removes the hearing aid. The stand-by mode is used so that any programmable parameters stored in the hearing aid  1405  are saved in memory by power supplied to the hearing aid memory. The programmable parameters are essential parameters that are stored in the hearing aid and should not be deleted with the power being turned off. The programmed parameters include the volume level. Thus, when the hearing aid  1405  is removed from the receptacle  1410 , the hearing aid is automatically powered by the normally on switch  1406  electrically reconnecting the hearing aid signal processing circuit to the power supply and the hearing aid  1405  returns to the stored volume level without the wearer being forced to manually adjust the volume level of the hearing aid. 
     The hearing aid storage system  1401 , in an embodiment, includes a magnetic field source  1415  that produces a magnetic field that is significantly greater, e.g., at least 3-4 times as great, as the constant magnetic field and/or the varying magnetic field of a telephone handset. This allows the hearing aid  1405  to include both the automatic switch  40  that alternates inputs based on a magnetic field of a first threshold and the automatic power-off switch  1406  that turns off the hearing aid based on a magnetic field of a higher threshold. Thus, hearing aid  1405  includes automatically switching between inputs, filters, settings, etc. as described herein and automatically powering down to preserve battery power when the hearing aid is in the storage receptacle  1410 . 
     In another embodiment of the present invention, the hearing aid  1405  further includes a rechargeable power supply  1407  and a magnetically actuated switching circuit  1406  as described herein. The rechargeable power supply  1407  includes at least one of a rechargeable battery. In an embodiment, rechargeable power supply  1407  includes a capacitor. In an embodiment, a power induction receiver is connected to the rechargeable power supply  1407  through the switching circuit  1406 . The receptacle  1410  includes a power induction transmitter  1417  and magnetic field source  1415 . When the hearing aid  1405  is positioned in the receptacle  1410 , the magnetic switch  1406  turns on a power induction receiver of the rechargeable power supply  1407 . The power induction receiver receives a power signal from the power induction transmitter  1417  to charge the power supply  1407 . Thus, whenever the hearing aid  1405  is stored in the receptacle  1410 , the hearing aid power supply  1407  is recharged. In an embodiment, the magnetically actuated switch  1406  electrically disconnects the hearing aid circuit from the hearing aid power supply  1407  and activates the power induction receiver to charge the hearing aid power supply. As a result, the hearing aid power supply  1407  is recharged when the hearing aid is not in use by the wearer. 
     In a further embodiment, the system  1401  includes a cleaning source  1430  connected to the storage receptacle  1410 . The cleaning source  1430  supplies sonic or ultrasonic cleaning waves inside the receptacle  1411 . The waves are adapted to clean the hearing aid  1405 . Accordingly, the hearing aid  1405  is automatically cleaned when placed in the receptacle  1411 . 
       FIG. 15  shows a further embodiment of the hearing aid switch  1406  that includes an indicator circuit  1450 . Indicator circuit  1450  is adapted to produce an indicator signal to the hearing aid user. In an embodiment, the indicator circuit  1450  is connected to a magnetic field sensor, e.g. sensor  115 ,  190  or  1300 . The indicator circuit provides an indication signal that indicates that the magnetic field sensor  190  or  1300  is sensing the magnetic field. In an embodiment, the indicator circuit indicates that the hearing aid has been disconnected from the power supply. In an embodiment, the indicator circuit indicates that the hearing aid power supply is being recharged by the recharging circuit  1417 . Indicator circuit  1450  includes a comparator  1455  that receives the output signal from the magnetic field sensor circuit  190  or  1300  and compares the received output signal to a threshold value and based on the comparison sends a signal to an indicator  1460  that produces the indicator signal. The indicator signal is a visual signal produced by a low power LED. 
       FIG. 16  shows a hearing aid switch circuit  1600 . Circuit  1600  switches the power from one input to another input. In an embodiment, one input is an induction input and the other input is an audio input. In an embodiment, circuit  1600  exclusively powers one of the inputs. Circuit  1600  includes a power supply  1601  connected to a resistor  1603  at node  1604 . Hence, node  1604  is at a high, non-groung potential. In an embodiment, the power supply is a hearing aid battery power supply. In an embodiment, the power supply is in the range of 1.5 to 0.9 volts. In an embodiment, the resistor  1603  is a 100 KOhm. The resistor  1603  is connected to a non-manual switch  1605  that is connected to ground. Switch  1605 , in an embodiment, is a magnetically actuatable switch as described herein. An input to first invertor  1607  is connected to node  1604 . The output of invertor  1607  is connected to the input of a first hearing aid input  1609  and an input of a second invertor  1611 . The output of the second invertor  1611  is connected to a second hearing aid input  1613 . In an embodiment, first and second invertors  1607  and  1611  are Fairchild ULP-A NC7SV04 invertors. The invertors have an input voltage range from 0.9V to 3.6V. 
     The circuit  1600  has two states. In the first state, which is illustrated, the switch  1605  is open. The node  1604  is at a high voltage. Invertor  1607  outputs a low signal, which is supplied to both the first input  1609  and the second invertor  1611 . The first input  1609  is off when it receives a low signal. The second invertor  1611  outputs a high, on signal to the second input  1613 . Accordingly, in the open switch state of circuit  1600 , the first input  1609  is off and the second input  1613  is on. When in the presence of a magnetic field, switch  1605  closes. Node  1604  is connected to ground and, hence, is at a low potential. Invertor  1607  outputs a high, on signal to the first input  1609  and second invertor  1611 . The first input  1609  is on, i.e., powered. The second invertor  1611  outputs a low, off signal to second input  1613 . Accordingly, in the closed switch state of circuit  1600 , the first input  1609  is on and the second input  1613  is off. In an embodiment, the first hearing aid input  1609  is an induction input and the second hearing aid input  1613  is an audio input. Thus, in the switch open state, the second, audio input  1613  is on or powered and the first, induction input  1609  is off or unpowered. In the switch closed state, the first, induction input  1609  is on or powered and the second, audio input  1613  is off. The circuit  1600  is used as an automatic, induction telephone signal input circuit. 
       FIG. 17  shows a hearing aid switch circuit  1700 . Circuit  1700  is similar to circuit  1600 , like elements are designated with the same two least significant digits and the two most significant digit refer to the FIG. on which they appear. In circuit  1700 , the switch  1705  is connected to the voltage supply  1701 . Resistor  1703  is connected between node  1704  and ground. The input of first invertor  1707  is connected to node  1704 . The output of first invertor  1707  is connected to the first input  1709  and the input of the second invertor  1711 . The output of the second invertor  1711  is connected to the second input  1713 . 
     The circuit  1700  has two states. In the first state, which is illustrated, the switch  1705  is open. The node  1704  is grounded by resistor  1703  and is at a low potential. Invertor  1707  outputs a high signal, which is supplied to both the first input  1709  and the second invertor  1711 . The first input  1709  is on when it receives a high signal. The second invertor  1711  outputs a low, off signal to the second input  1713 . Accordingly, in the open switch state of circuit  1700 , the first input  1709  is on and the second input  1713  is off. When in the presence of a magnetic field, switch  1705  closes. Node  1704  is connected to the voltage supply through closed switch  1705  and, hence, is at a high potential. Invertor  1707  outputs a low, off signal to the first input  1709  and second invertor  1711 . The first input  1709  is off, i.e., unpowered. The second invertor  1711  outputs a high, on signal to second input  1713 . Accordingly, in the closed switch state of circuit  1700 , the first input  1709  is off and the second input  1713  is on. In an embodiment, the first hearing aid input  1709  is an audio input and the second hearing aid input  1713  is an induction input. Thus, in the switch open state, the first, audio input  1709  is on or powered and the second, induction input  1713  is off or unpowered. In the switch closed state, the first, audio input  1709  is off and the second, induction input  1713  is on or powered. The circuit  1700  is used as an automatic, induction telephone signal input circuit. Further, circuit  1700  does not continually incur the loss associated with resistor  1703 . The default state of the circuit  1700  is with the resistor  1703  grounded and no power drain occurs across resistor  1703 . In circuit  1600 , there is a continuous power loss associated with resistor  1603 . Power conservation and judicious use of the battery power in a hearing aid is a significant design characteristic. 
       FIG. 18  shows a hearing aid switch circuit  1800 . Circuit  1800  includes a supply voltage  1801  connected to an induction, first hearing aid input  1809  and a non-manual switch  1805 . Switch  1805 , in an embodiment, is a magnetic field actuatable switch as described herein. A resistor  1803  connects a node  1804  to ground. Switch  1805  is connected to node  1804 . Invertor  1807  is connected to node  1810 . Both first input  1809  and an audio, second hearing aid input  1813  are connected to node  1810 . Second input  1813  is connected to ground. Circuit  1800  has two states. In a first, switch open state node  1804  is connected to ground through resistor  1803 . The invertor  1807  outputs a high signal to node  1810 . The high signal turns on or powers the second input  1813 . The high signal at node  1810  is a high enough voltage to hold the potential across the first input  1809  to be essential zero. In an embodiment, the high signal output by invertor  1807  is essentially equal to the supply voltage  1801 . Thus, the first input  1809  is off. In a second, switch closed state, node  1804  is at a high potential. Invertor  1807  outputs a low signal. In an embodiment, the low signal is essentially equal to ground. The potential across the first input  1809  is the difference between the supply voltage and the low signal. The potential across the first input  1809  is enough to turn on the first input. The low signal is low enough so that there is no potential across the second input  1813 . Thus, the first input  1809  is on and the second input  1813  is off in the closed switch state of circuit  1800 . 
     While the above embodiments described in conjunction with  FIGS. 16-18  include invertors, it will be recognized that the other logic circuit elements could be used. The logic circuit elements include NAND, NOR, AND and OR gates. The use of logic elements, invertors and other logic gates, is a preferred approach as these elements use less power than the transistor switch circuit as shown in  FIG. 3 . 
     The above embodiments described in conjunction with  FIGS. 16-18  include switching between hearing aid inputs by selectively powering the inputs based on the state of a switch. It will be recognized that the switching circuits are adaptable to the other switching applications described herein. For example, the switching circuits  1600 ,  1700 , or  1800  switch between an omni-directional input and a directional input. 
       FIG. 19  shows a hearing aid switch circuit  1900 . Circuit  1900  is similar to circuit  1600  described above with like elements being identified by reference numerals having the same two least significant digits and the two larger value digits being changed from 16 to 19. For example, the supply voltage is designated as  1601  in  FIGS. 16 and 1901  in  FIG. 19 . Switching circuit  1900  includes an electrical connection from the output of invertor  1907  to the signal processor  1922 . Consequently, invertor  1907  outputs a low signal to first input  1909 , second invertor  1911  and signal processor  1922  with the magnetic field sensing switch  1905  being open. Invertor  1907  outputs a high signal to first input  1909 , second invertor  1911  and signal processor  1922  with the magnetic field sensing switch  1905  being closed. Thus, the signal processor  1922  receives a hearing aid state signal from the invertor  1907 . In an embodiment, when the state signal is low, then the signal processor  1907  is adapted to optimize the hearing aid signal processing for a second (microphone) input from second input (microphone)  1913 . Second input (microphone)  1913  is in an active state as it has received a high or on signal from second invertor  1911 . The signal processing circuit  1922 , in an embodiment, optimizes the signal processing by selecting stored parameters, which are optimized for second input signal processing, from a memory. In an embodiment, the memory is an integrated circuit memory that is part of the signal processor  1922 . When the state signal is high, then the signal processor  1922  is adapted to optimize the hearing aid signal processing for a first input from first input (telecoil induction)  1909 . First input  1909  is in an active state as it has received a high or on signal from first invertor  1907 . The signal processing circuit  1922 , in an embodiment, optimizes the signal processing by selecting stored parameters, which are optimized for first input (induction) signal processing, from the memory. Other stored parameters in the memory of signal processor  1922  include automatic gain control, frequency response, and noise reduction for respective embodiments of the present disclosure. 
       FIG. 20  shows a hearing aid switch circuit  2000 . Circuit  2000  is similar to circuit  1700  described above with like elements being identified by reference numerals having the same two least significant digits and the two larger value digits being changed from 17 to 20. For example, the supply voltage is designated as  1701  in  FIGS. 17 and 2001  in  FIG. 20 . Switching circuit  2000  includes an electrical connection from the output of first invertor  2007  to the signal processor  2022 . Consequently, invertor  2007  outputs a high signal to first input  2009 , second invertor  2011  and signal processor  2022  with the magnetic field sensing switch  2005  being open. Invertor  2007  outputs a low signal to first input  2009 , second invertor  2011  and signal processor  2022  with the magnetic field sensing switch  2005  being closed. Thus, signal processor  2022  receives a hearing aid state signal from the invertor  2007 . In an embodiment, when the state signal is high, then the signal processor  2022  is adapted to optimize the hearing aid signal processing for a first input signal from first input (microphone)  2009 . First input  2009  is in an active state as it has received a high or on signal from first invertor  2007 . The signal processing circuit  2022 , in an embodiment, optimizes the signal processing by selecting stored parameters, which are optimized for microphone signal processing, from a memory. In an embodiment, the memory is an integrated circuit memory that is part of the signal processor  2022 . When the state signal is low or off, then the signal processor  2022  is adapted to optimize the hearing aid signal processing for a second input signal from second input (telecoil)  2013 . Second input  2013  is in an active state as it has received a high or on signal from second invertor  2011 . The signal processing circuit  2022 , in an embodiment, optimizes the signal processing by selecting stored parameters, which are optimized for second signal (induction) processing, from the memory. Other stored parameters in the memory of signal processor  2022  include automatic gain control, frequency response, and noise reduction for respective embodiments of the present disclosure. 
       FIG. 21  shows a hearing aid switch circuit  2100 . Circuit  2100  includes elements that are substantially similar to elements described above. Like elements are identified by reference numerals having the same two least significant digits and the two larger value digits being changed  21 . For example, the supply voltage is designated as  1601  in  FIG. 16 ,  1701  in  FIGS. 17 and 2101  in  FIG. 21 . Switching circuit  2100  includes a selection circuit that selects signal processing parameters. Selection circuit includes a logic gate  2107 . In the illustrated embodiment, the logic gate  2107  is a NAND gate. A first input of the NAND gate  2107  is connected to the power source  2101 . Thus, this input to the NAND gate is always high. A second input of the NAND gate  2107  is connected to the power source  2201  through a resistor and to a first terminal of magnetic field sensing switch  2105 . Consequently, the state of the switch  2105  determines the output of the NAND gate  2107  during operation of the hearing aid switch  2100 . Operation of hearing aid switch  2100  is defined as when the switch is powered. During the off or non-operational state of the hearing aid switch circuit  2100 , the supply voltage  2101  is turned off and the NAND gate  2107  will always produce a low output to conserve power, which is a consideration in designing hearing aid circuits. The switch  2105  is normally open. Thus, both inputs to the NAND gate  2107  are high and its output signal is high. The output of NAND gate  2107  is connected to signal processor  2122 . Signal processor  2122  includes a switch that upon the change of state of the NAND gate output signal changes a parameter setting in signal processor  2122 . In an embodiment, when the magnetic field sensing switch  2105  senses a magnetic field, switch  2105  closes. The second input to NAND gate  2107  goes low and NAND gate output goes low. This triggers the switch of signal processor  2122  to change parameter settings. In an embodiment, signal processor only changes its parameter settings when the signal from NAND gate  2107  shifts from high to low. In an embodiment, the parameter settings include parameters stored in a memory of signal processor  2122 . In an embodiment, a first parameter setting is adapted to process input from first input  2109 . A second parameter setting is adapted to process input from second input  2113 . In an embodiment, the first parameter setting is selected with the output signal from NAND gate  2107  being high. The second parameter setting is selected with the output signal from NAND gate  2107  being low. Accordingly, the switching circuit  2100  can select parameters that correspond to the type of input, e.g., microphone or induction inputs or directional and omni-directional inputs. The hearing aid thus more accurately produces sound for the hearing aid wearer. In an embodiment, the switch in signal processor  2122  is adapted to progress from one set of stored parameters to the next each time the signal from NAND gate  2107  goes low. 
       FIG. 22  shows a hearing aid switch circuit  2200 . Circuit  2200  includes elements that are substantially similar to elements described above. Like elements are identified by reference numerals having the same two least significant digits and the two larger value digits being changed  22 . For example, the supply voltage is designated as  2101  in  FIG. 21  is  2201  in  FIG. 22 . Switching circuit  2200  includes a selection circuit that is adapted to select parameters for signal processing. The selection circuit includes a logic gate  2207  having its output connected to signal processor  2222 . In the illustrated embodiment, the logic gate  2207  is a NAND gate. A first input of the NAND gate  2207  is connected to the power source  2201 . Thus, this input to the NAND gate is always high. A second input of the NAND gate  2207  is connected to the power source  2201  through a magnetic field sensing switch  2105 . The second input of NAND gate  2207  is also connected to ground through a resistor R. Consequently, the state of the switch  2205  determines the output of the NAND gate  2207  during operation of the hearing aid switch  2200 . Operation of hearing aid switch  2200  is defined as when the switch is powered. During the off or non-operational state of the hearing aid switch circuit  2200 , the supply voltage  2201  is turned off and the NAND gate  2207  will always produce a low output to conserve power, which is a consideration in designing hearing aid circuits. Switch  2205  is normally open. Thus, the first input to the NAND gate  2207  is high and the second input to NAND gate  2207  is low. Thus, the NAND gate output signal is low. Signal processor  2222  includes a switch that upon the change of state of the NAND gate output signal changes a parameter setting in signal processor  2222 . In an embodiment, when the magnetic field sensing switch  2205  senses a magnetic field, switch  2205  closes. The second input to NAND gate  2207  goes high and NAND gate output goes high. This triggers the switch of signal processor  2222  to change parameter settings. In an embodiment, signal processor only changes its parameter settings when the signal from NAND gate  2107  shifts from low to high. In an embodiment, the parameter settings include parameters stored in a memory of signal processor  2222 . In an embodiment, a first parameter setting is adapted to process input from first input  2209 . A second parameter setting is adapted to process input from second input  2213 . In an embodiment, the first parameter setting is selected with the output signal from NAND gate  2207  being low. The second parameter setting is selected with the output signal from NAND gate  2207  being high. Accordingly, the switching circuit  2200  can select parameters that correspond to the type of input, e.g., microphone or induction inputs. The hearing aid thus more accurately produces sound for the hearing aid wearer. 
     It will be appreciated that the selection of parameters for specific inputs can be combined with the  FIGS. 2-18  embodiments. For example, the magnetic field sensor changing state not only switches the input but also generates a signal, for example, through logic circuit elements, that triggers the signal processing circuit to change its operational parameters to match the type of input. 
     Possible applications of the technology include, but are not limited to, hearing aids. Various types of magnetic field sensors are described herein for use in hearing aids. One type is a mechanical reed switch. Another type is a solid state magnetic responsive sensor. Another type is a MEMS switch. Another type is a GMR sensor. Another type is a core saturation circuit. Another type is anisotropic magneto resistive circuit. Another type is magnetic field effect transistor. It is desirable to incorporate solid state devices into hearing aids as solid state devices typically are smaller, consume less power, produce less heat then discrete components. Further the solid state switching devices can sense and react to a varying magnetic field at a sufficient speed so that the magnetic field is used for supplying programming signals to the hearing aid. 
     Those skilled in the art will readily recognize how to realize different embodiments using the novel features of the present invention. Several other embodiments, applications and realizations are possible without departing from the present invention. Consequently, the embodiment described herein is not intended in an exclusive or limiting sense, and that scope of the invention is as claimed in the following claims and their equivalents.