Classifications

• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B3/00—Apparatus for testing the eyes; Instruments for examining the eyes
  • A61B3/0016—Operational features thereof
  • A61B3/0025—Operational features thereof characterised by electronic signal processing, e.g. eye models
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B3/00—Apparatus for testing the eyes; Instruments for examining the eyes
  • A61B3/10—Objective types, i.e. instruments for examining the eyes independent of the patients' perceptions or reactions
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B3/00—Apparatus for testing the eyes; Instruments for examining the eyes
  • A61B3/0016—Operational features thereof
  • A61B3/0041—Operational features thereof characterised by display arrangements
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B3/00—Apparatus for testing the eyes; Instruments for examining the eyes
  • A61B3/10—Objective types, i.e. instruments for examining the eyes independent of the patients' perceptions or reactions
  • A61B3/16—Objective types, i.e. instruments for examining the eyes independent of the patients' perceptions or reactions for measuring intraocular pressure, e.g. tonometers
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/0002—Remote monitoring of patients using telemetry, e.g. transmission of vital signals via a communication network
  • A61B5/0015—Remote monitoring of patients using telemetry, e.g. transmission of vital signals via a communication network characterised by features of the telemetry system
  • A61B5/0024—Remote monitoring of patients using telemetry, e.g. transmission of vital signals via a communication network characterised by features of the telemetry system for multiple sensor units attached to the patient, e.g. using a body or personal area network
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/02—Detecting, measuring or recording for evaluating the cardiovascular system, e.g. pulse, heart rate, blood pressure or blood flow
  • A61B5/021—Measuring pressure in heart or blood vessels
  • A61B5/022—Measuring pressure in heart or blood vessels by applying pressure to close blood vessels, e.g. against the skin; Ophthalmodynamometers
  • A61B5/02216—Ophthalmodynamometers
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/02—Detecting, measuring or recording for evaluating the cardiovascular system, e.g. pulse, heart rate, blood pressure or blood flow
  • A61B5/021—Measuring pressure in heart or blood vessels
  • A61B5/022—Measuring pressure in heart or blood vessels by applying pressure to close blood vessels, e.g. against the skin; Ophthalmodynamometers
  • A61B5/02225—Measuring pressure in heart or blood vessels by applying pressure to close blood vessels, e.g. against the skin; Ophthalmodynamometers using the oscillometric method
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/02—Detecting, measuring or recording for evaluating the cardiovascular system, e.g. pulse, heart rate, blood pressure or blood flow
  • A61B5/026—Measuring blood flow
  • A61B5/0295—Measuring blood flow using plethysmography, i.e. measuring the variations in the volume of a body part as modified by the circulation of blood therethrough, e.g. impedance plethysmography
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/72—Signal processing specially adapted for physiological signals or for diagnostic purposes
  • A61B5/7225—Details of analogue processing, e.g. isolation amplifier, gain or sensitivity adjustment, filtering, baseline or drift compensation
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/02—Detecting, measuring or recording for evaluating the cardiovascular system, e.g. pulse, heart rate, blood pressure or blood flow
  • A61B5/021—Measuring pressure in heart or blood vessels
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/40—Detecting, measuring or recording for evaluating the nervous system
  • A61B5/4076—Diagnosing or monitoring particular conditions of the nervous system
  • A61B5/4088—Diagnosing of monitoring cognitive diseases, e.g. Alzheimer, prion diseases or dementia

Definitions

• This disclosure relates to ophthalmology, and in particular, to the analysis of ocular blood flow and other workings of the human body. More specifically, this disclosure relates to a novel system that produces accurate, stable, and repeatable low distortion measurements of intraocular pressure (IOP) as a function of time from which reliable data about ocular blood flow (OBF) can be derived.
• IOP intraocular pressure
• OBF ocular blood flow
• a health care provider can assess the ocular blood flow data in conjunction with appropriate clinical correlation, such as additional tests, observations, and historical patient information, to detect abnormalities and diseases in the eye and other parts of the body. For example, a health care provider can detect glaucoma, carotid occlusive disease, and cerebral vascular disease. Also, a health care provider can detect changes in ocular blood flow and can provide appropriate therapeutic interventions such as pharmaceuticals and surgery.
• This disclosure relates to a method and apparatus for use in accurately measuring and recording minute and rapid fluid pressure fluctuations within a closed, fluid filled elastic chamber such as the eye.
• the variation in intraocular pressure occurs in response to each cardiac cycle that generates a bolus of blood that enters the eye and later drains from the eye.
• various diagnostic information including but not limited to, intraocular pressure pulse amplitudes, pulsatile ocular blood flow, pulsatile ocular blood volume, and equivalent Gosling pulsatility index may be ascertained.
• a tonometer is a device used by eye care providers to measure the fluid pressure inside the eye, commonly referred to as the intraocular pressure or IOP.
• the most common device for measuring IOP is the applanation tonometer, which directs a controlled force against the cornea to flatten or applanate the cornea. The IOP opposes and balances the applied force. When a predetermined area of the cornea has been applanated, the applied force is considered to be equal to the IOP and can be recorded as such.
• the alleged gold standard in applanation tonometers is the Goldman
• the Goldman tonometer and other applanation tonometers are limited in that they only supply a static IOP result averaged over time, whereas the device disclosed and claimed in this document reveals a time-resolved IOP.
• the prior applanation tonometers also do not provide any indication of ocular blood flow.
• Pneumatic ocular pressure probes measure and record variations in intraocular pressure over time.
• the literature gives a full description and theoretical analysis of the operating principles of the pneumatic ocular pressure probe (Langham, 1968) and an analysis of its ability to accurately and rapidly measure and record variations in intraocular pressure over time (Walker and Langham, 1975, Walker et al. 1975, Walker & Litovitz 1972).
• Ocular pressure is a reflection of blood flowing into the eye.
• Ocular blood flowing into the eye comes from the internal carotid artery, which arises directly from the heart.
• the ophthalmic artery is the first branch off the carotid artery in the cranium.
• the second branches from the carotid artery become the major arteries feeding the brain.
• Perturbations such as internal carotid artery stenosis can be seen in the measurement of ocular blood flow entering the eye via the internal carotid and ophthalmic arteries (Langham, 2009). Ocular blood flow within the eye arises from the ophthalmic artery.
• the ophthalmic artery gives off nine posterior ciliary arteries to supply the optic nerve and a vast rich plexus of blood vessels comprising the choroid, which nourishes the outer layers of the retina and pigment layer of the eye.
• the ophthalmic artery also gives rise to a single central retinal artery whose branches nourish the inner retina of the eye.
• OCT Ocular Computerized Tomography
• SLO Scanning Laser Ophthalmoscope
• Pulsatile ocular blood flow is due to the bolus of blood entering the eye with each beat of the heart. Because 90% of the ocular blood flow is derived from the choroidal circulation, the POBF is generally accepted as equivalent to the choroidal circulation (Zion, 2007).
• choroidal blood flow is difficult in optical based technologies because the pigmented layer of the eye prevents visualization of the choroid except in the absolute center of the retina or the macula.
• This visualized area comprises only about 5.5 square mm of the entire area of the retina and choroid, which measures about 1094 square mm in total (Schmetter, 2012 and Harris 2010).
• Newer optical based technologies such as swept source ocular computerized tomography and ocular computerized tomographic angiography (Jia, 2015) are now in development stages. These have the potential to measure larger areas of the choroid but are expensive to produce and develop.
• all optical based technologies are dependent on clear ocular media from the front of the eye or the cornea through the lens to the back of the eye where the retina and the optic nerve transmit visual information to the brain.
• common conditions such as cataracts may alter the measurement of ocular blood flow within the eye using optical based technologies.
• these technologies are not as accurate in many of the elderly patients who may have comorbid eye conditions such as cataracts.
• Measurement of retinal blood flow is primarily accomplished through optical based technologies (Schmetter 2012, Harris, 2010). Most of these techniques focus on the measurement of retinal blood flow (only 10% of ocular blood flow) because they are based on the use of the retinal vessel analyzer which cannot detect choroidal vessels and in most devices only measures a small segment of a retinal vessel, about 50 to 150 um. Statistical quantification techniques will increase the validity of these
• Perturbations of ocular blood flow to and within the eye are described in the scientific literature as a major contributor to the pathogenesis of multiple blinding conditions, for example, diabetic retinopathy, glaucoma, and age related macular degeneration. Others include ischemic optic neuropathy, retinal venous occlusive disease, and retinopathy of prematurity. In addition to ophthalmic conditions, changes in ocular blood flow have also been identified in systemic conditions such as
• Alzheimer's and carotid occlusive disease (Langham, 2009).
• POBF Planar Biharmonic adatoma
• the literature describes POBF values that are generally higher in males than in females in addition to an overlapping range of normal and abnormal values in glaucoma and other diseases (Yang, 1997). POBF is influenced by refractive error, axial length of the eye, central corneal thickness, age, sex, and ethnic origin (Zion, 2007). Because of variations between individuals, the technique is characterized by limited reproducibility (Spraul, 1998; Yang, 1997). These variations would make statistical evaluation of sensitivity and specificity using POBF as an indicator of a particular disease difficult. However, the high level of accuracy, reproducibility and repeatability of POBF measurements in individual patients is key to its clinical value. In this regard, the ability to measure changes in POBF with application of topical and systemic medications and to measure the change in POBF in individual patients over time are valuable indicators to clinicians providing eye care to their patients.
• instrument in accordance with this invention at all times delivers significantly lower pressures relative to IOP.
• a first component added after the pump in all systems was a pneumatic pressure regulator.
• Pressure regulators deliver a constant pressure differential across an entry and exit port. However, they limit the flow in so doing. As it turns out a pressure regulator at best produces a marginally useful envelope of flow over time when set to pressure that is safe for operation of the instrument tip.
• a thin stream of air also yielded a dynamic whereby the probe is hyper sensitive to the environmental variables present at the eye surface, causing difficulty in achieving a measurable condition quickly and maintaining the position for the duration of the measurement.
• Airflow in large enough volume over time is critical to proper operation of electronic measurement of pressure changes resulting from variations in flow.
• the second component used downstream from the pressure regulator was a needle valve arrangement to create a backpressure against the inflow of disturbed air from the regulator outlet. This smooths noise that was still unacceptably high in the airflow but further restricts the useful flow of gas or air needed to produce a rapid and reliable change in pressure in the probe in response to change in IOP over time.
• the inflow of air to the tip was constrained to the point where repeatable measurements taken with tips, even taken from the same manufacturing lot, resulted in different measurements of the same eye under the same conditions due to the minute variations in manufacture being sufficient relative to the small air flow to alter the tip to eye interface dynamics. This same weakness is evident in variations between operations of different probes due to minute wear and manufacturing variations.
• a third type of device was employed to meet this challenge, this being a chamber or plenum in the airstream just prior to the probe. This was in attempt to filter or smooth noise in the airstream. This particular instrument suppresses the measured waveforms while allowing easier and more repeatable measurements.
• Silver DM Geyer O. Pressure-volume relation for the living human eye. Curr Eye i?es 2000;20(2):115-120.
• Schmetterer L Dallinger S, Findl O, Eichler H-G, Wolzt M. A comparison between laser interferometric measurement of fundus pulsation and pneumotonometric measurement of pulsatile ocular blood flow. Eye 2000;14:39-45. Berisha F, Findl 0, Last M, Kiss B, Schmetterer L. A study comparing ocular pressure pulse and ocular fundus pulse in dependence of axial length and ocular volume. Acta Ophthalmoligica 2010;88:766-772.
• Harris A Jonescu-Cuypers CP, Kagemann L, Cuilla TA, Kreigelstein GK. Atlas of Ocular Blood Flow: Vascular Anatomy, Pathophysiology, and Metabolism.
• McLeod DS Grebe R, Bhutto I, Merges C, Baba T, Lutty GA. Relationship between RPE and choriocapillaris in age-related macular degeneration. Invest Ophthalmol Vis Sci 2009;50:4982-4991.
• the inventors of the subject matter of this disclosure have solved the problems of prior ocular blood flow analyzers by developing a novel air delivery system in combination with a pressure probe that can produce accurate, linear, and noise-free readings of intraocular pressure, which permits the production of accurate and repeatable data about ocular blood flow, particularly composite ocular blood flow, and more particularly, composite pulsatile ocular blood flow, which includes both retinal and choroidal components of the pulsatile ocular blood flow.
• the inventors also have developed novel electronic circuitry that quickly acquires and processes raw pressure data from the probe and produces data about IOP and ocular blood flow with an accuracy and precision significantly beyond that which has been available to date.
• the accurate IOP and pulsatile ocular blood flow data produced by the invention increases the practitioner's ability to detect and assess over time abnormalities in the eye and other parts of the body.
• the pump is a powerful brushless DC motor driven device, allowing high-speed operation with relatively smooth outflow of air.
• This type of motor is novel in that it maintains high torque so it does not tend to respond to increased resistance by slowing pump speed, pressure or volume.
• a constant supply of low noise pneumatic fluid at reliable pressure and flow is achieved.
• the air from the new pump arrangement is at a higher flow and pressure than fits within the safe operation envelope already established for the instrument/eye interface. This is solved with the use of a pressure compensated flow regulator device that can be purely mechanically operated.
• the pressure compensated flow regulator is also available as an electronic version, which can also be used.
• a flow regulator attempts to maintain a constant volumetric flow over time versus a constant pressure over time at the cost of flow.
• this device acts (as does a pressure regulator) as a noise filter, reducing the threshold of noise from the pump below the signal sought to be measured.
• the increased flow produced, accompanied by decreased pressure has been demonstrated to operate stably as low as 20 mmHg probe pressure: 10 mmHg IOP, with a linear and proportional ascent over a range of IOP from 10 to 40 mmHg.
• Novel hardware plays a significant role in reducing noise while maintaining enough energy in the fluid flow to permit detection of the IOP signal. A significantly higher margin of safety, accuracy, repeatability, and stability in use and responsiveness are shown by experimental results.
• a method and apparatus is involved for use in measuring and recording rapid and accurate fluid pressure within a closed organ without increasing the pressure within the organ itself.
• the system may measure the following indices, among others: (1) variation in intraocular pressure, (2) intraocular pressure pulse amplitudes, (3) pulsatile ocular blood flow, (4) pulsatile ocular blood volume, and (5) equivalent Gosling pulsatility index.
• a pneumatic probe and tip in the measurement apparatus contains a thin walled tube which initially is in contact with a flexible membrane covering the distal end of the tube and is connected to a pneumatic pump which increases the gauge pressure in the probe.
• Probe pressure measurements may be made at a rate of at least 100 times per second over a period of about 10 to 15 seconds or more for the measurement of indices (1-5) above. The pressure measurements can be made at even higher rates such as 200 times per second or more.
• a pump is connected to the probe with a soft plastic tube communicating with a pressure chamber encased in a probe handle.
• a pressure transducer communicates with the pressure chamber to measure probe pressure.
• the pressure changes measured by the pressure transducer are converted to electronic signals through a signal processor that creates an oscillating waveform representing the variation in intraocular pressure as a function of time.
• contact with the eye is at the ocular surface by means of a probe tip covered by a clean or sterile, single-use per patient, flexible membrane that separates the cornea from the pneumatic pressure inside the probe.
• the membrane transfers the pressure to the cornea.
• the pressure transducer may be integral with the probe housing rather than connected to the pressure chamber by a flexible tube.
• the observed oscillations in intraocular pressure occur at the surface of the eye as a result of accommodating changes in the intraocular blood volume induced by the bolus of arterial blood created by each cardiac cycle.
• the oscillations are recorded by a pressure-sensing device or pressure transducer. Analogue electrical signals
• the pulsatile ocular blood flow is derived from the continuous oscillatory pressure measurements acquired from the probe using predetermined relationships that link probe pressure to intraocular pressure, ocular volume change, and ocular volume flow.
• Figure 1 is block diagram of the main components of an ocular blood flow analyzer constructed in accordance with the invention.
• Figure 1A is a schematic block diagram of the pressure compensated flow controller of Figure 1.
• Figure 2 is a schematic cross sectional diagram of the pressure probe shown in Figure 1.
• Figures 3A, 3B, and 3C are flow charts depicting the operation of one example of the software in the computer shown in Figure 1.
• Figure 4 is a plot of probe pressure versus time obtained as a result of an illustrative first eye examination.
• Figure 5 is an illustrative plot of the derivative of probe pressure of Figure 4 as a function of time.
• Figure 6 is an illustrative plot of intraocular pressure as a function of time derived from the probe pressure data of Figure 4 to which a third order polynomial has been fit.
• Figure 7 is an illustrative plot of a flattened zero-based region of interest derived from the intraocular pressure plot of Figure 6 with the results of a peak finding process indicated on the plot.
• Figure 8 is an illustrative plot of a flattened IOP pulsatile region in an intraocular pressure plot based on the zero-based region of interest in the intraocular pressure plot of Figure 7 with the results of a peak finding process indicated on the plot.
• Figure 9 is an illustrative plot of the change in ocular volume as a function of time based on the IOP data shown in Figure 8.
• Figure 10 is an illustrative plot of the change in ocular volume as a function of time shown in Figure 9 with the results of a peak finding process indicated on the plot.
• Figure 11 is an illustrative plot of the derivative of the change in ocular volume as a function of time of Figure 10 with the results of a peak finding process indicated on the plot.
• Figure 12 is an example of numerically presented illustrative ocular blood flow data produced by a machine in accordance with the invention.
• Figure 13 is a flow chart depicting the operation of another example of the software in the computer shown in Figure 1.
• Figure 14 is a plot of probe pressure obtained as a result of an illustrative second eye examination.
• Figure 15 is an illustrative plot of the derivative of the probe pressure data in Figure 14.
• Figure 16 is an illustrative plot of the intraocular pressure data of Figure 14 to which a third order polynomial has been fit.
• Figure 17 is an illustrative plot of a flattened region of interest in an intraocular pressure plot of Figure 16.
• Figure 18 is an illustrative plot of the change in ocular volume as a function of time based on the IOP data of Figure 17.
• Figure 19 is an illustrative plot of the change in ocular volume as a function of time shown in Figure 18.
• Figure 20 is an illustrative flattened zero-based plot of intraocular pressure as a function of time showing the operation of a peak finding process.
• Figure 21 is a perspective view of a manometer that simulates the pressure behavior of a human eye.
• Figure 22 is a plot of pulsatile eye data obtained from a prior art legacy machine.
• Figure 23 is a plot of pulsatile eye data obtained from a machine in accordance with this invention.
• Figure 24 is a graph illustrating and comparing the non-linearity of a prior art machine and the linearity of a machine in accordance with this invention.
• Figure 1 shows an example of an ocular blood flow analyzer in accordance with the invention.
• the architecture of the OBF analyzer shown in Figure 1 comprises a system of electrical, electronic, and pneumatic components that produce ocular blood flow data that is superior to that produced by past and current devices and systems that purport to provide such data about ocular blood flow (OBF).
• OBF ocular blood flow
• the architecture shown in Figure 1 includes a power supply 100 that delivers a variable and controlled amount of electrical power to a brushless DC motor 102.
• a motor controller 103 interposed between the power supply 100 and the motor 102 determines the amount of electrical power delivered to the motor 102.
• the amount of power delivered to the motor 102 by the motor controller 103 is determined by the magnitude of a control voltage produced by a control signal generator 101.
• the control signal generator 101 may be implemented as a manually set tap on a potentiometer connected to a dc voltage source. The voltage set by the position of the tap on the potentiometer constitutes a desired amount of power to be delivered by the motor controller 103 to the motor 102.
• control signal generator 101 may be a computer that issues a voltage that constitutes a command to the motor controller 103 to deliver a predetermined amount of power to the motor 102.
• the computer may be the computer 122 described below or some other computer that controls the behavior of the motor 102 and the motor controller 103.
• the motor 102 provides a constant controlled output torque to drive a pneumatic pump 104 that produces a supply of air or other pneumatic fluid at a predetermined controlled and relatively pulsation free pneumatic pressure to a fluid supply line 106, which may be a flexible hose or tube.
• a fluid supply line 106 which may be a flexible hose or tube.
• This supply of fluid through the supply line 106 is directed to the input of a pressure compensated flow controller 108.
• the flow controller 108 provides a constant volumetric flow rate from its output regardless of pressure fluctuations.
• the flow controller 108 maintains a pneumatic fluid flow through the device of Figure 1 at about 0-160 cc/sec. at pressures from about 20-60 mmHg.
• the output 106 and the supply line 110 may be flexible hose or tubing that conveys pneumatic fluid from the pump 104 to the flow controller 108 and from the flow controller 108 to the probe 112, respectively.
• the output of the flow controller 108 is directed on pneumatic supply line 110, which also may be a flexible hose or tube, to the input of a pneumatic tonometric probe 112.
• a probe tip 114 on the probe 112 is placed in contact with an eye 115 being measured.
• a pressure detect line 116 from the probe 112 has a series connected pressure detector 118, which produces a raw analog pressure signal representing the level of pressure in the probe 112.
• the pressure detector 118 may be connected to the probe 112 by way of a line 116, illustratively in the form of a flexible hose or conduit. Alternatively, the pressure detector 118 may directly mounted on the probe 112 so that it is in direct communication with the pressure chamber in the probe 112. Contact of the tip 114 of probe 112 with an eye 115 causes minute fluctuations in the raw probe pressure that are processed by the electronics of the OBF analyzer to produce accurate, stable, and repeatable low distortion IOP and OBF readings.
• the pressure compensated flow controller 108 may be any control system that establishes a desired substantially constant fluid flow rate toward the pressure probe 112 and initiates compensatory action to return the flow rate to the desired rate in response to deviations of the flow rate from the desired value, particularly in response to pressure fluctuations that cause deviations from the desired flow rate.
• Flow rate controller 108 may comprise a controllable restriction, such as a flow control valve in series with the pneumatic supply line 106 of the pump 104 and the pneumatic supply line 110 between the flow controller 108 and the probe 112.
• a basic flow control valve consists of a changeable aperture that opens to increase the flow rate or closes to slow the flow rate.
• a needle valve that allows precision control of low fluid flow rates is preferred.
• These valves use an adjustable needle and valve stem to restrict or permit fluid flow.
• the flow rate through the flow control valve such as the aforementioned needle valve, generally is related to the pressure drop across the valve.
• the pressure compensated flow controller senses a change in the pressure drop across the flow control valve that causes a change in the flow rate through the valve.
• the flow rate controller 108 initiates compensatory action to return the pressure drop, and the flow rate, to the desired value.
• a valve in series with the supply line 106 and modulated by changes in pressure drop across the needle valve may perform the compensatory action.
• FIG 1A is a detailed block diagram of the pressure compensated flow controller 108 in Figure 1 that specifically illustrates of the points described above.
• the flow controller 108 comprises an inlet compensation valve 103 in series with the supply line 106 from the pump 104.
• a fluid passage 105 connects the inlet compensation valve 103 to the input of a flow regulation valve 107, which preferably can be the needle valve described above.
• the output of the flow regulation valve 107 is connected to the supply line 110 that directs pneumatic fluid to the input of the probe 112.
• the flow regulation valve 107 creates a pressure drop from the input of the valve 107 to the output of the valve 107 that determines the rate of fluid flow to the probe 112.
• the pressure drop across the valve 107 is maintained at a constant value to keep the fluid flow rate to the probe 112 constant in the face of pressure fluctuations on either side of the valve 107.
• the degree to which the valve 107 is opened determines the magnitude of the pressure drop across the valve 107 and thus the desired fluid flow rate to the probe 112.
• Maintaining this pressure drop constant, and thereby maintaining the flow rate to the probe 112 constant, is accomplished by the operation of a compensation structure 109 connected to the input of the valve 107 by way of a fluid passage 111 and to the output of the valve 107 by way of a fluid passage 113.
• the fluid passage 111 conveys information to the compensation structure 109 about the magnitude of the upstream fluid pressure on the inlet side of the valve 107.
• the fluid passage 113 conveys information to the compensation structure 109 about the magnitude of the downstream fluid pressure on the outlet side of the valve 107.
• the compensation structure 109 thus is responsive to the pressure drop across the valve 107.
• the compensation structure 109 has an output 117 that controls the opening of the input compensation valve 103.
• the compensation structure 109 is responsive to the pressure drop across the valve 107 by virtue of the fluid passages 111 and 113 and the output connection 117 to maintain the pressure drop across the valve 107 constant and thereby maintain the fluid flow rate to the probe 112 constant even though upstream or downstream fluid pressure may change.
• An example of a suitable pressure compensated constant flow rate valve and associated control structure is described in Beswick et al. US Patent 6,314,980.
• the selected pump 104 should produce an output as pulsation-free as possible.
• Completely pulse free airflow does not exist, especially not in the nano pressure levels being used here, but a suitable pump is a small diaphragm pump, which is less noisy than other alternatives in that such a diaphragm pump produces smaller, softened, output pulses.
• a brushless dc motor 102 drives the pump. Newer brushless dc motor technology produces an almost flat torque curve across rpm ranges. In this application, precise airflow is desirable as the pump bogs down less under variable loads. High or low rpm of the motor 102 both produce high torque. So the significance of the pump and motor is in providing a steady baseline pressure input to the flow control device.
• brushless dc motors do not spark like brushed motors. Sparks generated by brushed motors are generally unacceptable in medical devices. Sparks could introduce noise and distortion into signals produced by electronic circuitry like that used in measurement apparatus disclosed here. Sparking could also damage delicate electronic circuitry of this OBF analyzer and could be dangerous to the patient and the operator of the instrument. It, therefore, is advantageous to use the brushless dc motor 102 to drive the pump 104.
• the OBF analyzer of Figure 1 also includes a computer 122 that rapidly performs data acquisition followed by digital signal processing that converts the raw pressure data from the probe 112 into IOP data and OBF data used by the practitioner to assess the health of a patient.
• Computer 122 first converts the analog raw probe pressure signal on line 120 from the pressure detector 118 into a stream of digital samples representing the raw probe pressure as a function of time that populates a data table in the computer 122.
• the analog raw pressure signal produced by the pressure detector 118 may be converted to a digital raw probe pressure signal samples by an analog to digital converter integral with the pressure detector 118.
• the digital samples created from the analog output of the pressure detector 118 populate a data table in the computer 122.
• the computer 122 converts the digital raw probe pressure samples into digital samples representing measured IOP as a function of time.
• the computer 122 then converts the IOP data into OBF data.
• This OBF data may include one or more of the flow rate, flow volume, pulse amplitude, and pulse rate.
• the OBF data may also include pulsatility indices.
• the OBF data may be directed to a secure WiFi network 124 or any other secure network used by hospitals and healthcare institutions that is compliant with healthcare and medical laws and regulations such as the Health Insurance
• the POBF data may be shown to an authorized user on a display connected to a laptop or desktop computer 126.
• the data may be displayed numerically or as waveforms of measurements as a function of time.
• the computer 126 may be a Raspberry Pi running the Linux operating system illustratively programmed with suitable Python or IDL computer code compiled to binary code. The invention is not limited to
• Programmable can be used, such as any suitably programmed Microsoft Windows based personal computer, Apple personal computer, or other computer.
• Figure 2 shows a cross section of an example of the pressure probe 112 in Figure
• the probe 112 comprises a hollow elongated housing 200 defining a generally cylindrical pressure chamber 202.
• the pressure chamber 202 has proximal and distal ends 204 and 206 inside the housing 200.
• An inlet 208 into the housing 200 is adapted to admit air from the pump 104 and the flow rate controller 108 into the proximal end 204 of the pressure chamber 202.
• a port 210 in the housing 200 is adapted to communicate with the pressure sensor 118 that measures the air pressure in the chamber 202.
• a shaft 212 having proximal and distal ends 214 and 216, extends from inside the pressure chamber 202 through the distal end 206 of the housing 200 to the exterior of the probe 112.
• the shaft 212 axially slides with respect to the housing 200 by way of a cylindrical bearing 218 defined between the outer surface of the shaft 212 and the inner surface of an opening in a distal end wall 220 of the housing 200 through which the shaft 212 extends.
• An axially directed bore 222 is formed in the shaft 212.
• the bore 222 has proximal and distal ends 224 and 226, respectively.
• the bore 222 is in communication with the air in the pressure chamber 202 at its proximal end 224.
• a passage in the tip 228 attached to the distal end of the shaft 212 is coaxial with the bore 222 and forms a jet or nozzle 213 that directs air originating from the pump 104 toward the eye 115.
• the tip 228 has a cylindrical venting chamber 230 into which nozzle 213 extends.
• a circular flexible membrane 232 covers the distal end of the nozzle 213 and the open end of the venting chamber 230 thus sealing the nozzle 213 and the venting chamber 230.
• Vents 236 exhaust air from the venting chamber 230 when the pressure from the pump 104 is sufficient to cause the membrane 232 to separate from the distal end of the jet 213.
• the pump 104 directs pressurized pneumatic fluid to the input of the flow controller 108.
• the flow controller 108 outputs pneumatic fluid to the input port 208 of the probe 112 at a substantially constant volumetric flow rate.
• the pneumatic fluid flows into the pressure chamber 202, into the passage 222 in the rod
• the flow controller 108 maintains a substantially constant pressure drop across a flow restrictor inside the flow controller 108 to maintain a substantially constant rate of fluid flow to the probe 112.
• Pressure disturbances upstream or downstream of the flow controller 108 may change the pressure drop across the flow restrictor and thus may change the fluid flow rate to the probe.
• the membrane 232 will pressed against the nozzle 213 closing the nozzle 213 from the venting chamber 230. This will increase the fluid pressure on the downstream side of the flow regulator, decrease the pressure drop across the flow restrictor, and decrease the fluid flow rate to the probe 112.
• Compensation structure in the flow controller 108 opposes the reduction in pressure drop and restores it to a desired value, which thus restores the flow rate to the desired value.
• the compensation structure may be a valve in the inlet of the flow controller 108 that opens to increase the pressure on the inlet side of the flow restrictor by admitting more fluid into the controller 108 to thereby increase the aforementioned pressure drop.
• a decrease in the fluid pressure on the downstream side of the controller 108 will result in compensation in the opposite direction to maintain a substantially constant desired flow rate to the probe 112.
• Pressure disturbances on the inlet side of the controller 108 such as pulsations from the pump 104, will result in similar compensatory action.
• tonometric pressure data is measured to a precision of about 0.05 mmHg or better, and at a cadence of at least 50Hz (i.e. with a 20ms sampling interval), then valuable clinical information may be obtained.
• the sampling rate is at least 100 Hz or more, for example, 200 Hz and above.
• the device of Figures 1 and 2 obtains the raw probe input pressure from tee. From the tee, a new line (which is necessarily at the same pressure as the probe line) connects to a pressure measurement and digitization device 118. In concert with an embedded controller, the pressure data are repeatedly downloaded and logged.
• the data derived from the apparatus described above may be of very high quality (14 bit or higher pressure digitization at 100 Hz, or higher), exceeding the precision and temporal sampling requirements laid out above.
• Signal processing code in the computer 122 automatically reads the data generated from the device and extracts a number of key parameters. More specifically, programs stored in the computer 122 sense the onset and the end of the pulsatile region of the pressure data, then extract the pulse rate, pulse amplitude, pulse volume, and the ocular blood flow. The algorithms also return the mean IOP.
• the nozzle 213 at the end of the thin-walled tube 212 is initially in contact with the flexible membrane 232.
• the tube 212 is connected to a pneumatic pump 104 that increases the gauge pressure inside the pressure chamber 202.
• the membrane 232 transfers the pressure from the pump 104 to the cornea of the eye 115.
• the inward directed force produced by the air pressure overcomes the outward force induced by the IOP and is expressed by corneal deformation known as applanation.
• a gap between the tube 212 and the membrane 232 forms, releasing air from the tube 212 into venting chamber 230 and out of the probe 112 through the vents 236, and thereby stabilizing the probe pressure.
• the invention is not limited to this specific relationship between probe pressure and intraocular pressure.
• the relationship should be a relationship observed empirically with the specific equipment being used to implement an ocular blood flow analyzer. This empirical observation can be made by comparing the pressure measured by the blood flow analyzer to a known pressure produced by any accurate mechanism that simulates the pressure of an actual eye, such as a water or air manometer, or other model eye. Measuring in vivo an actual eye having known pressure characteristics may also be used. Use of an illustrative water manometer to accomplish this task is described below. Another possibility for a manometer is a pulsatile air manometer that produces square wave pressure pulses.
• the probe pressure oscillates at an amplitude of around 6 to 10 mmHg and at a frequency equal to the pulse rate.
• the pulsatile ocular pressure oscillates at an amplitude of about 3 to 5 mmHg.
• Figure 4 shows an example of raw pneumatic probe pressure data from an ocular blood flow analyzer in accordance with one example of the invention.
• Figure 6 shows an example of IOP derived from probe pressure data. Note in both cases the oscillatory signal impressed upon the overall pressure signal.
• the patient's pulse rate is about 77 beats per minute.
• the tonometric probe pressure oscillates in response to ocular blood vessels cyclically swelling due the systolic/diastolic cycle.
• the process may be summarized as follows. As the ocular blood vessels swell, the ocular volume is increased. The increased ocular volume is resisted by the elasticity of the eye thereby increasing the IOP. This is analogous to a balloon being inflated. At larger volumes, the membrane is more tightly stretched and the internal pressure is higher.
• Figures 3A, 3B, and 3C comprise a flow chart depicting in detail the operation of a first example of the software in the computer 122 that performs the signal processing part of this invention.
• the program begins in start block 300 in Figure 3A.
• the program loads into memory a text file of raw pressure data collected from the probe 112.
• the software in block 302, creates a plot of the raw probe pressure data as a function of time 400, as shown in Figure 4.
• the data from the probe 112 contains a significant amount of noise that could adversely influence the measurement of intraocular pressure and blood flow. At least some of the noise can be removed without adversely influencing the pressure measurements by applying a judicious amount of smoothing to the noisy data.
• a boxcar filter or running mean filter is used to smooth the data.
• Such a filter maps each data point in a set of unsmoothed data to another data point in a set of smoothed data.
• the boxcar filter averages the value of that data point with a predetermined number of other data points in the unsmoothed data set. That average then becomes one of the data points in the smoothed data set.
• a nine-element boxcar filter is used. The boxcar filter averages each pressure sample from the probe 112 with four of the immediately previous in time pressure samples from the probe 112 and four of the immediately next in time pressure samples from the probe 112.
• the average of these nine data points in the samples from the probe 112 then becomes a smoothed pressure sample in the smoothed data set.
• the degree of smoothing is tunable and is related to the length of the running mean (i.e. the length of the 'boxcar').
• the boxcar filter described here be used to smooth the probe pressure data, other smoothing circuits or filters that can remove noise in the probe pressure data may be used instead.
• the coherence length of a signal is the period of that waveform.
• the coherence length of the noise may be the period of the lowest unwanted frequency component of the noise.
• the coherence length could also be the average, or some other mathematical function, of the periods of the noise
• the smallest coherence length of interest in the signal is the period of the smallest frequency of interest, in this case, the pulse rate.
• the patient described above had a pulse rate of 77 beats per minute, or 1.28 beats per second. The period thus is 1/1.28 seconds or 780 milliseconds.
• the inventors have found that a suitable temporal width of the boxcar filter in this example of the invention is about 200 milliseconds, which is substantially smaller than the coherence length of the signal of interest. This translates into a boxcar temporal width of 9 or 11 consecutive samples when the sampling rate is 100 Hz.
• the coherence length of the noise is substantially below the 200 millisecond boxcar length, since good smoothing is obtained at this temporal boxcar width.
• the software automatically senses the onset and the end of the pulsatile regions of interest (ROI). This is done by differentiation of the smoothed input signal in block 304 in Figure 3A.
• a differentiated signal is one that represents the local slope of the original signal. The slope of the pulsatile region never exceeds a certain value, but when the measurement begins, and when the measurement ends, there is a rather large slope change, positive at the beginning of the measurement interval, then negative at the end of the
• FIG. 5 shows the smoothed and differentiated raw probe pressure data 500. Regions of high values 502 and regions of low values 504 in the differentiated raw data correspond to regions of highly positive and highly negative slope, respectively, in the raw probe pressure data. Such regions only occur at the onset and termination of the measurement, in other words at the beginning and end of the region of interest.
• These uniquely high slope regions may be sensed in block 306 in Figure 3A by a thresholding method, giving the locations of the onset and end of the pulsatile region of interest (ROI).
• the software compares the values of the differentiated raw data and compares them to positive and negative thresholds. When the positive threshold is exceeded, the software has identified the start of the measurement. When the differentiated raw data is less than the negative threshold, the end of the measurement has been identified. The pulsatile region of interest is between the beginning and end of the measurement.
• the software then truncates the raw pressure data by contracting the found ROI by about one second, for example, in block 308, to be sure that the region being sampled truly represents the pulsatile region, and does not contain any excessive signal due simply to onset or termination pressure fluctuations.
• the probe pressure is converted to intraocular pressure in block 309 in Figure 3A using a probe-specific lookup table or analytic expression.
• the IOP as a function of time 600 in the region of interest is shown in Figure 6.
• the program proceeds from Figure 3A to Figure 3B as indicated by block 310 in Figure 3A and block 311 in Figure 3B.
• a straight-line regression then is performed in block 312 in Figure 3B.
• a first order polynomial initially is fitted to the data.
• the mean IOP is determined in block 313.
• the IOP value at the center of the straight-line fit to the ROI is the mean IOP.
• the software then fits a 3rd order polynomial to the pulsatile ROI in block 314. See Figure 6, which shows a 3rd order polynomial curve 602 fitted to the ROI after the raw data are converted from probe pressure 400 to IOP 600.
• the 3rd order curve 602 generally runs through the inflection points of the oscillatory portion of the IOP waveform 600 in Figure 6.
• the 3rd order polynomial found above is subtracted from the truncated data in block 316 to remove the DC offset and any overall tendency of the data to slope or curve. This results in flattened and zero-based IOP data 700 in Figure 7.
• the peaks 704 and valleys 706 shown in the IOP data are identified in block 318.
• a copy of the data is created, which is then smoothed. That data is then differentiated. After differentiation, any zero-crossings correspond to peaks and valleys in the original data.
• the peak detection circuit has tunable amount of noise-suppression. All instances of smoothing in the code use the same subroutine that implements boxcar smoothing described above, with the possible exception of differing boxcar lengths depending on how much tolerance there is for smoothing. The subroutine thus takes a smoothing length as a parameter.
• the first step in the peak- finding process in block 318 is to smooth the IOP data 700 because the next step is to differentiate the data, which by its nature tends to amplify noise.
• the smoothing can be fairly aggressive at this point since the only interest here is in finding where in time the true peaks are, not what their amplitudes might be.
• Differentiation transforms a signal to something that represents the local slope. A local extremum (peak or valley) can only occur when the slope is zero.
• the computer 122 searches for where the signal goes to zero. Since there are only a limited number of data points, it is unlikely that any single data point will actually equal zero, and there could be several points competing for the title of closest to zero.
• the computer 122 determines whether a zero is a peak or a valley. To accomplish this, the computer 122 in block 318 constructs another boxcar, which it slides along the differentiated signal while monitoring whether its polarity is negative or positive. It does not matter what the average value is, just whether the average is positive or negative. At the moment that the average computed by the boxcar in block 318 changes polarity, the boxcar is centered on the zero point. If the result of the boxcar computation is going from positive to negative, then the boxcar is centered on a peak; if result of the boxcar computation is going from negative to positive, then boxcar is centered on a valley.
• the positive peaks in the IOP data found in block 318 are averaged to produce a mean positive peak; the negative peaks in the IOP data are averaged to produce a mean negative peak.
• the computer 122 subtracts the mean negative peak from the mean positive peak to obtain the pulse amplitude.
• the mean IOP computed in block 313 is added to the flattened zero-based IOP data of Figure 7 in block 321 to obtain a flattened IOP pulsatile region 800 shown in Figure 8.
• the horizontal line 802 in Figure 8 represents the mean IOP added to the flattened zero-based IOP data in Figure 7.
• the program proceeds from Figure 3B to Figure 3C via blocks 323 and 324.
• the mean pulse rate is found in block 325 in Figure 3C, for example, by counting the number of positive peaks in the ROI and dividing the number of positive peaks minus one by the time between the first and last peaks.
• the pulse rate PR is as follows:
• PR (np - l)/(t n - to), where n p is the number of peaks, t n is the time of the last peak, and to is the time of the first peak.
• Eye volume is based on averages given by Silver and Geyer for males and females, and sex selection in the software. According to Silver and Geyer, for the average human, the change in eye volume is:
• Figure 9 shows an IOP pulsatile signal that has been translated into change in eye volume curve 900 by applying the pressure-volume relationship (1) above.
• the positive peaks 1000 and the negative peaks 1002 in the AV data shown in Figure 10 are found and averaged.
• the computer uses the peaks found in block 322 to perform the averaging in block 328. Peaks in the IOP data occur at the same time coordinates as the peaks in the delta-V data, since the transformation to delta-V data affects only the dependent variable, not the time variable. So, the computer looks at the value of the delta-V data at the time coordinates found in block 322, The negative mean peak is subtracted from the positive mean peak to obtain the mean peak- to-peak pulsatile AV. In block 330, the mean AV above is divided by the mean pulse duration to obtain the Net Pulsatile Flow.
• d(AV)/dt the instantaneous rate of change of volume
• d(AV)/dt the results of which are shown as curve 1100 in Figure 11. Smoothing similar to that described above is performed here as well for noise suppression. Again, there is a trade-off between getting reliable data and impact on the magnitude of the data.
• the positive peaks of d(AV) /dt 1102 are found and averaged in block 334. This average is the mean instantaneous Peak Net Pulsatile Flow. This also is computed both in terms of ⁇ /s and ⁇ ,/ ⁇ .
• the Mean IOP, Pulse Rate, Pulse Amplitude, Net Pulsatile Flow, and the Peak Net Pulsatile Flow are presented for display in block 336.
• This blood flow data may be presented on any display, for example, a computer display, in graphical or numerical form. An illustrative numerical display is shown in Figure 12.
• the software in computer 122 computes various ocular perfusion pressures in block 336.
• the computer 122 acquires the systolic and diastolic components of the patient's arterial blood pressure, SBP and DBP, respectively. These numbers may be determined through the use of a traditional sphygmomanometer or electronic blood pressure measurement apparatus and entered into the computer 122 by the operator of the blood flow measurement system.
• the computer 122 generates the mean ocular perfusion pressure in accordance with the following relationship:
• the computer 122 also generates the systolic ocular perfusion pressure in accordance with the following relationship:
• diastolic OPP DBP - mean IOP.
• Figure 13 is a flow chart depicting in detail the operation of the software in the computer 122 that performs the signal processing part of this invention in accordance with a second example of the software.
• the program begins in start block 1300.
• the program loads into memory a text file of raw probe pressure data collected from the probe 112.
• the software in block 1304, creates a plot of the raw data as a function of time 1400, as shown in Figure 14.
• the software automatically senses the onset and the end of the pulsatile regions of interest (ROI). This is done by differentiation of a smoothed version of the input signal in block 1306.
• a differentiated signal is one that represents the local slope of the original signal. The slope of the pulsatile region never exceeds a certain value, but when the measurement begins, and when the measurement ends, there is a rather large slope change, positive at the beginning of the ROI, then negative at the end of the ROI, respectively.
• Figure 15 shows the differentiated raw data 1500. Regions of high value 1502 and regions of low value 1504 correspond to regions of highly positive and highly negative slope, respectively, in the raw probe pressure data. Such regions only occur at the onset and termination of the measurement, in other words at the beginning and end of the region of interest.
• These uniquely high slope regions may be sensed in block 1308 by a
• the software compares the values of the differentiated raw data and compares them to positive and negative thresholds. When the positive threshold is exceeded, the software has identified the start of the measurement. When the differentiated raw data is less than the negative threshold, the end of the measurement has been identified.
• the pulsatile region of interest is between the beginning and end of the measurement.
• the software then restricts the raw pressure data by contracting the found ROI by about one second, in block 1310, to be sure that the region being sampled truly represents the pulsatile region, and does not contain any excessive signal due simply to onset or termination pressure fluctuations.
• the probe pressure is converted to intraocular pressure in block 1312 either using a probe-specific lookup table or analytic expression.
• Figure 16 shows an illustrative intraocular pressure curve 1602
• the software then fits a 3rd order polynomial 1604 to the pulsatile ROI in block 1314. See Figure 16, which shows a 3rd order polynomial curve 1604 fitted to the ROI after the raw data are converted from probe pressure to IOP.
• the 3rd order curve generally runs through the inflection points of the oscillatory portion 1602 of the waveform in Figure 16.
• the software now truncates the data in block 1316 to include only the ROI, and the 3rd order polynomial found above is subtracted from the truncated data in block 1318 to remove the DC offset and any overall tendency of the data to slope or curve thereby creating a zero-based flattened ROI.
• the mean IOP across the ROI is then computed in block 1320 and added back to the signal in block 1322 resulting in a flattened ROI at the mean IOP pressure. See Figure 17, which shows a flattened ROI 1700.
• the mean IOP represented by the horizontal line 1702, has been added back to the zero-based ROI derived in block 1318 to obtain the flattened ROI 1700 in Figure 17.
• Eye volume is based on averages given by Silver and Geyer for males and females, and sex selection in the software. According to Silver and Geyer, for the average human, the change in eye volume is:
• Figure 18 shows an IOP pulsatile signal translated into change in eye volume curve 1800 by applying the pressure-volume relationship (1) above. Relationships (2) and
• the varying volume signal found above is now smoothed and differentiated in block 1326 to get the instantaneous pulsatile flow as a function of time. Then, the positive regions are integrated in block 1328 across the ROI to obtain the net pulsatile inflow, which is then divided by the duration of the ROI in block 1330 to get a pulsatile inflow rate in ⁇ /s. The same process is applied in block 1330 to the negative regions in block 1328 to obtain the net pulsatile outflow and the pulsatile outflow rate. See Fig.
• Figure 19 shows the instantaneous pulsatile flow curve 1900.
• the positive regions are integrated and averaged to get the average inflow rate; the average outflow rate is found in a similar manner with respect to the negative regions.
• the zero-based flattened IOP derived in block 1330 is smoothed and differentiated in block 1332 to identify the positive and negative pulsatile peaks.
• Figure 20 shows both the positive peaks 2000 and the negative peaks 2002 in the flattened, zero-based IOP region of interest identified by the peak-finding process. This yields both the pulse rate in block 1334 and the mean peak- to-peak pressure in block 1336.
• the blood flow data and the curves derived by the digital signal processing circuitry of Figure 13 may be displayed on a computer display or may be numerically displayed as illustrated by Figure 12.
• Gosling's pulsatility index data may be derived from the described data generated by the software in the computer 122.
• Gosling Pulsatility Index (PI) is a measure of the variability of blood velocity in a vessel, equal to the difference between the peak systolic and minimum diastolic velocities divided by the mean velocity during the cardiac cycle.
• the equivalent Gosling pulsatility index is based on flow rather than velocity.
• Illustrative electronic equipment that may be used to implement the ocular blood flow measurement system shown in Figures 1, 2, 3A-3C, and 13 include a computer 122 in the form of a Raspberry Pi microcomputer running a fully functional desktop operating system.
• the computer 122 may be implemented as a series connected Raspberry Pi and an chicken micro controller, the chicken being in series with the Raspberry Pi between the pressure sensor 112 and the Raspberry Pi to provide shielding for the Raspberry Pi.
• the electronics that processes the data generated by the apparatus of Figure 1, 2, 3A, 3B, 3C, and 12 does not have to be a computer.
• the suitable electronics may be any form of signal processing circuitry that is capable of producing the results described herein.
• the brushless DC motor 102 and pump 104 may be a Series 1410VD, Model 14100216, diaphragm pump with integral brushless DC motor, made by Gardner Denver Thomas, Inc. of Sheboygan, Wisconsin.
• the pressure compensated flow controller 108 may be a Model PCFCD-1N1-E BRS, pressure compensated flow controller, made by Beswick Engineering Co., Inc. of Greenland, New Hampshire.
• the pressure sensor 116 may be a Model MS4525 printed circuit board mounted pressure transducer, made by Measurement Specialties of Fremont, California. Operational Discussion
• the basis of the device operation is to contain and measure pressure within a column of gas, for example, air, that varies in flow and pressure only as a function of the source pressure and resistance encountered at the distal end of a hollow tube 212 and nozzle 213 vented to the atmosphere by holes 236 in the tip 228.
• a column of gas for example, air
• the gas enters the probe body 200 through an inlet tube 208. Initially it flows through the probe body 200 via a hollow passage or pressure chamber 202, and into and through a shaft 212, into and through the unobstructed probe tip via a jet or nozzle 213 at the distal end of the shaft 212.
• the pressure in the probe body is low at this point as no major obstruction is present other than the size restriction of central passage 222 and the nozzle 213.
• probe 112 Before eye contact is made, probe 112 is in a free flow condition, or as close to a free flow condition as is allowed by particular proportions of the probe used.
• a forward motion of the shaft 212 occurs at approximately 10-20 mmHg in this condition. This is because the difference in surface area of the proximal end of shaft 212 central passage 222 creates an opposing surface which allows a portion of pressure in the probe body 200 to exert itself against the rear of the shaft 212. Greater pressure and flow in free flow condition results in proportionally greater forward thrust of the shaft 212. Greater pressure and lower flow caused by opposing exit of air from tip 228 by contact with the eye causes a similar increase in thrust against the eye by probe tip 228.
• Equilibrium is achieved when air pressure causes deflection of the cornea allowing a gap between the membrane 232 and nozzle 213 in probe tip 228.
• a softer eye allows a gap to form sooner at a lower pressure and flow.
• a harder eye resists the formation of a gap more strongly, resulting in a gap forming at a higher pressure and flow, as well as a greater forward pressure of the entire probe/tip assembly 112/228.
• the pressure in probe body 200 increases in a known manner, for example, proportionally, in response to the resistance to flow caused by an individual eye against the tip 228 and nozzle 213.
• a standing column of air pressurizes tube 210
• Accuracy and time elapsed while acquiring a steady measurement are affected by user technique, unique characteristics of the eye, change in the eye caused by the measuring process, and constant changes in alignment of the probe nozzle 213 with the surface of the cornea causing variable pressures, and with changing IOP caused by the measuring pressure eliciting physiological changes in the IOP.
• the axis of nozzle 213 should intersect the cornea chord at 90 degrees to ensure a gap develops that is uniform around the entire perimeter of jet 213.
• a misalignment causes a leak-down of pressure readings, accompanied by initial difficulty in obtaining a repeatable, sustainable and measurable pressure rise and fluctuation over time. Effect of increased forward pressure of the shaft tip combination on the eye at higher pressures is to increase the ease of measuring, but it increases the applanation of the cornea and results in distortion of the eye and may cause egress of fluid through the outflow channels of the eye.
• the user must rely more on tactile feedback than any self-regulation of the amount of probe extension. Greater ease of measurement is desirable as it results in shorter contact time with the eye, less opportunity for the eye to adapt to the measuring forces, and increased safety margin. Greater ease of measurement is obtained at greater flow and pressure, but may be undesirable in terms of accuracy of measurement and distortion of the eye.
• Pressure changes are transmitted through tube 210 to the silicon based micro- machined pressure sensor.
• an on board ASIC digital signal processor
• a digital data stream is transmitted to a computer software module where the data is captured to a file. On obtaining a clean measurement, the probe is removed from the eye, pressure drops as probe resumes free flow mode, and recording ceases.
• a second software module is launched and auto loaded with the most recent data capture file.
• An analysis of the data is made by novel software. The results, including pulse amplitude, pulse rate, pulse blood flow/second, pulse blood flow/pulse, OPP, SPP, and DPP may be printed to screen, showing a user report along with a graphic depiction of the wave form analyzed.
• Apparatus in accordance with this invention immediately acquires measured pressure conditions, and produces a stable average pressure over time, as compared to the results obtained by the prior art devices.
• All the machines have a tip and membrane assembly that constrains the airflow by initial resistance of the tip itself when the tip is not in contact with the eye.
• the eye/membrane interface is deformed so as to remove the constraint against the escape of air from the probe.
• Pressure in the probe is intended to have a linear relationship to IOP at a higher pressure than the eye itself.
• An initial pressure and flow is set at the air source to provide a constant component of thrust of the probe/plunger toward the eye. Without this pressure the probe will not engage with the eye surface; no measurement would occur.
• the base pressure set with no eye contact is a critical element, as are the pressure and flow across a full scale of 10 through 40 mmHg eye pressure.
• the manometer 2100 comprises a base 2102 and a graduated cylinder 2104 set in the base 2102.
• the graduated cylinder 2104 contains a predetermined amount of fluid such as water.
• the bottom of the graduated cylinder communicates with one end of a horizontal passage 2106 formed in the base 2102.
• the other end of the passage 2106 is connected at its other end to a cylindrical pressure chamber 2108 formed in the base 2102.
• a flexible bladder 2110 analogous to the outer surface of a human eye seals the top of the pressure chamber 2108.
• the height of the fluid in the graduated cylinder 2104 is set such that the fluid pressure in the pressure chamber 2108 bearing against the bladder 2110 simulates the intraocular pressure in the human eye.
• the known fluid pressure against the bladder 2110 can be compared to readings obtained by the pressure measurement system described here when the pressure probe tip 114 is placed against the outer surface of the bladder 2110.
• Apparatus in accordance with the invention also is substantially more linear than prior art apparatus. Pressing the probe against a diaphragm type manometer that produces a known pressure that simulates actual IOP and plotting the pressure in the probe against the known manometer pressure illustrates the linearity of pressure measurements taken by a probe in accordance with the invention. The non-linearity of the prior art probe may be determined by doing the same thing with the prior art probe.
• Figure 24 is a plot of probe pressure versus manometer pressure for both the probe in accordance with the invention and a prior art probe referred to as a predicate device. Curve 2400 represents the probe pressure versus manometer pressure for a prior art probe. Curve 2402 represents the probe pressure versus manometer pressure for a probe in accordance with the invention.
• the curves 2400 and 2402 illustrate that a probe in accordance with the invention is substantially more linear than the prior art probe in a range of 10 - 30 mmHg manometer pressure, which is where actual patient IOP's are expected. Note the sharp deviation from linearity in the 15 to 10 mm Hg descending range of the prior art probe. Also, the pressure in the prior art probe is substantially higher than in the probe in accordance with the invention. For example, note the rather high probe pressure of 50 mmHg corresponding to 10 mmHg
• a new prototype machine was created with the objectives of testing the existing tips and probes with visibility into the nature of the wave forms representing ocular pressure, efficient data collection, and plotting in real time.
• Very accurate control over pressure and flow (volume) in the air supply was achieved by replacing the pressure regulator mentioned above with a pressure compensated flow control device 108 and using a more constant torque pump motor 102, specifically the newer generation DC motor called a Brushless DC Motor.
• the conjecture was that consistent flow control of air to the tip might result in substantially more repeatable accurate measurements without much effort or training on the part of the user, while providing full scale linear 10 to 40 mmHg measurements. This is counterintuitive, as a pressure compensated flow controller operates by increasing/decreasing pressure at the output to maintain a constant flow.
• the flow rate is pre-set using an integral user accessible needle valve and lock nut.
• a device that attempts to maintain constant flow in a system that measures pressure does not initially make sense. It would seem this technique would increase pressure independently of the measurement input from the probe tip. The hypothesis was that it would sustain repeatable constant flow across a full range of IOP/probe pressure while enabling a low pressure high flow initial calibration setting of the probe/air supply system. This theoretically would enable more consistent and repeatable measurements in the critical range of IOP from 15-17 mmHg. Results of experiments comparing the test prototype and the legacy machines proved that the test prototype device operates in a more consistent and reliable manner extending to lower IOP's, while also operating at an accurately controlled relatively low probe pressure. Manipulation of the probe shows a robust and repeatable contact with the eye that results in an immediate commencement of measureable pulse forms, without any special effort by the user.
• the pressure compensated device 108 dampens noise more effectively than a pressure regulator because it can be preset accurately at a low pressure, but is designed to operate over a range of output pressures while maintaining constant flow, versus the pressure regulator, which seeks to maintain a constant pressure against any outside force from the eye.
• Use of a pressure compensated flow controller 108 maintains a linear relationship between eye and probe pressure reliably into the most important 10-17 mmHg ranges of IOP, as well as consistent measurements over the entire range from 10-40 mmHg.
• Figure 24 shows precise correlation in slope ratio between manometer and probe pressure with probe pressure needed to read 10 mmHg is 30 mmHg below the legacy probe pressure of 50 mmHg. Lower pressures may equate to a better margin of safety when exerting pressure on the eye with a stream of air.
• the probe pressure is a much more linear function of the intraocular pressure in the region where most patients measure, that is, in the IOP range of about 10 to 30 mmHg. Compare curve 2400 with curve 2402 in Figure 24.
• the quality of pulsatile eye pressure signals obtained by the device described herein is markedly different from that of the legacy units.
• devices in accordance with this invention attain contact with the eye and begin to measure immediately, with a constant steady signal waveform plotted. See Figure 22. This waveform is of high resolution and maintains a steady value. Unlike the legacy device, the pulsations do not rise or fall once a measurement is attained.
• the present invention enables measurement of variations of intraocular pressure with high accuracy and repeatability to be obtained without undesirably large forces being exerted upon the eye.
• This patent document discloses a novel system that produces accurate, stable, and repeatable low distortion measurements of intraocular pressure as a function of time from which reliable data about ocular blood flow can be derived.
• Abnormal intraocular pressure and ocular blood flow can be an indication of abnormalities and diseases in not only the eye, but also in other parts of the body.
• a health care provider can assess the intraocular pressure and ocular blood flow data in conjunction with appropriate clinical correlation to identify those abnormalities and diseases. Clinical correlation may include other tests, observations, and historical patient information. For example, in the eye, a health care provider can detect glaucoma, macular
• a health care provider can also use ocular pressure and blood flow data to help detect Alzheimer's disease, carotid occlusive disease, systemic disease, and cerebral vascular disease. Additional conditions that can be monitored by analyzing ocular pressure and blood flow data are burned skin, and the cerebral vascular flow, edema, and pressure associated with traumatic brain injury. Intracranial pressure in newborns can also be monitored in this way. Also, a health care provider can detect changes in ocular blood flow and can provide appropriate therapeutic interventions such as pharmaceuticals and surgery in response to changes in ocular blood flow.
• Appendix 1 is a source code listing of a first example of a computer program written in the Python computer language that may be loaded onto the Raspberry Pi implementation of the computer 122 to accomplish the digital signal processing functionality described in Figures 3A, 3B, and 3C.
• Appendix 2 is a source code listing of a second example of a computer program that accomplishes the described digital signal processing functionality of Figure 13.
• smPulsPeak int(smPulsPeak/1000.0/tStep) # Now number of elements.
• boxPuls 9 # Boxcar detection length for peaks (best if odd-numbered).
• peaksPositive np.empty(O) # smoothLen: boxcar smoothing length pre-differentiation.
• peaksNegative np.empty(O) # tStep: used by differentiation (gradient) function.
• avg np.mean(data[0:window_len]) # windowjen: length of window in which to sense polarity change.
• peaksNegative np.appendfpeaksNegative, (itindex + window_len/2)-l)
• peaksPositive np.appendfpeaksPositive, (itindex + window_len/2)-l)
• posPeakValues np.appendfposPeakValues, data[posPeaks[i]])
• negPeakValues np.appendfnegPeakValues, data[negPeaks[i]])
• posPeakValues np.appendfposPeakValues, data[posPeaks[i]])
• negPeakValues np.appendfnegPeakValues, data[negPeaks[i]])
• posPeakValues np.appendfposPeakValues, data[posPeaks[i]])
• plt.fi gure(figsize (10,10)) plttitle (title)
• diastolic float(diastolic)
• meanlOPvec np.full(dataROI.size, meanlOP)
• pulseamplitude meanPeaktoPeak(dataROIzero, pksPosA, pksNegA)
• pulsetime (tROI[pksPosA[pksPosA.size-l]] - tROI[pksPosA[0]]) / (pksPosA.size - 1) elif pksPosA.size ⁇ pksNegA.size:
• deltaVdat deltaV(dataROIadded, sex)
• pulsevolume meanDeltaV(deltaVdat, pksPosA, pksNegA)
• netPulsFlow pulsevolume * (pulserate/60.0) # Differentiate deltaV to get the instantaneous rate of volume change (which is the instantaneous blood flow).
• deltaVdatSmooth smooth(deltaVdat,smPuls) # Smooth the input signal
• pulsFlow np.gradient(deltaVdatSmooth,tStep) # Differentiate the smoothed signal to obtain flow.
• pksPosB peakfind(pulsFlow[0:tROI.size],smPulsPeak, tStep, boxPuls)
• PeakNetPulsatileFlow meanPeakPos(pulsFlow, pksPosB)
• Peak Net Pulsatile Flow (uL/s): " + str(np.around(peakNetPulsatileFlow,l)) print Peak Net Pulsatile Flow (uL/m): " + str(np.around(peakNetPulsatileFlow*60,l)) if OPP
• Diastolic BP (mmHg) : " + str(int(diastolic))
• diastolicOPP diastolic - meanlOP
• AddDataName raw_input("Do you want the datafile name to appear on the plots? (Y/N)? ")
• plt.savefig('IOP.png', transparent True) # Flattened Zero-Based ROI with peaks annotated.
• circlesNegX np.zeros(pksNegA.size)
• circlesNegY np.zeros(pksNegA.size)
• circlesPosX np.zeros(pksPosA.size)
• circlesPosY np.zeros(pksPosA.size)
• circlesNegX[i] tROI [pksNegA[i]]
• circlesNegY[i] dataROIzero[pksNegA[i]]
• circlesPosX[i] tROI[pksPosA[i]]
• circlesPosY[i] dataROIzero[pksPosA[i]]
• circlesNegX np.zeros(pksNegA.size)
• circlesNegY np.zeros(pksNegA.size)
• circlesPosX np.zeros(pksPosA.size)
• circlesPosY np.zeros(pksPosA.size)
• circlesNegX[i] tROI [pksNegA[i]]
• circlesNegY[i] dataROIadded [pksNegA[i]]
• circlesNegX np.zeros(pksNegA.size)
• circlesNegY np.zeros(pksNegA.size)
• circlesPosX np.zeros(pksPosA.size)
• circlesPosY np.zeros(pksPosA.size)
• circlesNegX[i] tROI [pksNegA[i]]
• circlesPosX[i] tROI[pksPosA[i]]
• circlesPosY np.zeros(pksPosB.size)
• circlesNegX[i] tROI [pksNegB[i]]
• circlesNegY[i] pulsFlow[pksNegB[i]]
• circlesPosX[i] tROI[pksPosB[i]]
• circlesPosY[i] pulsFlow[pksPosB[i]]
• plt.savefig('InstPulsFlow.png', transparent True) pltshowQ
• time_sample 0.02 # time of sample 20 ms
• timeArray np.arange(0, xRange, time_sample)
• btn2 QtGui.QPushButton('Stop')
• btn3 QtGui.QPushButton('Save')
• timeArray np.arange(0, xRange, time_sample)
• timer2 QtCore.QTimerQ
• timer2.start(l) # read data as soon as possible
• fileName '7home/pi/instrument/data/''+datetime.datetime.nowQ.strftime(''%m ⁇ %d-%Y_%H:%M:%S")+txtFName.textO + txtLName.text() + ".png"
• xAxis.setTicks [[[(0,int(s-(xRange+l))),(l,”),(2,int(s-(xRange-l))),(3,”),(4,int(s- (xRange-3))),(5,”),(6,int(s-(xRange-5))),(7,")],[]])
• # smooth A simple boxcar average smoothing function.
• # is the size of the boxcar filter.
• peaksNegative np.appendfpeaksNegative, (itindex + window_len/2)+l) elif it[0] > 0:
• posPeakValues np.appendfposPeakValues, data[posPeaks[i]])
• negPeakValues np.appendfnegPeakValues, data[negPeaks[i]])
• time np.linspace(0.0, data.size*timestep, data.size)
• timeRoi time[roileft:roiright]
• dataROIzerobased dataRoi - cubic(timeRoi)
• deltaV deltaV(dataROIadded, sex) if drawplots:
• avgPulsFlowPos totalPulsFlowPos / (np.max(timeRoi) - np.min(timeRoi))
• avgPulsFlowNeg totalPulsFlowNeg / (np.max(timeRoi) - np.min(timeRoi))
• npeakpairs floatfpeaksPos.size + peaksNeg.size) / 2
• pulsetime (timeRoi[peaksPos[peaksPos.size-l]] - timeRoi[peaksPos[0]]) / (peaksPos.size - 1)
• pulserate 60.0 / pulsetime
• circlesPT np.zeros(peaksPos.size)
• circlesNT np.zeros(peaksNeg.size)
• circlesN [i] dataROIzerobased[peaksNeg[i]]
• circlesP[i] dataROIzerobased[peaksPos[i]]
• circlesPT[i] timeRoi[peaksPos[i]]

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