Ultrasound ,n Med. & Bml. Vol. 17. No. 5, pp. 445452, Printed in the U.S.A.

0301-5629/91 $3.00 + .oO 0 1991 Pergamon Press plc

1991

@Original Contribution EVALUATION

OF DOPPLER ULTRASOUND PERFUSION MEASUREMENTS

FOR BLOOD

Department of Electrical Measurements, Lund Institute of Technology, P.O. Box 118, S-221 00 Lund, Sweden Abstract-The need to develop clinical methods for the noninvasive monitoring of regional blood perfusion, i.e., the blood flow through the very fine capillaries in body tissue, has long been felt. Hitherto existing methods exhibit limitations, such as insufficient measurement depth and poor time- or space-resolution, which restrict the measurements that can be performed. Dymling (1982) introduced a new CW Doppler ultrasound method for noninvasive blood perfusion measurement which might be one possible solution to this problem. Preliminary experiments indicated a correlation between blood flow and measured perfusion value. Unexpectedly large variations in the recorded perfusion values lead to further investigation of the method, both in vitro using a specially designed flow phantom and in viva This study indicates that at least some of the large variations recorded are the result of measurement errors caused by movement artifacts or ultrasonic signal interferences. Methods to diminish the effects of these artifacts are discussed. Key Words: Ultrasonic Doppler, Blood perfusion, Artefact suppression.

INTRODUCI'ION

Unfortunately, a commercially available clinical instrument which in a simple, reproducible and affordable way, can perform continuous in-vivo regional blood perfusion measurements, especially at selected spatial measuring sites situated deeply below the body surface, is yet to be developed. Plethysmography and tracer techniques are clinical methods which are sometimes utilized, but with strong limitations in time as well as spatial resolution. The laser Doppler flowmeter (Nilsson et al. 1980) shows great promise, but can only be used to measure blood perfusion at depths of less than 1 mm from the body surface, due to the limited penetration of the laser beam in the tissue. Today’s ultrasonic Doppler instruments are not suited for blood perfusion measurements. First, they are designed to present the mean, maximum or eventually the spectrum of the blood flow velocities in a vessel, rather than the net inflow of blood to a selected body tissue. Second, the possible degradation of the Doppler signal resulting from different motional artifacts is normally reduced by addition of a steep high-pass filter with a comer frequency corresponding to a blood flow velocity around 5 cm/s. In this way, the measurement of normal arterial blood flow velocity is hardly affected at all, whereas the lowfrequency nonflow-originating Doppler signals, such as those caused by reflections from pulsating blood vessels or from structures affected by breathing move-

The function of blood circulation has excited man’s curiosity from the earliest times. William Harvey was the first to demonstrate the circulation of blood in 1628, but it was not until the beginning of the present century that quantitative methods become available for the study of blood flow in man. Recent additions to the battery of clinical and research tools for noninvasive measurement of blood flow include ultrasonic Doppler and magnetic resonance imaging techniques. The use of such advanced instruments now makes it possible to make noninvasive semiquantitative clinical recordings of the blood flow in the larger vessels inside the human body. However, it is the blood flow through the very fine capillaries in body tissue, i.e., the blood perfusion, that controls the transport of nutrients and cell products to and from the tissue. The performance of this transport mechanism is, therefore, of utmost importance for the biological function of the said body tissue. An accurate and reproducible measure of the regional blood perfusion would bring knowledge of the clinical status in the examined tissue and also better knowledge of the clinical status of the patient.

Address all correspondence

to Roger Eriksson. 445

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Ultrasound in Medicine and Biology

ments, are strongly attenuated. Unfortunately, this will also result in adversely distorted Doppler recordings of venous blood flow and a total rejection of Doppler signals from the very low velocity perfusion blood flow (in the order of mm/s). Dymling (1982) suggested a different ultrasonic CW Doppler method for measurement of regional blood perfusion. This ultrasonic blood perfusion (UBP) method overcomes some of the shortcomings associated with the conventional Doppler method and shows many advantages as compared to other existing blood perfusion measurement methods: it is continuous, noninvasive, offers a good control of the measuring sample volume and can be applied to tissues deep inside the body. Early experiments to record the blood perfusion of the index finger (Dymling 1985, 199 1) indicated a dependency between the ultrasonically measured blood perfusion value and the actual blood perfusion: a complete arterial occlusion of the blood flow to the finger resulted in a decrease of the blood perfusion value which again increased when the blood flow in the tissue was restored. However, the measured perfusion value from a given sample volume showed surprisingly large variations with time. During an experiment, lasting for a few minutes, the variation between measured maximum and minimum perfusion values could be as large as a factor of ten or more. Similar results were reported for blood perfusion measurements with the laser Doppler flowmeter (Tenland 1982), where the large variations were explained as local changes in the diameter of small arterioles (vasomotion) causing variations in the local perfusion. But as the sampling volume of the UBP-meter is many times larger than that for the laser Doppler flowmeter, one would expect the ultrasonically obtained measurement to average the regional variations in blood perfusion, and it is therefore questionable if the recorded changes in the measured blood perfusion can be explained only in terms of local vasomotion. Obviously, the UBP-method needs further evaluation, and the objective of this work is to perform experiments, in vitro and in vivo, to investigate whether the large recorded variations in the measured perfusion values are local physiological variations in the perfusion or measurement artefacts. Finally, the outcome of this evaluation will form a basis for the development of new instrumentation hardware, designed to adapt the UBP-method for clinical use. MATERIALS

AND METHODS

Blood perfusion Blood perfusion of tissue is normally defined as the blood volume flowing through a unit volume of

Volume 17, Number 5, 1991

tissue per second, i.e., m l/(cm 3.s) . Anatomically, this corresponds to the total blood flow, in the investigated tissue volume, through the microvascular exchange vessels, i.e., the capillaries. From this definition Dymling (1985, 199 1) has shown that blood perfusion in general can be expressed as

P - N,E{V]

(1)

where P is the blood perfusion of the tissue, N, is the number of red blood cells in the investigated tissue volume, and E{ V) is the mean velocity of the red cells in the investigated tissue volume. By examining the auto correlation of the receiver signal in a Doppler flowmeter, blood perfusion can be calculated from the Doppler spectrum according to

where f is the frequency and S(f) is the Doppler power spectrum. This relation is derived and valid for a vascular network where all flow directions are equally probable.

Instrumentation The instrumentation used in the study was partly specifically designed by us and partly standard components and instruments. Except for the transducers, the same instrumentation was used in all experiments. A block diagram of the UBP-meter can be seen in Fig. 1. The electrical signal from the oscillator is applied to the transmitting ultrasound transducer to generate an ultrasound wave propagating into the measurement region. Part of the wave is reflected or scattered back to the receiving transducer, where the ultrasound wave is transformed into an electrical signal, which is amplified in the subsequent low-noise preamplifier. Its output signal contains the flow information modulated on a high-frequency carrier, and the demodulator is used to extract the Doppler shifted signal. The Doppler signal processor performs the mathematical operation described in eqn (2) to estimate the blood perfusion value. The master oscillator was a Hewlett-Packard model HP3325A quartz-controlled frequency synthesizer. Coherent demodulation (principle described in Atkinson and Woodcock 1982) was performed with a double balanced mixer (Mini Circuits SRA-1). The Doppler signal processing can be performed in either the frequency domain (Dymling 1985) or in the time domain (Nilsson et al. 1980). The frequency-domain

Evaluation of Doppler ultrasound 0 R. ERIKSONet al.

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Oscillator

Amplifier

Doppler Signal Processor

Demodulator

Display

Fig. 1. Ultrasonic blood perfusion system based on the Doppler principle.

surements of blood perfusion flow velocities up to about 6 mm/s using an ultrasound frequency of 10 MHz. In the next block the signal is squared using a fast analog multiplier before it finally is time-averaged in an electronic operational integrator with a suitable selected time constant. As a result, this Doppler signal processor can present an estimated blood perfusion value with a relatively short delay, mainly dependent on the selected time constant in the integrator averager. It is an important issue to select the right averaging time constant in the integrator, as it will not only affect the response time of the Doppler signal processor but also the accuracy of the weight taken to lowflow velocity components in the measured blood perfusion value. A short response time will reduce the accuracy at low blood perfusion flow velocities as well as make it possible to easily identify short-duration motional artifacts. As a compromise, the time constant was set corresponding to a 3-dB cut of frequency of 2.9 Hz for the Doppler signal processor during invivo measurements. Doppler shifts around and below this cut of frequency will not be completely averaged but cause ripple on the UBP-signal after the integrator.

method is a straightforward approach easily implemented but time consuming. To produce a fast-responding UBP-meter, suitable also for the clinical applications, signal processing in the time domain (see below) was chosen for our experiments. The estimated blood perfusion value was during in-vitro experiments measured with a Hewlett-Packard model HP3478A multimeter and during in-vivo experiments on a Y-T thermal recorder, Graphtec model WR7600. A more detailed description of the analog circuitry implementation of the time domain Doppler signal processor is shown in Fig. 2. The output signal from the demodulator is low-pass filtered in a first order 200-Hz low-pass filter to remove unwanted high-frequency modulation products and noise. The next operation is a passage through a specially designed filter bank which produces an output signal proportional to the square root of the frequency. This multiplication has the adverse effect to further enhance the remaining high-frequency noise components in the signal. Therefore, the signal is again filtered through another second order 200-Hz low-pass filter to suppress these noise components. The cut-off frequency of these filters was selected to allow for mea-

Doppler difference signal

4

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Fig. 2. Block diagram of Doppler signal processor in the frequency domain (a) and in the time domain (b).

i-Perfusion I

(a)

Doppler difference signal

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(W

Squarer

-W Averager

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Two continuous-wave (CW) Doppler ultrasound transducers, with resonance frequencies of 5 and 10 MHz, were manufactured for the experiments. Each transducer has two separate transducer elements, one transmitter and one receiver. The elements (piezoelectric ceramic Pz27, Ferroperm A/S, Denmark) were carefully mounted and shielded to minimize the ultrasonic and electric cross-talk between the transmitter and the receiver elements. High sensitivity was obtained by using airbacked transducer elements minimizing the ultrasound losses inside the transducer housing. The ~-MHZ transducer was used for the in-vitro experiments. This transducer has two disc-shaped elements, 5 mm in diameter, mounted in two separate housings to be able to arrange the transmitting and receiving areas for different measuring sample volumes. The 1O-MHz transducer was designed for in-vivo experiments. To enable a measuring sample volume located close to the surface of the transducer, a discshaped piezoelectric element, 5 mm in diameter, was cut into two half-discs. These two elements were mounted in the same housing, slightly separated at an angle of 135”. The center of the measuring sample volume is located at a depth of about 5 mm under the transducer surface, with a resulting sample volume of about 0.2 cm3. During some of the experiments, the frequency of the transmitted ultrasound will be changed, so it is important that the bandwidth of the measurement system be sufficiently large to permit these changes of the frequency. The combined relative 3-dB bandwidth of the ultrasound transducers used and attached amplifiers was measured to be larger than 20% of the employed center frequency. The maximum frequency deviation from the center frequency used in any of the experiments was less than 5%. In-vitro experiments A special test object was manufactured for the in-vitro assessment of the performance of the ultrasound blood perfusion method (see Fig. 3). Taking the human blood perfusion system as the design prototype, it is obvious that the test object in addition to the normal flow-containing tube should be fabricated from a reasonably tissue mimicking material, and also contain a number of echo reflecting structures. The dimensions of the test object is 55 X 40 X 35 mm and it is molded from Wacker@ SilGel 604. A thinwalled silicon tube (external diameter 3 mm) extends through the model, 15 mm below the upper surface. Further, 10 mm below the tube is a layer of randomly distributed 2 X 2 X 3 mm Duretan@ granules, which

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Ultrasonic

transducers

Seohadex-alvcerol mi&’ Magnetic

Fig. 3. Measurement

configuration ments.

stirrer

for the in-vitroexperi-

will be the prime source for non-Doppler-shifted ultrasound reflections back to the transducer. The resonance frequency of the broadband ultrasound transducer used was 5 MHz, and was driven with a constant oscillator amplitude of 0.5 V,, . The two parts, one transmitter and one receiver, of the Doppler transducer, were applied to the test object with an angle of 45” to the flow direction in the silicon tube and an intermediate angle of 90” between the two transducers. Good acoustic coupling to the test object was ensured by specially shaped transducer mounts on the model and the use of ample amounts of ultrasonic transmission gel. During the experiments, both the reflected ultrasound signal received by the Doppler transducer and the resulting blood perfusion values were recorded. Forty contiguous values were collected for each measurement and the mean as well as the standard deviation of these were calculated. Instead of blood, which can present certain practical problems during longer measurement sessions, a scattering fluid which comprises of a mixture of glycerol and Sephadex@ particles has been used. These spherical dextran gel particles, with a distribution of diameters between 10 and 40 pm, were added in the ratio 5 g Sephadex@ to 80 cl glycerol. A magnetic stirrer was used to keep the particles floating, and thus maintain the concentration uniform throughout the fluid. The circulation of the mixture through the test object was handled by a peristaltic roller pump which generated an average flow velocity of 4 mm/s. Two different measurements were performed. In the first series of experiments, the frequency was set to 5 MHz, and a screw clamp was attached right across the test object and tightened step-wise so as to slightly deform it, and thereby change the structure and positions of the reflecting interfaces inside the flow model. The maximum change in width of the model was

Evaluation of Doppler ultrasound 0 R. ERIKSSON et al.

about 3 mm, or less than 10% of the original width. In the second sequence of experiments, the oscillator frequency of the ultrasound blood perfusion measurement system was increased from 4.9 to 5.1 MHz in incremental steps of 10 kHz between each reading. In-vivo experiments These experiments were carried out to investigate the possible existence of correlation between the blood perfusion in vivo, and the signal obtained from the UBP-meter described above. The measurements were performed on a healthy young male subject. Before the measurements, the subject was resting for 20 min. The IO-MHz ultrasound Doppler transducer was applied halfway up from the ankle joint, on the dorsal part of the calf. The transducer was treated with a surplus of ultrasonic gel and fixed to the calf with an adhesive tape. The transmitter element was driven with a signal from the oscillator with an amplitude of 0.5 V,,. During the experiments, the subject was lying down on his back resting as comfortably as possible. Arterial occlusion was ensured by inflation of a blood pressure cuff, and positioned around the thigh to a pressure of 230 mmHg. RESULTS AND DISCUSSION

The objective of the first in-vivo experiment was to study the presence of a relationship between measured and in-vivo blood perfusion, and to accomplish this in the most basic way. The ultrasound transducer was positioned on the calf of a young healthy male. Fig. 4 shows a UBP-meter recording from these measurements. One minute after the start of the experiment, the pressure cuff was inflated to inhibit the blood supply to the region. The inflation of the cuff took about 15 s, and is easily noticed on the recording due to the large motional artifact it caused. The blood flow was then inhibited for about 1 min before the

Inflation Fig. 4. In-vivo

of cuff

Arterial

449

cuff was again deflated. In Fig. 4, it can be seen that, before the occlusion, there were relatively large UBP signals, synchronous to the heart beat, which almost disappeared during the occlusion, except for occasional large-amplitude movement artifacts. After the deflation of the cuff, the perfusion signals returned, but not to the same extent as before the occlusion. Since the sample volume is placed at a depth of about 5 mm below the skin surface, one possible explanation could be that the measurement was made in the subcutaneous layer, a region not likely to have an increased blood supply after an occlusion. Fig. 5 shows the recording from the next in-vivo experiment, also made on the calf, where careful precautions were taken not to interfere with the perfusion during the experiment. Also here, the occasional large peaks going to the top of the diagram are easily identified as motional artifacts. It is important to notice the large, almost rhythmic variations in the trend of the perfusion signal. We know, from the previous experiment, that the recorded perfusion signal is related to the blood perfusion flow, but does the perfusion in a selected sample volume vary to the extent indicated in Fig. 5, or is the variation induced by some artifacts in the UBP-method? The next experiment was designed to answer that question. During the two following in-vitro experiments, both the amplitude of the returning ultrasound echoes at the receiving transducer and the estimated ultrasound blood perfusion signal were recorded (see Figs. 6 and 7). The amplitude measurements turned out to be very reproducible with a standard deviation below 2%. The perfusion measurements had a significant larger standard deviation as compared to the amplitude measurements. Each measurement point shown in Figs. 6 and 7 is a mean value with the standard deviation added as an error bar. The first in-vitro experiment is performed with constant flow through the test object, but with increas-

occlusion

Deflation

of cuff

perfusion measurement on the right calf. After 1 min, an arterial occlusion was induced by inflating a blood pressure cuff. The inflation took 15 s, and the arterial occlusion was maintained for 1 min.

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0

60

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120

Fig. 5. Blood perfusion measurements

180

240

time (s)

on the calf of a resting subject. The recording shows large low-frequency variation in the measured values.

ingly distorted geometry as a result of the screw clamp being pulled tighter. Fig. 6a shows the amplitude of the received ultrasound signal, and Fig. 6b the estimated blood perfusion value. There clearly occurs a change in the measured perfusion signal during the

deformation of the flow model, and this is not correlated to a corresponding change in the intensity of the ultrasound signal returning to the transducer. The second in-vitro experiment is performed with constant flow through the test object and without

100 80 60 40 20 0 0

1

100

; cl3

. g

80 60

I

I

I

0

1

2

Distortion

(mm)

3

03 Fig. 6. Perfusion measurement recordings of (a) the amplitude

on a flow model which was increasingly geometrically distorted. Simultaneous of the received signal at the Doppler transducer, and (b) the perfusion value.

Evaluation of Doppler ultrasound 0 R. ERIKSSONet al.

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10.1 to 9.9 MHz in two lOO-kHz steps. It can be seen that the estimated blood perfusion value is lower at 10.0 MHz than at 9.9 or 10.1 MHz. Figure 8b shows a perfusion recording when the ultrasound frequency first is raised from 10.3 to 10.4 MHz for about 30 s, and then back to 10.3 MHz. In this experiment, it is the higher ultrasound frequency that estimates the lower perfusion value. When the frequency again is lowered to 10.3 MHz, the estimated blood perfusion value returns to a higher level. It should be pointed out that the peak value in the perfusion signal recorded immediately after a change of ultrasound frequency is an artifact caused by the frequency change itself.

any distorsion of the geometry. Instead, the frequency

of the transmitted ultrasound was varied from the start frequency of 4.9 MHz to the stop frequency of 5.1 MHz in increments of 10 kHz. Fig. 7a shows the amplitude of the received ultrasound signal, and Fig. 7b the estimated blood perfusion value. Also, this experiment results in a change in estimated blood perfusion value even though nothing else is changed except for the transmitted ultrasound frequency. Actually, one would expect the estimated perfusion value to increase as a function of frequency, as the ultrasound Doppler shift is proportional to the transmitted frequency. The increase in frequency is about 4%, and should therefore give a rise in Doppler frequency shift with about the same amount. On the other hand, the increased frequency might also influence the attenuation of the sound beam. Finally, the frequency of the transmitted ultrasound was changed in steps of 100 kHz during an ongoing in-viva experiment. Fig. 8a shows the perfusion recording when the frequency was lowered from

CONCLUSION This study has investigated the performance of the UBP-meter, focusing especially on the large Iluctuations in the estimated blood perfusion value recorded during ongoing experiments. The accomplished model experiments indicate a relation be-

- _

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Transmitted

0

frequency

(MHz)

60 time

30

(s)

(a) frequency

Transmitted

0

3’0

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(s)

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Fig. 8. In-vivo blood perfusion measurements. During the measurement, the frequency of the transmitted ultrasound was (a) reduced from 10.1 to 9.9 MHz in three steps; and (b) raised from 10.3 to 10.4 MHz, and back to the original 10.3 MHz.

tween the UBP-value and the true blood perfusion for a given body tissue, but the large fluctuations in measured perfusion values can also be a result from influences other than an actual change in the blood perfusion. For example, it was shown that a minor change in the geometry of the test object was sufficient to cause a large fluctuation in the recorded perfusion value. A possible explanation to this is in the influence from wave interference, caused by reflections of ultrasound from boundaries and interfaces inside and outside the measuring sample volume, causing local maxima and minima in the ultrasound intensity pattern. Due to this interference pattern, blood perfusion in different parts of the measuring sample volume will not produce the same contribution to the estimated perfusion value. Instead, there will be parts giving higher contributions, as well as parts giving lower contributions. A small change of the geometry of the volume, through which the ultrasound beam is passing, can then alter the ultrasound interference pattern, causing other parts of the measuring sample volume to give a higher contribution to the measured blood perfusion value. The interference pattern can also be altered by a shift in the transmitted ultrasound frequency. A change in the transmitted ultrasound frequency also

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means a change in the propagated ultrasound wavelength. This will, of course, create new interference patterns. Thus, the fluctuation shown in Fig. 8 is also probably a result of a frequency-induced change of the existing ultrasound interference pattern. Why will the interference pattern change during an ongoing experiment, and thus change the recorded perfusion value? There can be many different causes for this. One reason in vivo might be slow movements of internal structures and boundaries in the ultrasonic field, caused by changes in the blood pressure or by muscles movements. Another reason can be a changed direction for the ultrasound beam or some other movement of the ultrasound transducer. As a result of the described ultrasonic interference patterns, it seems to be difficult to produce totally reproducible in-vim measurements with the present UBP-meter. On the other hand, if the effects of this interference could be reduced, we will get a clinical method with many interesting features. The last in-vim measurement showed how a change in the transmitted ultrasound frequency caused a change to the recorded perfusion, a change that is probably more or less randomly distributed. If this is the case, it should be possible to reduce the variation and improve the reproducibility through the use of many simultaneously transmitted ultrasound frequencies and subsequent averaging using the estimated perfusion value from each frequency. Preliminary results from such a new multifrequency ultrasound blood perfusion measurement system have shown promise. Acknowledgements-This work was supported by the Swedish National Board For Technical Development (grant nos. 8802403 and 8901742).

REFERENCES Atkinson, P.; Woodcock, J. P. Doppler ultrasound and its use in clinical measurement. London: Academic Press; 1982:22-74. Basler, S.; Vie& A.; Anliker, M. Measurement of tissue blood flow by high frequency Doppler ultrasound. In: Huch, H.; Huch, R.; Rooth, G., eds. Advances in experimental medicine and biology, Vol. 220-Continuous transcutaneous monitoring. New York: Plenum; I987:223-226. Dymling, S. 0.; Hertz, C. H.; Persson, H. W. The measurement of blood perfusion in tissue. Proceedings of the Fifth World Congress of Ultrasound in Medicine and Biology, Brighton, UK 26-30 July 1982. Dymling, S. 0. Measurement of blood perfusion in tissue using Doooler ultrasound. PhD Thesis. Deoartment ofElectrical Measurements, Lund Institute of Technology, Sweden. LUTEDX/ (TEEM-1027)/l-4/(1985). Dymling, S. 0.; Persson, H. W.; Hertz, C. H. Measurement of blood perfusion in tissue using Doppler ultrasound. Ultrasound Med. Biol. 17:433-444; 199). Nilsson, G. E.; Tenland, T.; Oberg, P. A. Evaluation of a laser Doppler flow-meter for measurement of tissue blood flow. IEEE Trans. Biomed. Eng. 10~597-604; 1980. Tenland. T. On laser Doooler flowmetrv. Ph.D. Thesis. Deoartment. of Biomedical Engineering, Lidkoping University, Sweden. No. 83; 1982.

Evaluation of Doppler ultrasound for blood perfusion measurements.

The need to develop clinical methods for the noninvasive monitoring of regional blood perfusion, i.e., the blood flow through the very fine capillarie...
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