102

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. BME-24, NO. 2, MARCH 1977

A Staged, Tortuous Microcapillary System: Hemodialysis and Ultrafiltration A. ZELMAN, M. CATHEY, M. SCOTT, AND R. RHODES

Abstract-A microcapillary system has been constructed by etching staged, tortuous microchannels onto stainless steel sheets; a flat membrane over the channel completes the microcapillary. The microcapillaries are sinusoidal and terminate every X to form a region where the output of all capillaries mix. There are three stages per centimeter. Tortuosity and staging greatly reduce boundary layers and increase mass transfer through the membrane. There are 430 000 capillaries per square meter of membrane. The capillaries are oval (125 Am X 250gim); thus membrane area per unit volume is only 36 percent less than a hollow-fiber capillary. Initial testing indicates satisfactory clearances and hydraulic conductivities for Cuprophane PT-150 membrane. This microcapillary system should find many industrial and biomedical applications where prime volume per unit membrane area and concentration polarization are major concerns of design.

fluid flows it tumbles over the mesh, forming secondary currents perpendicular to the membrane surface. Solutes concentrated on the surface of the membrane are washed away by the secondary currents. Screens, however, may present problems in blood-membrane systems for two reasons. First, there could be a region of high shear where the blood path was partially occluded by the screen. Hemolysis and other blood damage result from regions of high shear [2]. Secondly, screens could allow small stagnant pools of blood to form on the sides of the mesh not parallel to the flowing blood; stagnant pools generate thrombi [3]. Most of the commercially available hemodialyzers have incorporated a means of inducing secondary currents. This ha. been accomplished by providing on the dialysate side a mem NOM ENCLATURE brane support which causes the membrane to form a nonAi slope of least square straight line for species i, defined planar surface. The membrane supports have generally been of two varieties: plastic mesh or screen type and the multiby (2), min/ml; point type. Blood is sandwiched between two layers of memCi clearance of species i, ml/min; solute clearance of species i extrapolated to infinite brane. The hydrostatic pressure of the blood forces the Cl membrane into the pattern of the screen or multipoint surface; blood flow rate, ml/min; thus a somewhat tortuous blood path is generated. The gentle (Q)P concentration of species i in plasma, moles/I; Jv volume flux across the membrane, ml minm m-2 mixing induced by this means has significantly decreased the Lp hydraulic conductivity evaluated from altering trans- mass transfer resistance of the boundary layer. A list of memmembrane hydrostatic pressure, ml - m 2 minm - brane support types and their manufacturers may be found torr-1 in [14]. A different approach to membrane transport in artificial hydraulic conductivity evaluated from altering trans(Lp), membrane osmotic pressure, ml m2 * min' torr-1; organs is to neglect the formation of secondary currents APPm arithmetic average transmembrane pressure, torr; altogether and concentrate 'engineering on minimizing the size of the system. A very simple artificial kidney has been QB flow of whole blood, ml/min; marketed as a result of the development of thin-walled hollow QD flow of dialysate, ml/min; fibers only 200 pm in diameter (Dow Chemical Corp., Walnut r radius of microchannel, cm; Ai\ m arithmetic average, transmembrane osmotic pressure, Creek, CA). The hollow fibers are placed in a cylindrical tube so that the individual tubules are straight. The boundary layers torr. will surely be maximized in these devices [4]. Artificial kidneys have evolved along two basic patterns: C ONCENTRATION polarization is the nemesis of the memsheet membrane systems with a mat next to the membrane to _ branologist. No matter what the device or method used, induce secondary currents and hollow-fiber systems which do the boundary layer persists [1]. However, there are many denot induce secondary flows, but have a minimal prime volume. vices which have been developed to mitigate the effect of the of the research reported here was to combine the The purpose boundary layer on the overall membrane transport coefficients. best engineering aspects from the flat-sheet and hollow-fiber For example, for reverse osmosis or electrodialysis processes, We devices. wished to develop a system with greater tortuosity a screen is generally placed in the path of the fluid. As the than previously achieved with the flat-sheet membrane systems yet retain the character of the low prime volume of the Manuscript received April 23, 1975; revised September 23, 1975 and hollow-fiber system. Hence we developed a staged, tortuous February 17, 1976. microcapillary system (STMS). A. Zelman is with the Center for Biomedical Engineering, Rensselaer Polytechnic Institute, Troy, NY 12181. STMS has many positive features. First, the microcapillaries M. Cathey, M. Scott, and R. Rhodes are with the Departments of have been formed with tortuous blood paths which induce Physiology and Internal Medicine, Meharry Medical College, Nashville, TN 37208. secondary currents in the flowing blood. Second, the micro-

ZELMAN et al.: MICROCAPILLARY SYSTEM FOR HEMODIALYSIS

capillaries are internally staged; that is to say, we provide a short section of tortuous capillaries and then a very short section where the output of all capillaries can mix. The mixed solution then enters another section of tortuous capillaries. The method of staging channels offers both the advantage of having mixed blood enter each new microcapillary (thus the time for boundary layer buildup is minimized) and the great advantage of not having a single thrombus occlude an entire capillary. If a hollow-fiber tubule becomes occluded, the entire tubule stops operating. However,if a staged microcapillary is occluded, only the flow in a small segment is stopped and the fluid merely shunts around the occlusion with all capillaries thereafter open and operating normally. Third, these staged, tortuous microcapillaries can be used with any flat membrane, so as new developments in membrane technology occur, the developed hardware will not become obsolete. Fourth, these microcapillaries are oval and thin; a large area of membrane is exposed per unit capillary volume. METHODS

Production of the Staged, Tortuous Microcapillary System Fifteen mil thick, 316 type stainless steel sheet was chemically etched such that the center of the channel depth was 125 ± 25 gim and the breadth of the channel on the uppermost surface was 250 ± 25 ,gm. The standard deviation in the depth and breadth is due to intrinsic variations in the stainless steel's resistance to chemical etching. Etching was done on both sides of the sheet for the blood channels and on one side for the dialysate channels. Fig. 1 is a schematic of a cross section through one module of the STMS. Blood flows on both sides of the central metal sheet within the microchannels. When the module is clamped together, a nylon screen presses against the membrane to hold it firmly against the microchannels. The blood is completely confined to flow in the tortuous microchannel. Note that the cross section of each microchannel is equivalent to approximately half that of a round tubule. The membrane surface area of a microchannel is proportional to 4r whereas the surface area of a round hollow fiber is proportional to 2irr = 6.3r. The etched microcapillaries have only 36 percent less surface area per unit volume of capillary than hollow fibers of comparable diameter. The device was designed to be modular. Two modules were manufactured for this prototype. The modular unit is depicted schematically in Figs. 2 and 3. Blood enters the device via a blood manifold. Two grooves have been machined in the manifold such that blood is delivered to 1 -in-OD tubes and dialysate is delivered to port holes in each support plate. A silicone rubber gasket ensured separation of blood from the dialysate. The 1 -in-OD tube is sandwiched in between two membrane layers. When the device is clamped together, the membrane acts as a gasket. The inflowing blood forces the membrane into tapered grooves machined in the support plates; thus, a tapered channel surrounded by membrane is formed which is initially about i-in in diameter and linearly tapers to zero depth across the device. Blood flows out of the tapered channel across both sides of the etched plate into a

103

DIALYSATE--

0.4&40AFRy~ -

MICROCHANNEL SCREEN

*--MEMBRANE

7 '4NMICROCHANNELS

BLOOD

4-MEMBRANE DIALYSATE -SCREEN ,~ /7 ~ 7 MICROCHANNEL

Fig. 1. Schematic cross section through one module. TAPERED BLOOD GROOVE

MICROCHANINEL PLATE

SUPPORT

PLATE-7-

/ /

W

SCR

W

MEMBRANEe

NHN~ SCR

E

E

N

MICROCI

PLATE

DIALYSATE INPUT BLOOD INPUT TUBES

Fig. 2. Detailed drawing of one module of STMS. PRESSURE PLATE

POSITIONING

BODBLOOD GASKET MANIFOLD

DIALYSATE

LMS

CO

SUPPORT

PLATES

BLOOD

BASE

BLOOD

BLOOD MANIFOLD

GASKET

Fig. 3. Unit as assembled with two modules of STMS.

tapered groove on the other side of the microcapillary array and exits through another short length of i-in tubing. Dialysate enters the device, through the blood manifold and is distributed to a plate with microchannels covered by a nylon net. The dialysate flows perpendicular to the blood flow. As shown in Fig. 3, the device became an extremely simple assemblage of parts. The support plates and blood manifold were machined from Lexan sheet stock (General Electric Co., Pittsfeld, MA). All machined surfaces were polished. The support columns held the support plates evenly aligned and were constructed of black-anodized aluminum. Plasma Clearances Clearances were evaluated from experiments with less than 20-torr transmembrane pressure difference; hence, essentially no ultrafiltration occurred. The clearance equation used was

=i (i)n-

(ign

out

QB

(1)

0.119 Ci is the clearance of the ith species as indicated from plasma determination, and QB is the volume flow of whole blood.

104

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, MARCH 1977

* 105 mlmin, in vivo Y

o 105 ml/min, in

*

A

D3 -Z t

ml/min,

A 225

ml/min, in vitro

in

v0vo

* 305 ml/min, in vivo

V7

t

vitro

A 225

40[

a 305

ml/min, in vitro

Y 435

ml/min, in vivo

V 435 ml/min, in vitro 0

ImI/min)m-2 MEMBRANE

Fig. 4. Hydrostatic pressure drop across blood path. This figure demonstrates the pressure increase at the blood inlet as a function of blood flow. The blood outlet is at atmospheric pressure and the dialysate flow is constant at (500 ml/min) * m-2 membrane. All flow values are normalized to 1 m2 of membrane.

Clearance values reported herein are normalized for 1 m2 of membrane, i.e., the membrane area of the device was 0.1 19 M2; there are no corrections for "effective" membrane area. Blood from anephric dogs was taken from the carotid artery and returned to the jugular vein. An occlusive pump (model 7561-20, Cole-Parmer Co., Chicago, IL) maintained constant blood flow. Blood and dialysate pressure were measured by a pressure gauge (model 11-1-3000, Sweden Freezer Manufacturing Co., Seattle, WA) and a pressure transducer (model P 7 D, Celesco Corp., Canoga Park, CA). Cuprophane PT-150 membrane was used throughout the experiments. The system was set up for single-pass dialysis. Dialysate was an NaCl solution at 0.9 gm% and 390C. For the in vitro experiments either human or dog blood was diluted to (20 ± 0.5) percent Hct in order to simulate the hematocrits of renal failure patients. Uric acid, urea, glucose, and phosphate buffer were added to produce initial whole blood concentrations in the range of 10 mgm% uric acid, 100 mgm% B.U.N., and 200 mgm% glucose. All animal experiments used heparin as the anticoagulant. Initial dosage was 120 USP units of heparin per kilogram of body weight and 25 units/kgm additionally per hour thereafter.

20

20

40

60

80

PD (torr)

Fig. 5. Blood and dialysate hydrostatic pressure as a function of blood flow. In vitro experiments used blood at 20 percent Hct. Dog experiments had Hct's of about 40 percent. The hydraulic resistance of the dog's circulation and higher blood viscosity produces higher pressures in the dog experiments as compared to the in vitro experiments.

of the dialysate path. Hence when the blood pressure is increased at constant dialysate flow rate, the dialysate pressure increases due to the membrane stretching into the path of the dialysate. This effect is depicted in Fig. 5.

Solute Clearances Figs. 6 and 7 depict the clearance values as defined under the section on methods. The value of the clearance at infinite blood flow, Cl, was obtained as the Y-intercept for the least square straight line of the linear equation. In Ci In Cj"'o =

Ai (QP

(2)

where Ai is the slope of the straight line and In C, is the Yintercept. QD was maintained at 500 (ml/min) * m2 while QB was increased.

Ultrafiltration

The hydraulic conductivity, Lp, of membrane systems has been developed for multicomponent membrane systems in a Ultrafiltration experiments used freshly drawn human or general way by Zelman [5 ]. However, the complexity of most dog blood set at 20 percent hematocrit with isotonic saline. hemodialyzer systems prevents complete characterization. For these experiments approximately 1000 gin of blood was The hydraulic conductivity, Lp, can be accurately defined by placed in a container on a top loading balance (model P1210, [ A/v Mettler Instrument Corp., Princeton, NJ). The loss of weight r1m )v] (3) was recorded as a function of time and transmembrane A(APm))i~rmvoconst. Aorm=0 pressure. The prime volume of the system was evaluated from where .v4-is the volume flux across the membrane (ml/min) the initial weight loss and the volume of blood in the tubing. m 2, APPm is the average transmembrane hydrostatic pressure RESULTS difference in torr, ATm is the average osmotic pressure difference across the membrane in torr. The equality in the above Blood Flow Resistance Fig. 4 depicts the pressure drop across the device as a func- equation for Lp holds when Lp is not a strong function of the tion of blood flow at a constant dialysate flow of (500 ml/ concentration profile across the membrane. The hydraulic min) m-2 membrane. The low resistance to hydraulic flow conductivity cane also be approximated by the relationship of this device is typical of multiparallel plate dialyzers. The flow of blood and dialysate is in adjacent microchannels. The AJv (4) (Lp )7 [.A(A1Tm~)J APM=const. hydraulic resistance of the blood path is nearly equal to that

Ultrafiltration and Prime Volume

M

ZELMAN et al.: MICROCAPILLARY SYSTEM FOR HEMODIALYSIS

TABLE I HYDRAULIC CONDUCTIVITY, LP X 10 (ml -m min-

INOR. PHOS.

*

o ca 2 320

O

z

240

200 a E

.E E

[Lp] Am=°0

APm

F

a

CD

F

160

A.,

120

~

~ 0

A~~~~~~

* 0~~~ a0~~~~~~~~~~~

80 F

0

40

100

200

300

OR (ml/min)m72

400

0

MEMBRANE

torr

GLUCOSE 0

BUN

0

URIC ACID

a 0

0

z m

ej

100

0

F

80F

IE

.E E

60

[Lpr

2.07 ± 0.19 1.79 ± 0.16 2.61 ± 0.23 3.20 ± 0.29

39

.36

1.38 ± 0.12 2.48 ± 0.22 3.91 ± 0.35

Average value

2.42 ± 0.62

2.87 ± 1.54

2.59 ± 1.27

Note:

1.78 ±

.96

±

0.16

03

[Lp] AlrO

was evaluated at A irm =0, using distilled

-cons [LpJvirAim=Clst.*

was evaluated with normal whole blood diluted with 0.9 gm% NaCl to 20% Hct

[Lp]

H20.

and dialysate of 0.9 gm% NaCl. evaluated with blood as above and between dialysate concentrations of 0.85 and 0.9 gm% NaCl.

was

TABLE II COMPARED UREA CLEARANCE

APm torr

Ultrafiltration Rate

Urea Clearance

< 20

< 0.5

95

50

2

90

50*

1.2*

77

ml/min

ml/min

0

120

MU2 (a

[LpJ]Arm=const.

50 100 150 200

Fig. 6. Clearances of ionic solutes. All clearance values were from dog experiments (35 percent 6 Hct 6 48 percent) except Cl- which was obtained by flowing 0.85 gm% NaCl on the blood side and distilled water on dialysate side. Clearance values and blood flow normalized to 1 m2 membrane. Dialysate (0.85 gm% NaCl) was flowing at (500 ml/min) * m 2 membrane. Transmembrane pressure < 20 torr. These are average values from nine experiments.

140

torr')

K+

A'Cr

280

105

STMS with Cuprophane Standard Kiil with Cuprophane [8] Standard Kill with Cuprophane [7] Twin Coil UF-100 [8] Dow HFAK model#3 [8]

150

4

149

50

1

128

0

Note: QB (200 ml/min) m 2 and QD = (500 ml/min) - m 2; all values normalized to 1 m2 of membrane. Asterisk indicates estimated values. =

6

40

F

100

300

200

Q

(ml/min)m

400

CD

MEMBRANE

Fig. 7. Clearance of organic solutes. These are average clearance values from nine dog experiments. Clearance and blood flow normalized to 1 m2 membrane. Dialysate (0.85 gm% NaCl) was flowing at (500 mvmin) * m 2 membrane.

These values are listed in Table I. Note that the hydraulic conductivity increases with increased transmembrane pressure as would be expected. Hydrostatic pressure and osmotic pressure have essentially the same effect on ultrafiltration which is indicative of a membrane highly permeable to water. DIscUSSION

This paper demonstrates a valuable technique for inducing secondary flows in membrane-solution systems. These secondary flows can greatly reduce the boundary layer thickness resulting in enhanced mass transport through the membrane. The staged, tortuous microcapillary system (STMS) as described herein is an attempt to combine effectively the best design features of the hollow-fiber kidney and the flat-plate kidney. In terms of practical application, a very important feature of the STMS is that the continuous blood mixing in-

hibits coagulation. Although mitigated by recently employed quality controls, coagulation continues to be a problem in the hollow-fiber kidney [6]. Coagulation in the STMS did occur under two conditions: 1) when using outdated human blood; and 2) when the dog experiments were carried out to such lengths that the dog expired (>11 h). In both these cases capillaries were occluded and the beneficial effects of "staged" capillaries were evident. The clots would form in the first inch of the capillary array, but the blood simply shunted around and over the clots and flowed into the remaining capillaries free of clots. The staged capillary array acts as its own filter. Most commercial dialyzers induce a transmembrane hydrostatic pressure as a consequence of flowing blood through the device. For example at QB = 200 ml/min and QD = 500 ml/ min, the pressure drop across the membrane for Dow HFAK model 3 and standard Kiil is 50 torr and for the Twin Coil UF-100 it is 150 torr; whereas for STMS it is less than 20 torr [7], [8]. Thus, unlike STMS, the commercial dialyzers have an undesirable "minimum ultrafiltration rate." The higher the ultrafiltration rate, the larger the clearance value [91; this effect must be considered when comparing dialyzer clearance values for different devices. Table II lists the urea clearance values and ultrafiltration rates for a few dialyzers for com-

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, MARCH 1977

106

TABLE III HYDRAULIC CONDUCTIVITY OF CUPROPHANE MEMBRANE L X 102 (MI/min) * m-2

STMS

Standard Kiil [8] Standard Kiil [71 Laboratory test cell [7] Average values

Lp (saline)

Lp (blood)

2.42 ± 0.62 -

2.87 ± 1.54 4.0 ± ? 1.9 ± ?

2.3 +? 2.38 + 0.30 2.37 + 0.06

2.92 ± 1.05

Note: Question mark indicates standard deviation not given.

parison [71, [8]. STMS has a higher urea clearance than the standard Kiil even though the Kiil has a higher ultrafiltration rate. This indicates that the STMS has either a larger effective membrane area exposed to blood or that the more efficient blood mixing of STMS increases urea transfer by decreasing the boundary layer or both. Table III compares the hydraulic conductivity of Cuprophane evaluated with different devices [71, [8]. There appears to be considerable spread in the values of Lp for blood while Lp values for saline are consistent. There are numerous possible explanations for this including: pin holes in the membrane, hydration-dehydration of the membrane due to altered blood osmolality, protein adhesion, nonuniform blood flow paths and possibly variations in blood hematocrits. The overall performance of STMS was not expected to produce clearance values greater than those of commercially available hemodialysis systems. It is well known that hemodialysis is primarily a membrane-limited process rather than a boundary-layer-limited process [10]. However, we are looking forward to the "third generation artificial kidney" which will perform in a more biological mode. The blood in future systems will be ultrafiltered and the filtrate treated and returned to the patient just as the biological kidney does [111, [12]. When ultrafiltration becomes the predominant transport mechanism STMS will be able to offer a distinct advantage. STMS is a unique system and should find its way into many industrial and biomedical applications where prime volume and concentration polarization are major concerns of design. ACKNOWLEDGMENT The staged, tortuous microcapillary system was first designed as a membrane blood oxygenator by Dr. A. Zelman as part of his postdoctoral training under the direction of Dr. M. Weissman at Carnegie-Mellon University, Pittsburgh, PA. Research support was from Dr. Weissman's PHS grant HL12714; salary support for Dr. A. Zelman from NIH Bioengineering Training Grant 5TOlGM01455-05. It was agreed that Dr. Zelman would further investigate the potential of the staged, tortuous

microcapillary system for use as an artificial kidney and that oxygenator development would remain at Carnegie-Mellon [131. Development as an artificial kidney was initiated by Dr. Zelman at Meharry Medical College, Nashville, TN, under his PHS grant AM16802 with salary support from MBS grant 5506RR08037. The project is continuing testing as an artificial kidney at RPI under Dr. Zelman with his research support by PHS grant AM19198. The silicone rubber gasket was the courtesy of J. Aikens, IBM Corporation, Endicott, NY. Precision machining was by Gatan, Inc., Pittsburgh, PA, with R. Swann in charge. Chemical etching was by Magnetics, Inc., Butler, PA, with N. Sundel in charge. Our special thanks to medical illustrator G. Card, Meharry Medical College, Nashville, TN. REFERENCES

[1] H. L. Evans, Laminar Boundary-Layer Theory. Reading, MA: Addison-Wesley, p. 3, 1968. [2] E. E. Spaeth, G. W. Roberts, S. R. Yadwadkar, P. K. Ng, and C. M. Jackson, "The influence of fluid shear on the kinetics of blood coagulation reactions," Trans. Amer. Soc. Art. Int. Organs, vol. 19, pp. 179-187, 1973. [3] C. J. Glover, et al., "Effect of shear stress on clot structure formation," Trans. Amer. Soc. Art. Int. Organs, vol. 20, pp. 463468, 1974. [4] M. H. Weissman and L. E. Mockros, Med. and Biol. Eng., vol. 7, pp. 169-183, 1969. [5] A. Zelman, "Membrane transport: Generalization of the reflection coefficient method of describing volume and solute fluxes," Biophysical J., vol. 12, pp. 414-419, 1972. [6] J. W. Eschbach, "Evaluation of the Dow hollow fiber artificial kidney," Dep. Health, Education and Welfare, Publication (NIH) 74-248, p. 90. [7] E. Klein, J. K. Smith, F. F. Holland, and R. E. Flagg, "Membrane and materials evaluation; Permeabilities, physical and mechanical properties of hemodialysis membranes-Bember Cuprophane PT-150 Membrane," Nat. Tech. Information Service, U.S. Dep. Commerce, PB-225 070, 1973. [8] A. L. Babb, P. C. Farrell, D. A. Uvelli, and B. H. Schribner, "Hemodialyzer evaluation by examination of solute molecular spectra," Trans. Amer. Soc. Art. Int. Organs, vol. 18, pp. 98-105, 1972. [9] F. A. Gotch, J. Autian, C. K. Colton, H. E. Ginn, B. J. Lipps, and E. Lowrie, "The evaluation of hemodialyzers," Dep. Health, Education and Welfare, Publication (NIH) 73-103, 1971. [10] K. Keller, "Fluid mechanics and mass transfer in artificial organs," Special Publication, Trans. Amer. Soc. Art. Int. Organs, 1973. [111 S. G. Dharnidharka, R. Kirkham, and W. J. Kolfel, "Toward a wearable artificial kidney using ultrafiltrate as dialysate," Trans. Amer. Soc. Art. Int. Organs, vol. 19, pp. 92-97, 1973. [12] F. Lai, R. Tankersley, M. Scott, R. Rhodes, and A. Zelman, "Third generation artificial kidney: pH and concentration control," Trans. Amer. Soc. Art. Int. Organs, vol. 21, pp. 346-351, 1975. [13] M. H. Weissman, T. K. Hung, and H. S. Borovetz, "Development of the etched channel membrane oxygenator," in Proc. Ann. Conf. on Medicine in Engineering and Biology, vol. 17, p. 240, 1975. [14] The Hemodialysis Manual, Dep. Health, Education and Welfare Manual (HMS) 72-7002.