Ind. Eng. Chem. Prod, Res. Dev. 1081, 20, 1-5
1
SYMPOSIA SECTION
I.
Symposium on Advances in Polymeric Controlled Release Formulations F. W. Harris, Chairman 179th National Meeting of the American Chemical Society Houston, Texas, March 1980
Controlled Release Micropumping of Insulin at Variable Rates Mlchael V. Seflon' and Kevln J. Burns Department of Chemical Engineering and Applied Chemistry, Universtty of Toronto, Toronto, Ontario, Canada, M5S 1A4
The insulin delivery characteristics of a solenoiddriven controlled release micropump have been determined. Basal delivery obeyed Darcy's law with a permeability of 5 X lo-' cm2 for a 3/8-in. i.d. Lucite pump with a 1 cm long foam membrane. Repeated compression of this membrane by the core of the solenoid resulted in augmentatbn of the delivery rate. Augmentations in this pump were low, however, but do exhibit the expected Increases with decreasing basal rate, increasing solenoid power and increasing compression frequency. Furthermore, pump delivery was reproducible (f8%) over 25 4-h cycles of basaVaugmented operation. The incorporation of a second rate-controlling membrane (e.g., Nucleopore) in series with the foam reduced the basal rate and increased the degree of augmentation to practlcal levels. However, the formation of insulin precipitates on this membrane is a serious limitation that warrants further study.
Diabetes is the third leading cause of death in North America. Diabetics are 25 times more prone to blindness, 17 times more prone to kidney disease, five times more prone to gangrene, and twice as prone to cardiovascular disease (Canadian Diabetic Association, 1978). These degenerative complications have been linked to the accumulated effect of the periodic variation in blood glucose levels associated with conventional insulin therapy (Cahill et al., 1976). To restore normoglycemia,both open-loopand closedloop artificial pancreata have been developed. These have been recently reviewed (Santiagoet al., 1979). Subsequent to this review, other researchers have reported their success using open-loop insulin delivery systems to maintain normal glucose levels in dogs (Goriya et al., 1979; Blackshear et al., 1979) and in humans (Irsigler and Kritz, 1979; Pickup et al., 1979; Tamborlane et al., 1979; Kolendorf et al., 1980). Although there is some evidence supporting the relationship between normoglycemia and the degenerative sequelae of diabetes (Albisser and Leibel, 1977; Tchobroutsky, 1978) longer-term experiments are required before the beneficial effects of normoglycemia can be demonstrated. Although controlled release formulations have been modified for continuous insulin delivery (Creque et al., 1980; Langer et al., 1980) these formulations cannot satisfy the variable rate delivery requirements of the artificial pancreas. A controlled release micropump which has variable rate delivery characteristics is shown in Figure 1. With a suitable supply of insulin connected to the pump, the concentration and/or pressure difference across the membrane results in diffusion or bulk transport through
the membrane(& This is the basal delivery and requires no external power source. Augmented delivery is achieved by repeated compression of the foam membrane by the coated mild steel piston. The piston is the core of the solenoid and compression is effected when current is applied to the solenoid coil. Interruption of the current causes the membrane to relax, drawing more drug into the membrane in preparation for the next compression cycle. The basal rate is determined by the magnitude of the concentration and/or pressure difference and by the permeability of the membrane&). The augmented rate is a function of the elastic properties of the membrane, the force applied by the solenoid piston, and the frequency of compression. Two advantages are suggested for the controlled release micropump relative to the more conventional peristaltic pump or valve arrangements that are being used for insulin delivery (e.g., Santiago et al., 1979; Goriya et al., 1979; Blackshear et al., 1979). First, augmented delivery is achieved without valves and hence the mechanical unreliability associated with the inadvertent opening and sticking of valves or with complex pumps is avoided. (However, the long-term stability of the foam membrane, especially after repeated compression, is important and is being investigated.) Furthermore, basal delivery is not achieved by regulating a larger flow rate. Failure of the micropump, therefore, cannot result in uncontrolled delivery at the maximum rate and cause an inadvertent insulin overdose. The significance of these suggested advantages, however, remains to be demonstrated. Energy consumption is low since power is consumed only during the postprandial delivery phase and even then only 0 1981 American Chemical Society
2
Ind. Eng. Chem. Prod. Res. Dev., VoI. 20, No. 1, 1981
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SOLENOID CORElPlSTON
PUMP BARREL
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LUCITE ROD FOAM “MEMBRANE POROUS SUPPORT
Figure 1. Schematic diagram of controlled-releasemicropump. The rate controlling membrane is not present in every prototype.
during the compression portion of the cycle (power for typical therapeutic use: 60 W peak X 5 ms “on” time = 0.3 J/compression stroke, with a frequency of 30 strokes/min). The device is small enough (7 cm3, 35 g) to be easily implanted. Although the initial prototypes will be used as transcutaneous devices (with only the outlet under the skin), an implantable device can be envisaged complete with percutaneously refillable drug reservoir and rechargeable power pack. A metal bellows device (Blackshear et al., 1979) or a stretched elastomer bladder (Leeper et al., 1977) would be appropriate for maintaining a constant, though high, pressure difference for intravenous delivery. More likely, however, a constant concentration (unit activity) reservoir would be used; for example, the insulin controlled release formulation of Creque et al. (1980) might be appropriate as a combination of the reservoir and rate-controlling membrane. Materials and Methods Experimental work was directed to the characterization of the prototype shown in Figure 1 for insulin delivery. Membrane. Hypo1 nonwicking hydrophilic polyurethane foam (W. R. Grace and Co., Lexington, Mass.) was used as the membrane in all experiments because of ita high compressibility and hydraulic permeability. The foaming mixture consisted of 100 parta by weight of FHP 3000 prepolymer, 1part of L-520 surfactant (Union Carbide, Toronto) and 70 parts of distilled water. The porosity of the dry unswollen foam was 0.933-0.940 and the foam absorbed approximately 22.5 g of water/g of dry foam. The Darcian permeability without lateral constraint of the foam and a t zero head was approximately 3 X lo4 cm2 (Sefton and Lusher, 1980). The 4.6 mm diameter rods of foam were drilled from liquid nitrogen frozen slabs of cured foam. Only disks of foam with an air permeability of 5 f 1 X cm2at an air flow rate of 850 cm3/min were used in subsequent experiments. Foam lengths ranged from 0.9 to 1.1cm.No specific conditioning of the foam was done nor was any attempt made to remove extractable from the cured foam. No additional ratecontrolling membrane was used in t h e e experiments, except where noted.
Figure 2. Falling head permeameter.
Micropump. The prototype which has received the greatest attention consisted of a short length of 3/16-in.i.d. Lucite tubing capped at one end by a in. thick disk of Med 12 Porex porous polyethylene (Glssmk, Fairbum, Ga); 2500 turns of 34 gauge Teflon-mated wire were wound around the Lucite tube and then enclosed in a mild steel housing. Coil resistance was 100 R. The piston was a 4.57 mm diameter, 2.5 em long Lucite rod glued to an identically sized polyester coated steel rod. Power Supply. For augmented delivery, a variable voltage dc supply was produced by rectifying and smoothing the output from a powerstat (Superior Electric Co., Bristol, Conn.). A “555” timer was used to provide an interrupted dc current for the solenoid with an “on” time of 5 ms a t a variable frequency. Except where otherwise noted, the frequency was 30 strokeslmin. Flow Rate Measurement. The rate of decrease of insulin solution in a graduated 0.5-mL pipet (cross-sectional area = 0.323 mm2)acting as the feed reservoir was used as a ‘Talling head permeameter” (Figure 2) to measure both the basal and augmented flow rates. The slope of a semilogarithmic plot of pressure head against time is directly proportional to the permeability at that pressure head. The volumetric flow rate a t any head is equal to the product of the pressure and this permeability. A linear plot indicated that the pump permeability is constant and Darcy’s law is observed. Correction must be made for capillarity; the capillary height relative to the level in the downstream reservoir is subtracted from the measured height to get the true head. No detectable difference in permeability was found using this method relative to that using radio-labelled insulin and a constant head feed reservoir. The tubing and valves accounted for about 15% of the total resistance of the pump in both methods. The downstream reservoir liquid level was maintained constant through the use of an overflow tube. The feed reservoir consisted of 0.4 U/mL mixed bovine/porcine insulin (Toronto insulin, Connaught Laboratories Ltd., Toronto, Canada) in 0.05 M phosphate buffered saline (pH 7.4), containing 1% formaldehyde. Reproducibility. An accelerated use sequence consisting of 2 h of basal delivery followed by 2 h of augmented operation at a pressure drop of 9.3 X lo2 dyn/cm2 and a time averaged power of 340 mW (90V) was used to aswa the short term (25 cycles) reproducibility of the device. The cycle was interrupted periodically for 30 min to characterize the delivery using the falling head permeameter. Delivery Improvement. A 1%” diameter, 1-pmpore size polycarbonate membrane (Nucleopore Corp., Pleaeanton, Calif.) held in a Swinnex fdter chamber (Millipore Corp., Bedford, Mass.)was inserted in the delivery line
Ind. Eng. Chem. Prod. Res. Dev., Vol. 20, No. 1, 1981 3
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TIME AVERAGED POWER (mW)
Figure 3. Effect of pressure difference and time averaged power on augmentation (i.e., augmented/basal rate, Q/Qo).Lucite pump, 2500 turns, 30 strokes/min, foam length = 0.9 cm. Time averaged power = peak power X “on time” X frequency.
between the insulin reservoir and the micropump to lower the basal rate to practical levels. The effective membrane area was 0.7 cm2. The pump was similar to that described above except that it consisted of 2000 turns of 36 gauge wire (R = 80 0)wrapped around a 7-mm 0.d. glass tube, capped with a sintered glass disk. A tapered glass adapter was added to the outlet and the Lucite rod was 5 mm long. To measure the lower flow rates in this apparatus a length of 0.015 ID Intramedic tubing (PE 20, Clay Adams) was used instead of the pipet in the falling head permeameter. Results and Discussion Basal Delivery. Under basal delivery conditions, (i.e., no applied power), the semilogarithmicplot of feed solution height in the reservoir against time was linear over a pressure range of 1 X lo3to 15 X 103 dyn/cm2,indicating the validity of Darcy’s law for the description of basal flow. At lower liquid heights, the correction for the capillary height in the pipet introduces some error causing an apparent deviation from linearity. Basal permeability was 5 X cm2based on the area and length of the foam using the falling head permeameter. The difference between this value and that determined by steady-state measurements using radiolabelled insulin (1.5 X lo-’ cm2) was attributed to the presence of an axial cavity in the piston of the earlier more permeable pump (Sefton et al., 1979). Augmented Delivery. The effect of repeated compression of the foam is shown in Figures 3 and 4. Augmentation is defined as the ratio of augmented delivery rate (i.e., with repeated compression)to the basal delivery rate at zero power. It can be seen from Figure 3 that increasing the current to the solenoid (applied power) increased the force applied by the solenoid core to the membrane and increased the augmented delivery rate and the degree of augmentation. Increasing the pressure difference across the membrane caused an increase in the augmented rate but reduced the degree of augmentation at a given power input. At higher delivery rates, the compression resulting from a given power input had a relatively smaller effect, producing smaller degrees of augmentation.
Increasing the frequency of compression resulted in an increased augmentation, as expected (Figure 4); that is, increasing the number of strokes per minute increased the net volume flow per minute. However, the apparent stroke volume V,, calculated from (Sefton et al., 1979) Q = 80+ fv, where Q = augmented rate, Qo = basal rate, and f = frequency, was not independent of frequency: the stroke volume decreased from 1.5 pL at the normally used frequency of 30 strokes/min to 0.38 pL at a frequency of 1500 strokes/min. This suggests that at the higher frequencies, reexpansion of the foam became impaired and the foam was unable to “refill” with fluid from the reservoir before the next compression stroke. Evidence of decreased augmentations at high frequencies has been obtained from pure water delivery (Lusher, 1978). The reduced stroke volume is compensated for by the increased number of strokes to give an increased augmentation, until the stroke volume is too small to be offset by further increases in frequency. Control of both frequency and current gives great flexibility to the delivery control mechanism: one parameter can be used to account for patient variability, the other used for the purpose of physiological control. Unlike the delivery of amaranth by a similar glass pump (Sefton et al., 19791, the stroke volume decreased with increasing pressure drop (or basal rate) at a given frequency and applied power. Also stroke volumes were approximately an order of magnitude smaller at comparable basal rates and peak powers, though covering a narrower range of basal rate than those reported earlier. These discrepancies underscore the inadequacy of the simple apparent stroke volume calculation. The apparent stroke volume represents the net delivered volume per stroke and is the difference between the total stroke volume and the volume per stroke of retrograde flow which is assumed to occur during each compression. The situation is clearly more complex than this, since the pumping efficiency (the fraction of fluid delivered downstream relative to the total stroke volume) is not necessarily constant and the basal rate is lowered during compression because of the reduced porosity. Furthermore, compression does not simply cause a quantity of fluid to be “squeezed out from the membrane as would be implied by this simple calculation. A more likely partial explanation for the mechanism of augmentation may be the change in resistance past the solenoid piston during rapid movement relative to that when it is stationary (basal flow). Rapid movement causes a reduction in the form drag as the fluid enters and exits
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Ind. Eng. Chem. Prod. Res. Dev., Vol. 20, No. 1. 1981 20
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Figure 5. Short-term reproducibility of the controlled-release micropump, as indicated by the cycle-to-cycle variation of basal and augmented flow rates at a pressure difference of 10s dyn/cmz and an average power of 340 mW (90V). Each cycle consisted of 2 h of basal operation followed by 2 h of augmented operation at the -e pressure and power. Same pump as in Figure 3. The standard deviations of the mean rates are also indicated.
the annular space about the piston and changes the pressure profile through the pump. This model of flow enhancement through rapid movement of the piston, predicta that augmentation should increase as the annulus size parameter, z (z = piston length/annulus thickness) decreases, which is consistent with the differences in augmentation between the earlier glass prototype and the current Lucite device. Furthermore, this model implicates the relative permeability between foam membrane and porous support as an important factor controlling augmentation. Although experimental support for this hypothesis is available (Treen, 1979) further work is required to account for all of the experimental results. Reproducibility. Figure 5 shows the effect of cyclic operation on the delivery characteristics of the pump over 25 4-h cycles. During this period there was no decay of pump performance. The cycle-to-cycle variation indicated in Figure 5 for basal (*6.4%) and augmented flow rate (*7.9%) was not significantly different, at 99% confidence, from the variation in flow rate measurement within a cycle (both F-statistics less than 3.46). This latter variation was attributed to a random measurement error, quantified over a 10-min period when no change in real permeability could be presumed. A larger variation in flow rate was apparent over 1000 h of continuous augmented delivery but this appeared related to the need to replace insulin reservoirs on a regular basis (Bums,1980); this larger variation was not apparent in the ratio between augmented and basal flow rates (i.e., augmentation), suggesting an approximately constant effect of piston movement. Delivery Improvement. The volumetric flow rates of Figures 3 and 4 are too high for practical application of the current prototype in insulin delivery (the corresponding insulin concentration would be too low, making the drug reservoir very large). Furthermore, the augmentations are insufficient to meet the needs of the artificial pancreas. More recent prototypes incorporate a thin rate-controlling membrane in series with the Hypo1 foam. Not only was the basal flow rate reduced dramatically, but the augmentation was much greater (Figure 6). Even though the stroke volumes were approximately an order of magnitude less than those in the foam-only device, the basal rate was lowered to a greater extent, resulting in greater augmentations. The utility of this membrane arrangement in making a practical device is further in-
Figure 7. Scanning electron micrograph of precipitates ou I-pm Nucleopore membrane collected after 2 h use. Unlike Figure 1 or the pump of Figure 6, this membrane was placed between the foam membrane and the porous support. Energy dispersive X-ray analysis of crystals indicated the absence of metallic elements except for traces of potasaium and silicon suggesting that the crystals may be primarily organic (e.g., insulin?).
dicated in Figure 6 since a t the desired basal rate of 0.2 mL/day (1.4 X lO-'mL/min), the augmentation a t maximum power is more than 10. At this basal rate a 25 mL, 500 U/mL reservoir could last a year depending on the insulin demand. Even higher augmentations have been obtained using other membranes (Cahill, 1980). Unfortunately, Nucleopore membranes would not be suitable for insulin delivery because of the extensive amount of insulin precipitates which form on these membranes during use and which ultimately block the pores of these membrane (Figure 7). The basal rates shown in Figure 6 were the initial rates; within 5 h these rates had dropped by a fador of 4. However, the permeability of hydrophilic membranes to insulin were measured at steady state (Sefton and Nishimura, 1980) without any reduction in transfer rate with time. This suggests that precipitation of insulii may he secondary to an adsorption process which is dependent on the surface energy of the underlying substrate. This precipitation process is characteristic of all such insulin delivery devices (Albisser, 1979) and currently limits the application of the artifical pancreas. Both
Ind. Eng. Chem. Prod. Res. Dev. 1981, 20, 5-12
the use of cuprophane membranes in the micropump and the insulin adsorption process are being studied in more detail. Conclusions A controlled release micropump without valves has been developed for the administration of insulin at variable rates. The device is small and consumes little power. The degree of augmentation depended on pressure drop, applied power, and the frequency of compression. The basal flow rate has been reduced and augmentation increased to practical levels using a rate-controlling membrane in series with the compressiblefoam membrane. The delivery characteristics were reproducible, with the short-term variation indistinguisable from measurement error. Although some modifications and improvements are required, the continued development of this device will enable biomedical researchers and diabetologists to investigate the relationship between metabolic homeostasis and the degenerative complications of diabetes. The potential for restoration of homeostasis and the prevention of diabetic sequelae will benefit the millions who suffer or will suffer from this disease. Acknowledgment The authors express their thanks to the J. P. Bickell Foundation and the Natural Sciences and Engineering Research Council of Canada for their support of this project. They also thank P. J. Cahill and A. To for their experimental assistance. Literature Cited Albisser, A. M., Leibel. B. S., Clln. Endocrlnol. M t a b . , 8(2), 457 (1977). Albisser, A. M., Hospital for Sick Chlldren, Toronto, Canada, personal communication, 1979.
5
Blackshear, P. J., Rohde, T. D., (fotteng, J. C., Dorman, F. D., P&hs, P. R., Varco, R. L., Buckweld, H., Dlebstes, 28, 834-639 (1979). Burns, K. J., M.A.Sc. l h d s , Depnrtmmt of chemlcel E m n g and A p p W Chemistry, Unlverdty of Toronto, 1980. Cahlll, 0.F., E t z w k , D. D., Frdnkel. N., DlekStcM, 26, 237 (1978). W, P. J., B.A.Sc. The&, -of Chemkal En@mhg and AppHed Chemistry, Unlverdty of Toronto, 1980. Canadlan DIabeUc Assodatkn, “Dkbac#, Stat)stlcs”, 1978. a w ,R., Folkmen, J., Dfektss, 29, 37-40 (1980). Creque, H. M., L (krlya, Y., Bahorlc, A., Marliss, E. B., Zinman, B., Alblsser, A. M., Dlebstas, 28, 558-584 (1979). Irslgler, K., K r k , H., Dlebetes, 28, 196-203 (1979). Kdendorf, K., Bolsen, J., Lorup, B., Dlabetkgh, 18, 141-145 (1980). Langer, R. S., R h b , W. D., Hsieh, D. S. T., Fdkman, J., paper presented at ths 179th Natbnal Meeting of the American Chmlcal Society, Dhrbion of Industrlal and Engbwerlng ChemtstF/, Houston, TX, Mar 23-28, 1980. Leeper, H. M., Buckles, R. Q., aumerd,G. V., Lorberbaum, M. A., Sevilla, E. R., Yum, S. I., paper presented at the 111th Meeting of the Rubber Mvc slon, American chemlcel Sodety chlcego, Ill,May 3-6, 1977. Lusher, H. M., MA.&. Thesis, Department of chemlcai Engineering and Applied Chemistry, W e r d t y of Toronto, 1978. Plckup, J. C., White, M. C., Keen, H., Parsons, J. A., Albertl, K. G. M. M., Lancet, 870-873 (Oct 27, 1979). SanUago, J. V., Clemens, Clarke, W. L., Klpnis, D. M., Dlebetes. 28, 71-84 (1979). Sefton, M. V., Lusher, H. M., Flrth, S. R., Waher, M. U., Ann. Biomed. Eng., 7, 303-317 (1979). Sefton, M. V., Ntehlmura, E., J. pharm. Scl., 89, 208-209 (1980). Sefton, M. V., Lusher, H. M., J. Appl. fb&m. Scl., In press, 19SO. Tamborlene, W. V., Sherwln, R. S., Qenel, M., FeUg, P., NewEngl. J . A M . , SOW1 l), 573-578 (1979). Tchobroutsky, G.,Dlabstbgk, 16, 143-152 (1978). Treen, M. E., B.A.Sc. Thesis, Department of Ghemlcal Englneerlng and A p plled Chemistry, Unlverslty of Toronto, 1979.
Received for review May 16, 1980 Accepted June 24, 1980
Presented at the 179th National Meeting of the American Chemical Society, Houston, Texas, Mar 23-28,1980, in the Division of Polymer Chemistry Symposium,Advancea in Polymeric Controlled-Release Formulations.
Controlled Release of Polymeric Organometal Toxicants Max Kronsteln Chemistty Department, Manhattan &/&e,
Riverdale, New Yo&
10471
The mechanism of antifouling paints consists of the interreaction of the Inorganic toxic metal oxide pigments, or of the organometallic toxicants, with the low-pdymeric sweller fraction in polymeric vehicles. The reaction product, in the applied coatlng, migrates gradually through the higher-polymeric matrix to the surface, there providing antifouling protection. In under-water exposure, the reaction product from there enters the water. Organometalik toxicants give a lower release rate than the metal oxide toxicants. The organic components in the released reaction products are studied by their infrared spectra; the metals, by atomic absorptlon spectroscopy.
The paper is concerned with the mechanism of polymeric antifouling paints, in particular those with organometal toxicants. The polymers used in their formulation consist of more than one fraction, one being higherpolymerized and insoluble but dispersible in organic solvents; the other, lower polymerized, is still fluid and soluble, and still reactive with other reactive materials. In the applied polymeric antifouling paints the high-polymer fraction represents the matrix of the applied protective layer and the lower polymeric fraction represents a sweller phase which is capable of migrating through the polymer matrix, coming gradually to the surface of the protective layer and from there entering gradually the immersion water. 0 196-432 118 1I1 220-OOO58O 1.OOlO
The organometal toxicants are capable of reacting with the low polymer fraction and become an actual component of the sweller fraction and of the polymeric material as a whole. When this fraction migrates to the coating surface, the toxicant groupings offer a protective effect against the growth of fouling matter on the surface, and, because these toxicants are not independent enclosures but actual components which can be released by migration through the matrix, their release is a gradual one, a “controlled” one, and a limited one, before they enter the immersion waters of the environment. Before discussingthis mechanism and ita results in more detail and before discussing the methods of measuring the rate of gradual release and the total amount of the released 0 1981 American Chemical Society
