Closed-loop delivery systems for insulin therapy
Insulin replacement therapy is one of the most debated issues in modern diabetology. The aim of the therapy is twofold: first, to prevent acute compli-cations such as ketoacidosis, and second, to prevent long-term micro- and macro-vascular complications.
Prevention of acute complications has been readily achievable since the introduction of insulin in 1922, as demonstrated by the huge reduction of the number of deaths due to diabetic ketoacidosis during the following decades (1). Unfortunately the second, name-ly the prevention of micro- and macro-vascular com-plications, has not been achieved yet.
The natural history of insulin-dependent diabetes (100M) is associated, despite insulin therapy, with a progressive onset of microangiopathic complications leading to retinopathy, nephropathy, and accelerated atherosclerosis. More than 10 years ago, Pirart had already suggested that not only the duration of diabetes but also the degree of hyperglycaemia was responsible for the frequency and the severity of retinopathy (2). Now a number of prospective studies have been undertaken to investigate the effects of prolonged near-normoglycaemia on the development or the reversal of diabetic complications. To this end it is worth considering the results of two trials which are already concluded, the Steno and Oslo studies (3, 4). These studies, even though conducted on a limited number of patients, demonstrated a lower incidence of retinopathy in patients on intensified insulin therapy. This observation justifies the applica-tion of an intensified insulin therapeutic regimen in young type 1 diabetics of recent onset, to obtain a glycaemic control close to the physiological model.
Insulin therapy differs from other forms of hormone replacement since not only the total dose but also the timing and the means of administration are
crucial to fulfil the aim of the therapy. Intensified insulin replacement can be accomplished nowadays using two therapeutic options: conventional insulin therapy based on multiple injections (regular insulin is administered before each meal and intermediate insulin is given at bedtime); and, alternatively, contin-uous subcutaneous insulin infusion by means of external pumps, superimposing preprandial boluses (similar to the preprandial injection mentioned earlier) to a continuous basal infusion, often at variable rates, thus avoiding the administration of intermediate insulin (5-8). The critical point of these therapeutic options lies in the fact that insulin delivery is programmed externally on the basis of frequent and painful capil-lary blood glucose self-monitoring using the finger-prick technique (9, 10). The outcome of these thera-peutic approaches can be considered satisfactory in terms of mean daily blood glucose and glycosylated haemoglobin, but they require constant medical su-pervision and optimal patient compliance (11,12). Furthermore, the physiological model of blood glucose regulation cannot be entirely reached.
Physiological blood glucose regulation can only be attained if insulin delivery is modulated by glucose concentration itself via a feedback mechanism as it exists in the healthy pancreas. Therefore what is needed is a closed-loop system, or artificial pancreas, capable of restoring a physiological model of insulin release.
The artificial pancreas
blood glucose concentration is maintained within a narrow range by appropriately increasing plasma insulin levels after meals and by decreasing insulin delivery between meals and during the night. As a result of the ability to increase and decrease insulin delivery both hyper- and hypoglycaemia are prevented. On the other hand, in diabetic patients blood glucose appeared to fluctuate chaotically and unrelated to circulating insulin levels (13). These insights empha-sised the compelling necessity for a more rational insulin therapy. Thus, many researchers began to consider the possibility of recreating the biological feedback loop between blood glucose concentration and insulin release by means of an artificial system. The first attempt to control blood glucose levels on the basis of continuous monitoring by means of a "servomechanism" was described by Kadish in 1964 (14). However, the first computer controlled artificial pancreas was developed in the early 70's at the Hospital for Sick Children in Toronto, Canada, by a team led by Michael Albisser. The first report on its clinical application was given to the medical audience at the annual meeting of the American Diabetes Association in May 1973 (15). The system consisted of three basic components: a continuous flow glucose analyser (a modified Technicon Autoanalyzer), a com-puter controller implemented with algorithms for in-sulin delivery, and a pumping mechanism. These components were connected in such a way as to form a loop around the patient. Blood was continu-ously withdrawn from a peripheral vein by means of an indwelling double lumen cannula, diluted with an anticoagulant (heparin), and conveyed to the glucose analyser. The electrical signal arising from the analy-ser, after being converted into digital format, was delivered to the computer, and interpreted to calculate the appropriate insulin amount to be infused into another vein by the pumping mechanism. The enthu-siastic reaction to the prototype prompted world-wide investigations and a plenitude of bedside apparatus were developed in many laboratories during the following years (16-19).
However, only two commercial versions were made available. The Biostator GCIIS, a cumbersome bedside device weighing over 60 Kg, and the more recent Betalike, a desk-top device weighing 19 Kg (20, 21). The operating principles of these closed-loop insulin delivery systems are virtually unchanged with respect
Fig. 1 - Schematic representation of the glucose sensor used in currently available closed-loop insulin delivery systems. Glucose diffuses through an outer porous membrane into the enzyme layer where the enzymatic reaction catalysed by glucose oxidase (GOD) occurs. The reaction product hydrogen peroxide (HP2) in turn diffuses through an inner permselective membrane and is oxidised on the surface of the platinum electrode poised at suitable potential (+ 650 mV versus AgIAgCI).
to the Toronto prototype (only minor changes to the mathematical formulation of the algorithms) however they both rely on a glucose sensor rather than the cumbersome continuous flow analyser (22, 23). The glucose sensor consists of an enzymatic membrane (the enzyme is immobilised between two membranes, the outer is made of polycarbonate and the inner of cellulose acetate) assembled on top of an ampero-metric electrode (24). The basic operating principle of this enzyme electrode (25) is easily understood (Fig. 1). Glucose diffuses through an outer porous membrane into the enzyme layer where the enzymatic reaction catalysed by glucose oxidase (GOD) occurs. The reaction product hydrogen peroxide (H202) in turn diffuses through an inner permselective mem-brane and is oxidised on the surface of the platinum
TABLE I - FACTORS STILL LIMITING THE DEVEL-OPMENT OF AN IMPLANTABLE CLOSED-LOOP INSULIN DELIVERY SYSTEM
• Implantable glucose sensor • Control algorithms
electrode poised at suitable potential (+650 mV versus Ag/AgCI). The resulting electric current is directly proportional to the amount of glucose present in the solution. The sensor has a linear range of 0 to 30 mmol.r' and a response time of less than one minute for a change of 11 rnrnol.t' in glucose concentration. During the past decade, bedside closed-loop insulin infusion systems have been exploited for both clinical and research applications (26). However, they cannot be considered as routine therapeutic tools. To this end, an innovative change is represented by the so called "wearable" artificial pancreas. This apparatus weighs only 400 g and was developed by Shichiri and colleagues in the early 80's (27). This portable system represents a step towards the achievement of the ultimate goal: the implantable artifical pancreas. Though, several problems remain to be solved before a safe and effective device will be available world-wide. Listed in Table I are the major issues that are still without a definitive answer. The first, and more substantial problem, is related to the unavailability of an implantable glucose sensor.
Implantable glucose sensor
A chemical sensor might be viewed as a device incorporating a sensing element either intimately con-nected or integrated within a transducer (Fig. 2). The usual aim is to obtain a digital electronic signal which is proportional to the concentration of a specific chemical (28). Sensors, or biosensors when molecular recognition is carried out by a biological sensing element such as an enzyme, may be classified according to the basic physics of the transducer (29). Therefore, four main categories can be defined:
a) Electrochemical, based on organic or inorganic catalysts exploiting either amperometric or poten-tiometric devices.
b) Optical, relying on reversible or irreversible chemi-luminescence detected by means of optical fibres. c) Piezo-electric, based on the vibrational properties
of some crystals.
d) Calorimetric, based on heat detection by means of a thermistor.
Theoretically, all the previously mentioned tech-niques, may be used for glucose sensors assemblage (30-33). However, most of them are pure research matter that, even though under active investigation,
Fig. 2 - Diagram illustrating the operating principle of an enzyme-based electrochemical biosensor (see text for further detail).
Fig. 3 - Diagram of a 0.25 mm needle-type glucose sensor (from Ref.36,modified).
accomp-Fe
Fig. 4 -Ferrocene structure (bis-[n5-cyclopentadienyl] iron).
lished by means of outer membrane pre-treatment with organosilanes allowing for a limited glucose diffusion without interfering with oxygen diffusion, thereby increasing the apparent Km of the enzyme and thus reducing oxygen dependency (36, 37): the second procedure used to overcome oxygen depen-dency deals with 2nd generation amperometric
bio-sensors where oxygen independence is achieved by including within the enzyme layer a reaction mediator for electron transfer thereby excluding oxygen from the reaction cascade (38). A well known and widely used mediator is ferrocene (bis-jn't-cyclopentadienyl] iron) and its derivatives. Ferrocene is a transition metal rr-arene complex which consists of an iron atom sandwiched between two cyclopentadienyl rings (Fig. 4). This class of biosensors also has other advantages over first generation devices (39), in particular: the low potential needed to reduce ferro-cene (220 mV vs Ag/AgCI) tends to minimise interfer-ence from other electroactive species; because the electron transfer between the mediator and the re-duced enzyme is fast, the response time of these biosensors is lower than first generation devices. Ferrocene-mediated glucose biosensors normally ex-hibit a linear current response over the range of glucose concentrations commonly found in diabetics (1-30 rnrnol.r") and needle-shaped devices have been successfully implanted in humans for continuous subcutaneous glucose monitoring during glucose in-fusion (40).
Further advantages in terms of response time and stability might come from the 3rd generation
arnpero-metric sensors (41). These devices provide a simple method to determine glucose concentration. They allow the direct oxidation of the reduced enzyme at the surface of an electrode based on conducting organic salts. Suitable materials for such modified electrodes are 7,7,8,8,-tetracyano-p-q uinod imethane (TCNQ) as electron acceptor, and N-methylphenazi-nium (NMP) as electron donor. Albery and colleagues using a third generation electrode based on TCNQ and NMP obtained a remarkable long-term stability: 28 days of continuous operation with only 20% loss of response to glucose (42).
The electrochemical aspect of biosensor technology represents a major challenge for scientists, however, other fundamental issues still limit a full in vivo application of these devices.
A crucial question is related to the selection of an adequate implant site (43), in other words: where should the sensor be placed to provide a measure-ment of glucose concentration likely to be used to control insulin delivery?
Given the needle shape of current sensors the most likely site seems to be the subcutaneous tissue since intravascular or intramuscular implants appear to be too hazardous and painful. Both under steady and non-steady state conditions, interstitial glucose con-centration reflects that of venous blood, even though a rapid change from euglycaemia to hyperglycaemia is associated with a significant time delay (44,45).
Besides, the choice of the subcutaneous tissue is also related to the minimal tissue response evoked by the wire or filament shape (46), and the relatively short operational life of these sensors (currently ranging from less than one up to three days) such that they require frequent replacement. The causes of the poor long-term stability are associated with:
a) the electrode materials b) the enzyme
c) the often inadequate biocompatibility.
More reproduciible electrode assemblage and con-struction techniques could contribute to improve the long-term performances (47).
TABLE 11- BIOSENSORS STERILISATION METHODS: ADVANTAGES AND DISADVANTAGES
Sterilant Ethylene oxide Propylene oxide
Ionizing radiation (gamma) (From Ref. 53, modified)
Sterilisation time 6-8 hours days hours Safety Toxic, flammable Non-toxic Hazardous Availability Hospitals Chemical vendors Special service companies
However, a primary factor responsible for the im-pairment of sensor performances, is related to the deposition of proteinaceous material inside the outer membrane pores which takes place soon after sensor implantation (49). These proteins are precursors of the latter arrival of polymorphonuclear leucocytes and macrophages at the implant site (50). These events cause a severe restriction to substrate diffusion which results in a sluggish and decaying sensor response. To improve biocompatibility a number of attempts have been made to produce new synthetic membranes using several types of polymers, modifi-cation of surface physical properties, and chemical derivatization. Polyurethanes have been widely used for membrane construction in view of their relatively good biocompatibility. Other potentially suitable mate-rials for biocompatible membrane are polytetrafluo-roethylene (PTFE or Teflon) and polyalginate (51). It is also worth remembering that an implant might not have just local effects but also systemic ones. For instance, toxic compounds released from the enzyme matrix might cause liver damage in the long-term. Heterologous proteins, e.g. glucose oxidase, are im-munogenic and an accidental leakage can cause hypersensitivity reactions. Unfortunately, the systemic effects of sensor implantation have so far received little attention and these vital questions lack definitive answers (52).
Finally, when placing a sensor into the body one must be sure that the patient is protected from the introduction of infectious agents, i.e, the device must be sterile. Sterilisation procedures should provide full patient safety and not compromise the integrity of the sensor components. Several procedures, using gaseous and liquid chemicals as well as ionizing radiation, have been employed for biosensor
sterilisa-tion. The advantages and disadvantages of the most widely used methods are summarised in Table II (53).
Control algorithms
The development of robust and effective algorithms is another problem to be tackled to improve the capabilities of a closed-loop insulin delivery system. Over the past 20 years several algorithms have been proposed and used for closed-loop blood glucose regulation. Irrespective of the strict mathematical formulation and the addition of pseudo-constants they are all proportional-derivative algorithms. The proportional element, also referred to as static in some formulations, is usually related to the actual and/or "projected" glucose concentration. The deriva-tive element, also referred to as "dynamic", is related to rate of change of glucose concentration. The algebraic sum of these two elements allows an insulin infusion profile in response to a glucose challenge mimicking the physiological biphasic pattern of insulin release (54). All the proposed algorithms perform similarly in terms of blood glucose control, however none of them can restore a normal blood glucose control (55, 56). Furthermore, closed-loop insulin delivery regulated according to these algo-rithms usually resulted in overestimation of insulin requirements and consequent hyperinsulinaemia. The causes of hyperinsulinaemia during closed-loop insulin infusion depend upon;
a) the peripheral route of infusion
b) the absence of anticipatory influences on insulin delivery
c) the time delay associated with blood glucose measurement
Cobelli and co-workers demonstrated with an ele-gant simulation study that peripheral insulin infusion inevitably resulted in elevated circulating insulin levels during closed-loop therapy (57). Whilst the same algorithms could restore a near-normal insulin profile when intraportal insulin delivery was simulated. How-ever, intraportal insulin delivery is scarcely feasible, therefore, in an attempt to identify other causes responsible for insulin requirement overestimation we focused on the fact that insulin delivery from these closed-loop systems was merely regulated by blood glucose dynamics, while the physiological response of human pancreas beta-cells is anticipated because of the influence of gastroenteric hormones and auto-nomic nervous system. Hence, we added to the control loop a feed-forward element simulating the anticipated response of the beta-cells by means of a pre-programmed insulin infusion. At the beginning of the meal a two-step pre-programmed insulin infusion lasting 24 min was begun. The total amount of insulin infused during this phase was calculated to equal 20% of the patient's insulin requirement with respect to the meal. At the end of the pre-pro-grammed infusion the conventional feedback insulin delivery was restored. Pre-programmed infusion com-pared to conventional feedback alone allowed a significant reduction of the total insulin dose and a better glycaemic control (58).
Closed-loop Insulin delivery performed with feed-back controlled devices is associated with a measure-ment delay, mainly caused by fluid transportation, that might be responsible for insulin requirement overestimation and controller instability (59). To in-vestigate the effects of time delay on insulin infusion rate, insulin dose and blood glucose behaviour, and the effectiveness of a time delay compensator (Smith's predictor) (60) we undertook a simulation study. The program was written in Simnon language and con-sisted of 3 main components: a continuous system (patient model), a discrete system (digital controller), and a connecting system allowing for sampling and subsystem interconnections. A validated control sys-tem model was used to simulate the patient (61, 62) and a proportional-derivate algorithm to control insulin infusion (20). Simulated Intravenous Glucose Toler-ance Test (IVGTT [0.33 g/kg]), 60 min intravenous glucose infusion (10 mg/kg/min), and Oral Glucose Tolerance Test (OGTT [75 g]) were used as
perturba-tion schemes. Simulaperturba-tions were performed without delay, with 3 and 7 min delay, and with time delay compensation. During the IVGTT time delay, up to 7 min, scarcely influenced insulin infusion rate, total insulin dose, and blood glucose. During intravenous glucose infusion no major differences were observed when a 3 min delay was simulated, but a 7 min delay resulted in large fluctuations of both blood glucose and plasma insulin. Significant delay related incre-ments of the insulin dose were not observed, but time delay compensation reduced the amplitude of plasma insulin oscillations. During the simulated OGTT, a 7 min delay resulted in higher peak value of plasma insulin than without delay, and the total delivered insulin increased by 8.5%. Time delay com-pensation was effective in reducing both peak plasma value and total insulin dose, while no differences were observed in blood glucose with respect to simulations performed without time delay (63).
Fig. 5 - Block diagram ofaModel-Reference Adaptive Controller (conventional nomenclature and symbols have been used; from Ref.65,modified).
Insulin pumps and infusion route
Several types of pumps have been developed and used for insulin delivery during the past decade, namely, syringe pumps, peristaltic or roller pumps, and bellows pumps (71). All of these devices have known thoroughly, as in the case of blood glucose regulation, simple adaptation mechanisms do not perform satisfactorily and self-adaptive controllers must be used because they can perform on-line parameters (while the system is operating) estimation and adjustment (65). A typical example of Model-Reference based Adaptive Controller (MRAC) in shown in Figure 5. By means of an outer loop added to the conventional control loop (inner loop) a suitable reference model is used to evaluate the ideal response of the controller which is in turn compared with the actual output of the system. The difference (error Em) is then used to adjust the parameters of the controller. This type of controller works well and encouraging results have been obtained in short-term animal eximents (66). However, little is known about the per-formances of such controllers in the long-term, thus it is difficult to anticipate whether or not this control approach will do any better than conventional feed-back algorithms (67). Computer simulation (68) might be expremely helpful to evaluate "non-invasively" the long-term performances of new algorithms (69, 70), even though the definitive answer will only come from extensivein-vivo studies.
reached a sufficient miniturization and sophisticated electronics allows optimal operational control.
They probably represent the most advanced com-ponent of either a portable or implantable artificial pancreas. Devices featuring fixed or variable rates have been designed for both external and implantable pumps. The pumping mechanism is usually controlled by a microprocessor and remote control units have also been developed (72). The advantages of a remote controlled pump are immense. For instance, this will allow the use of external sensors (such as s.c. needle) to control an implanted pump via a telemetric system.
Portable pumps are usually carried attached to the belt or to another harness and deliver insulin, through small cannulae and needles, either into the subcu-taneous tissue or intravenously. A portable, or weara-ble device exploiting subcutaneous insulin delivery might appear feasible in a reasonable time, however this infusion route is scarcely suited for closed-loop controlled systems. Insulin absorption kinetics from the subcutaneous tissue is heavily influenced by a number of variables (73) (local degradation, skin temperature, blood flow, etc.) that might preclude feedback controlled delivery. The intravenous route is certainly superior, more effective than the subcutane-ous and has been already used by Shichiri and colleagues for their "wearable artifical pancreas" (74). The efficiency of this infusion route is well established and documented in a large number of diabetic pa-tients. However, blood is an extremely hostile envir-onment and intravenous delivery is frequently asso-ciated with catheter obstruction by blood clotting, and elevated thrombotic risk (75). On the other hand, an extracorporeal pump is exempt from biocompati-bility problems, whilst implantable devices must be biocompatible. All of the components of an implanta-ble pump, with the exception of the delivery catheters, are usually housed in a highly biocompatible case. The preferential site for implantable pumps is the subcutaneous tissue of the lower abdominal wall, and, through special catheters, insulin is delivered intraperitoneally. The pump is well tolerated and the local reaction is usually minimal. However, to improve both biocompatibility and cosmetic results (the di-mensions of a current implantable pump are those of a first generation pace-maker) smaller and lighter pumps should be developed (76). When it comes to
inplanted insulin delivery systems other issues deserve further attention. Current implantable pumps deliver insulin into the peritoneal cavity. This infusion route offers several advantages over the subcutaneous and intravenous routes, but additional studies are needed to clarify the pharmacokinetics of intraperitoneally delivered insulin (77). Another major problem to be tackled is related to delivery catheter occlusion. This represents the prevailing cause of pump explant as recently reported by the International Study Group on Implantable Insulin Delivery Devices (78). Catheter occlusion depends on poor biocompatibility, with consequent fibrin deposition and tissue growth, the imperfect geometry, and insulin deposition inside the lumen.
Insulin stability
Insulin is an inherently fragile protein that exhibits a marked tendency to polymerise and to form macro-molecular aggregates (79, 80) causing serious prob-lems to hormone delivery (81). We can distinguish between the formation of macromolecular aggregates due to the association of insulin hexamers, and the association of multiple insulin fibrils with macromole-cules arranged along the central axis of the fibril (82, 83). Insulin fibrils in an aqueous medium are ex-tremely stable and biologically inactive. Furthermore, the infusion of altered insulin has been associated with raised serum amyloid A in insulin-dependent diabetics treated with infusion pumps (84). Both the tendency towards aggregation and that towards fibril formation are exacerbated by movement, temperature and contact materials (85, 86). To date there is not a definitive answer to the problem of insulin aggrega-tion, though, encouraging results have been obtained both with low pH insulin preparations and with the addition to insulin preparations of glycerol or surface-active substances (87).
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CONCLUSIONS
In the early 70's the implantable artificial pancreas was considered feasible in a few years. Regretably it is not yet available. Bedside and even portable feed-back controlled devices are in use, and the pumping part of the system has been implanted in humans successfully. Nevertheless, before a complete and safe device can be developed, either portable or implantable, the problems related to glucose sensors, control algorithms, infusion pumps, route and cathe-ters, and insulin preparations must be settled.
ACKNOWLEDGEMENTS
The editorial assistance of Mrs. Georgina Hoddle Stoppini is gratefUlly acknowledged.
P. BRUNETTI, M.MASSI BENEDETTI,
G.
CALABRESE, G.P. REBOLDI1st. di Patologia Speciale Medica e Metodologia Clinica, Universita di Perugia, Perugia - Italy
Reprint requests to: Prof. P. Brunetti,
Istituto di Patologia Speciale Medica Universita di Perugia
Via E. dal Pozzo 06100 Perugia, Italy
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