Clinical Need for New Biohybrid Pulmonary Assist Devices
Current approaches to treating respiratory failure have not significantly improved survival or quality of life of patients with severe pulmonary disorders, such as acute respiratory distress syndrome (ARDS) and late stages of chronic obstructive pulmonary disease (COPD).1 Despite continual progresses in supportive care, the mortality rate for ARDS is indeed still in the range of 40%. Chronic obstructive pulmonary disease, a pathologic condition with prevalence as high as 6–8% in the general population, is becoming the third leading cause of death worldwide.2 For acute critical respiratory illnesses, like COPD, ARDS, but also for chronic hypoventilatory respiratory failure, or fibrosis, the only options are lung transplantation or some forms of supportive mechanical ventilation. However, the number of potential recipients outpaces by far that of transplantable donor organs. Globally there are an estimated 700,000 patients who are in critical need of and could potentially benefit from a lung transplant. Yet, less than 6,000 procedures are done each year worldwide because of the limited availability of quality lungs. That means less than 1% of patients are being helped and of those listed for transplant, 20% will die waiting for a suitable donor organ (Tom Waddell, personal communication, April 21, 2017). There is an urgent need for therapeutic interventions for such critically sick patient where insufficient pulmonary function can be remedied by using (bio)artificial extracorporeal respiratory support devices different from conventional mechanical ventilation.
Extracorporeal technologies that supplement mechanical ventilation were first tested in the late 1970s with the aim to minimize the need for continuous positive-pressure ventilation and mitigate the subsequent ventilator-associated lung injury in severe hypoxemic respiratory failure.3 Early success with the technique prompted a randomized trial of extracorporeal membrane oxygenation (ECMO) in ARDS that demonstrated no survival benefit over standard-of-care ventilator management.4 This was, at least in part, attributed to complications from the device, including hemorrhage and thrombosis.
Technological innovations, such as using smaller-caliber double-lumen cannulae with singular venous access and centrifugal pumps, reducing blood flow rates, and improving hemocompatibility of the biomaterials used, have significantly raised the safety profile of the newer devices in patients with hypercapnic ventilatory failure.5 Recently, devices for extracorporeal CO2 removal have been used in both arteriovenous (AV) and venovenous (VV) attachment modes at blood flow rates that are lower than for ECMO. The goals for using these devices are to 1) prevent intubation in acute exacerbated COPD, when noninvasive ventilation is not tolerated or fails,1 2) discontinue invasive mechanical ventilation, or 3) enable ventilation with low tidal volumes and inflation pressure for patients with milder forms of acute lung injury and ARDS.1 Lung-protective mechanical ventilation has improved survival compared with traditional ventilatory support.6 The beneficial effects of lung-protective mechanical ventilation can further be enhanced in concert with extracorporeal CO2 removal by respiratory support devices.7 Avoiding mechanical ventilation while providing extrapulmonary lung assist to unload the lungs and provide relief for respiratory muscles may open the door to chronic support and improvement in the patients’ quality of life.
The limitations of the standard treatment for patients with COPD and other chronic conditions characterized by hypercapnic ventilatory failure underline the unmet medical need for newer therapeutic approaches that will directly remove CO2 from the blood. CO2 dialysis will permit to supply patients with high oxygen flow without causing central ventilatory depression, thus avoiding common complications of mechanical ventilation, and allow the patients to more fully enjoy the benefits of physical exercise and pulmonary rehabilitation. To meet these goals, novel respiratory assist devices with an adequate safety profile are being developed. The long-term goal is to use such devices in both the intensive care setting and for long-term domiciliary therapy in patients with chronic ventilatory failure.8
Pulmonary Assist Devices
Artificial lungs can be divided into four groups: extracorporeal devices, paracorporeal wearable devices, implantable, and intravenous artificial lungs positioned in the vena cava.9 These diverse designs are currently in various stages of development and commercialization. In the past, ECMO applications were often based on the off-label use of heart-lung machine-derived devices. Recently, more advanced systems have been developed for extrapulmonary gas exchange in clinical situations in which lung protection needs to be maintained in spite of invasive ventilation. Several of these newer extracorporeal devices are currently being tested in clinical trials. A number of paracorporeal, i.e., wearable, devices are expected to become available for patients within the next few years.10–12 We consider bioengineering approaches toward optimizing such devices as an important stepping-stone toward providing long-term lung assist that may ultimately lead to implantable bioartificial lungs. The development of implantable or intravenous bioartificial lungs has recently been reviewed10–12 and is beyond the scope of this review.
Extracorporeal lung support devices have provided predictable gas exchange in patients with acute respiratory failure since 1972.13 In these devices, the hollow fibrous, blood-contacting gas exchange membranes are conventionally made of plasma-tight polymethylpentene (PMP), which is suitable for prolonged extracorporeal lung assist. Alternatively, some gas exchange membranes are made of polypropylene (PP) and are mostly used for short-term (up to 6 hours) applications, such as during cardiopulmonary bypass surgery. Both fiber types show excellent gas exchange characteristics, but the PP fibers exhibit several inherent drawbacks, such as plasma leakage.14,15 The primary goal in using such extracorporeal devices is to reduce intubation and mechanical ventilation while providing effective gas exchange. This kind of lung assist device (LAD) enables lung-protective ventilation with lower tidal volumes, lower respiratory rates and pressures and facilitates short-term therapeutic interventions, e.g., pulmonary resection in patients with severely compromised pulmonary function.16 Lung assist devices are also used for extended periods of time and may serve as bridge to palliation and recovery or to lung transplantation. Lung assist devices provide lifesaving or lung-protecting gas exchange by eliminating CO2 from the circulation, thus correcting respiratory acidosis.17 Depending on the clinical presentation (e.g., pulmonary hypertension versus COPD), pump-driven or pumpless extracorporeal systems are employed, in concert with different modes of cannulation and various types of cannulas. This versatility in the available hardware enables a personalized therapy tailored to the patients’ needs, providing both respiratory and circulatory support.18
Initially, LADs utilized pumpless AV cannulation. More recently, pump-driven VV minimally invasive low-flow CO2 removal devices are being employed.17,19 This development in pulmonary assist is analogous to the recent progress in acute renal dialysis therapies, which has progressed from continuous AV to continuous VV hemofiltration.20 In short-term cardiac applications, such as cardiopulmonary bypass, which last up to several hours, there is no need to replace the gas exchanger. By contrast, given that extracorporeal lung assist can last several weeks up to several months, necessitating periodic replacement of the gas exchange units during the treatment period. Among the potential issues with such protracted procedures are the long-term contact of the patients’ blood with diverse components of an extracorporeal LAD, such as gas exchange membranes, mechanical blood pumps, tubing, and cannulas, all of which may lead to hemolysis and clotting.21 Examples for current extracorporeal devices on the market are the CardioHelp system by Maquet (Rastatt, Germany) or the iLA (interventional lung assist) system by Novalung (Heilbronn, Germany).
Paracorporeal Wearable Devices: Initial Concepts
A key therapeutic goal of LADs is to obviate sedation and immobility, two of the main drivers of morbidity and mortality.22 A major advantage of current VV LADs is their potential for use in awake, spontaneously breathing patients.23 Increasing the mobility and enabling spontaneous breathing of patients on extracorporeal lung support will prevent atrophy of their respiratory muscles.24 Keeping patients awake and mobile during extracorporeal lung assist therapy results in significant physical and psychosocial improvements.25,26 In contrast to the advances in the design and development of durable implantable cardiac assist devices that serve as a bridge to heart transplantation27 and for destination therapy, no comparable LADs exist for chronic lung failure patients.
Currently, some early-stage biotechnology companies are developing their own versions of a bioartificial lung. For example, the initial concept of Biolung (MC3 Cardiopulmonary, Dexter, MI), a compliant thoracic artificial lung, entails a “soda can–sized” gas exchange device will be connected to the right ventricle of the heart.28 By contrast, the TandemLung (CardiacAssist/TandemLife, Pittsburgh, PA) kit contains a coaxial dual lumen cannula for VV bypass and a quick-prime membrane oxygenator and blood pump and has recently been approved by the Food and Drug Administration for use as a temporary cardiopulmonary support for up to 6 hours.
Some established manufacturers of extracorporeal devices are also focusing on the miniaturization and wearability of existing LADs and on extending the lifetime of the gas exchange units by reducing their thrombogenicity and avoiding fouling. In terms of enhancing patient mobility, promising strategies for the miniaturization of existing LADs and their reconfiguration as paracorporeal devices require a reduction in the overall weight and size of existing clinical systems. Examples of current projects aimed at redesigning existing LAD systems are the miniaturization of the iLA (Novalung, Heilbronn, Germany) or the Hemolung (A-Lung Technologies, Pittsburgh, PA).29 Miniaturization and total wearability require a design that will allow carrying the entire support system on the patients’ body without the need for bulky oxygen cylinders or compressors, which currently still keep the patient tethered to an ambulatory oxygen supply.29–31
Miniaturization of Extracorporeal Respiratory Assist Devices: The Challenges
The control systems and hardware components of current artificial lungs were originally designed for use in the operating room or the intensive care unit. They are far too large and heavy (>15 kg) to be wearable, particularly not by weakened patients. In miniaturizing current LADs, the major technical challenges include redesigning the permanent hardware components and reengineering the disposable parts of the gas exchange unit. A key step toward improving wearability is to identify an alternative sweep gas supply, which in the intensive care unit is usually provided by the hospital’s central oxygen supply or a mobile rack carrying oxygen tanks.29
Zhou et al.31 recently developed a wearable artificial pump lung, in which the controller, together with the battery source, is placed onto the patient’s body, whereas the oxygen source, usually a tank or oxygen concentrator, is integrated into a mobile rack. More recently, our team advanced the wearability concept of “i-lung,” a novel miniaturized integrated ambulatory lung device designed after the iLA, by incorporating all parts of this LAD into a dedicated apparel. The first prototype of the i-lung was recently tested in an animal study in pigs.32 As part of this innovative concept of integrative miniaturization, we used only a few components of the “iLA activve” without modifications. For instance, existing components and algorithms were utilized for the development of the central unit and the blood pump control. All other components and control systems were developed from scratch. For example, we developed a miniaturized suction pump, optimized for small size, weight, and noise levels, which provides the gas exchange unit with ambient air as sweep gas, thus avoiding the need for bulky external oxygen tanks. For this, the Novalung DP3 pump, which drives, inter alia, the “iLA activve” system, was radically reduced in weight by 65% and size by 50%. This integrated approach will allow for complete mobility and enhanced quality of life of the patient. Figure 1 shows an overview of our concept for downscaling and biolizing an existing ambulatory LAD.
The miniaturization of the gas exchanger itself is essential for restoring patient mobility. For example, Kopp et al.30 integrated a rotatory blood pump into the center of a membrane gas exchanger and thus reduced the amount of tubing and connectors in the system. More recently, Zhang et al.33 developed an ultra-compact, disposable pump-based lung device, by integrating a magnetically levitated impeller into the housing of a hollow fiber membrane. Miniaturization can be achieved by providing homogeneous blood flow fields, which incidentally will also result in enhanced efficiency of gas exchange. Any new design will have to be modeled mathematically to predict and optimize the gas exchange, while taking into consideration the ability to reduce blood coagulation or the requirement for anticoagulation therapy.34
Optimization of flow conditions can be achieved, at least in part, by physically modifying the gas exchange unit. For example, Schewe et al.35 modeled the effects of the blood inlet and outlet expansion angles on device impedance and blood flow pattern and showed experimentally that for blood flow rates up to 7 L/min arranging the inflow/outflow tracts at an angle of 45° resulted an ideal combination of low impedance and uniform blood flow patterns. Similarly, redesigning the shape of an existing clinical gas exchange device, such as the iLA, and the inlet and outflow tracks yielded a miniaturized device with a better gas exchange efficiency than the parent model, in spite of a significant reduction in gas flow.36 Most recently, a first limited side-by-side comparison of a commercially available LAD (iLA) with its prototype miniaturized ambulatory version (i-lung) in a pig model indicated no discernable differences in gas exchange efficiency between the two devices over a 72 hour period.32
Beyond State of the Art: Microfluidic Approaches for Artificial Lungs of the Future
Advancing beyond the current ideas for miniaturizing wearable LADs, some innovative, yet futuristic concepts focus on microfabricating small-scale, bio-inspired, bioartificial lungs. Using Rapid Prototyping technologies, e.g., soft lithography, microfluidic gas exchange devices have been manufactured from elastomeric, gas-permeable synthetic polymers, such as polydimethylsiloxane.37,38 These microfluidic devices generally comprise two communicating compartments—one accommodating the blood flow, and the other harboring gas exchange channels.39
In general, microfluidic approaches have distinct technological advantages: the biomimetic channel design provides improved fluid dynamic conditions in comparison to devices based on hollow fibers. For example, the design by Kovach et al.40 maximizes gas exchange rates by mimicking the natural blood-vessel architecture in a complex chip the size of a credit card. Scaled-up versions of multilayers of microfluidic devices, based on “flat gas exchange membranes,” are being considered as alternatives to more conventional hollow fiber gas exchange units.41
While ingenious as a concept, scale-up to clinically relevant levels of blood flow might require stacking of several thousand of these microchips, as well as careful consideration of the biocompatibility.38 Further research is required to translate these innovative technologies into realistic clinical applications.
Interactions of Surfaces with Blood: Antithrombogenic Modifications of Artificial Surfaces
Long-term use of wearable artificial lungs critically depends on the improved biocompatibility of their blood-contacting surfaces. In addition to patient-specific complications and comorbidities, such as hypercoagulability or sepsis, which modulate overall clot formation on an individual, case-by-case basis, device-induced thrombogenicity is governed by overall hemodynamic conditions in the device and the specific properties of the blood-contacting surfaces.42 Blood responses to artificial surfaces include activation of the intrinsic coagulation cascade, which can be initiated by the contact of plasma-contained soluble factors with negatively charged surfaces, or by induction of the extrinsic cascade, i.e., the upregulation of tissue factor on damaged endothelial cells (ECs) at the site of vascular injury and also on activated platelets. Any activation of the coagulation pathways finally results in the conversion of fibrinogen into fibrin and formation of a blood clot.43
Strategies to Modify Artificial Surfaces
Antithrombogenic surface modifications that increase the hemocompatibility of blood-contacting devices have been repeatedly reviewed.44,45 In the past 25 years, three such strategies have been pursued (Figure 2): surface passivation, biochemical surface functionalization, and coating the blood-contacting surfaces with a monolayer of functional (quiescent) ECs (endothelialization).
Surface Passivation of Blood-Contacting Surfaces
Surface passivation aims at minimizing the interactions of plasma proteins and blood cells with artificial surfaces by applying highly hydrophilic synthetic coatings, such as polyethylene glycol.46 Coating with albumin is another common way to attempt passivation of surfaces: serum albumin, readily adsorbs to most surfaces and does not activate the coagulation cascades. However, exchange processes at the surface47 and physiological degradation of the protein within 2 to 8 weeks, depending on the mode of crosslinking,48 limit the functionality of albumin-based coatings for long-term applications.
Biofunctionalization of Blood-Contacting Surfaces
Biofunctionalization of blood-contacting surfaces is an alternative biomimetic approach that actively promotes anticoagulation mechanisms by simulating some of the antithrombotic functions of quiescent endothelial surfaces of the luminal vessel wall. Heparin remains the most widely used molecule for this purpose.49,50 Heparin-coating leads to increased hemocompatibility by reducing the activation of humoral and cellular responses.51 Although immobilization prolongs the half-life of heparin compared with soluble heparin, metabolic breakdown limits the time span in which the biological function is maintained.52
Currently, passivation of most commercial extracorporeal membrane gas exchange systems is based on ionically or covalently immobilized heparin coatings, such as the albumin-heparin multilayer coating of the iLA System (Novalung, Heilbronn, Germany), BioLine polypeptide-heparin coating of the PLS System (Maquet GmbH Rastatt, Germany),53 or end-point immobilized heparin of the Carmeda Bioactive surface coating (Medtronics, Minneapolis, MN).54 The more complex coating of the Trillium BioSurface in the Affinity NT system (Medtronic, Minneapolis, MN) comprises covalently coupled heparin, polyethylene oxide, and sulfonate groups.55
Endothelialization of Blood-Contacting Surfaces
The concept of improving the hemocompatibility of bioartificial implants by endothelialization has been discussed for more than 3 decades, initially in the context of vascular grafts and cardiac assist devices,56,57 and more recently also for bioartificial lungs.58,59 The goal of endothelializing bioartificial devices is to recreate a quiescent endothelium by coating all blood-contacting surfaces with a durable monolayer of functional ECs.56 In general, the three main problems with endothelializing complex blood-contacting surfaces are 1) creation of a functional monolayer of quiescent ECs, 2) retention of its intactness, and 3) maintenance of its quiescence under diverse flow conditions.
The ECs lining the external surfaces of the hollow fibers forming the mats of the gas exchange membranes in an LAD will be exposed to physiological levels of blood flow.59,60 Mechanical forces, in particular fluid shear stress, regulate EC genotype and phenotype,61 specifically their pro- and antithrombogenic and anti-inflammatory properties.62
One option for increasing retention and functionality of the EC monolayer lining the gas exchange membranes is to dynamically precondition the seeded cells by gradual adaptation to increasing levels of fluid shear stress.59 For example, culturing ECs for up to 14 days under static conditions before exposure to flow,60 or preconditioning/maturation of the cells by gradually increasing the levels of shear stress, significantly improved EC retention,63 yielding EC monolayers that can eventually be exposed to physiological shear rates without any cell loss.58–60
In order to reduce LAD-induced thrombogenicity, several groups have attempted to endothelialize the blood-contacting surfaces of the gas exchange membranes in current LADs.64 For example, Hess et al.65 described the coverage of PMP membranes with human cord blood-derived late outgrowth endothelial colony-forming cells, using a double-layer heparin-albumin coating, and showed reduced platelet adhesion and activation compared with unseeded membranes. In addition to this heparin-albumin coating, which is already in clinical use, other groups have tested adsorption of proteins like collagen, fibrinogen, gelatin, or fibronectin,60,66,67 or covalent attachment of short peptides that selectively support adhesion of ECs, but not of platelets or circulating monocytes.
Taken together, physicochemical modifications of artificial (gas exchange) surfaces as described above effectively promoted initial EC adhesion and subsequent formation of a complete circumferential endothelial monolayer. However, retention of the endothelial monolayer under dynamic physiological flow conditions, especially for prolonged periods of time, remains problematic, in spite of recent significant progress.58,68 Shown in Figure 3 are examples of our own work, where we successfully endothelialized stacks of 8 or 66 gas exchange membranes and retained an essentially intact EC coating for 14 days under dynamic flow conditions. As a caveat, complete and successful endothelialization of a gas exchange membrane with ECs will require several days in the best case. As with all cell-seeded biomedical devices, a general drawback of this approach is that the initial in vitro steps before implantation, i.e., precoating, preseeding and preconditioning, are labor intensive and time-consuming.
In addition to cost and duration of the procedures, cell sourcing may pose another serious, yet surmountable challenge. It has been estimated that it takes 3—6 × 108 ECs to cover the surface of a single miniaturized gas exchange membrane (approximately 0.5–1 m2). One possible solution to this bottleneck is the use of pluripotent stem cells as a convenient source which may yield the required numbers of functional autologous or engineered allogeneic ECs. Given the controversy surrounding the use of human embryonic stem cells, human induced pluripotent stem cells (hiPSCs) are preferred as primary cell source. With the introduction of more efficient methods for their generation, hiPSCs as readily available cell source have become a realistic alternative, specifically for autologous cells derived from prospective patients. For example, Sahara et al.69 (2014) screened over 60 bioactive small bioactive molecules that might promote the differentiation of hiPSCs into endothelial progenitor cells (EPCs). Under optimized conditions, this approach generated over 50% conversion of hiPSCs into EPCs with a high angiogenic and clonogenic proliferation potential and efficient subsequent maturation into functional ECs.69 More recent studies described the direct transdifferentiation of somatic cells into ECs by defined factors and the observation that manipulation of innate immune signaling may modify differentiative decisions and the fate of hiPSC-derived cells.70
Risks and Challenges for Bioartificial Lungs
Current therapeutic modalities of extracorporeal respiratory assist, while offering significant benefits, are also associated with considerable challenges, including primary safety concerns, such as bleeding or thrombosis. Some of these challenges are iatrogenic or technical in nature and can be minimized by choosing the right cannulation71 and avoiding introduction of air-bubbles into the circulation.72 Other risks are possible infections at the site of transcutaneous entry of the extra- or paracorporeal lines into the patients’ body.
Hemocompatibility of the devices is an ongoing cause for concern. Once the blood-contacting surfaces have been modified with proteins, lipids, or ECs, these biolized surfaces need to remain hemocompatible during the entire intended duration of using an LAD. Possible surface passivation strategies, which involve the use of ECs, carry special hazards, despite all advantages described above. One potential pitfall is the reduction in the performance of the gas exchange surfaces after coating with cells.60 Although Hess et al.65 did not find any significant changes in gas transfer before and after seeding PMP fibers with ECs, others reported up to 25% decrease in the transfer rates for CO2 after cell seeding.73 A more recent study from the same group shows a significant augmentation of oxygen transfer through endothelialized gas exchange membranes. Further studies are needed to resolve this issue.
Finally, the day-to-day handling of future LADs will have to be simplified in order to facilitate general physician and patient acceptance for use in an intermediate care setting beyond their availability in specialized ECMO centers. Cost-effectiveness and attractive reimbursement strategies will strongly influence the translation of this promising concept into a routine modality of care for patients in the clinic and at home.
The current version of the miniaturized respiratory assist device (i-lung) is the tangible outcome of European Union (EU) funding for the “Ambulung” Project (http://cordis.europa.eu/result/rcn/175848_en.html).
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