Current Trends and Future Developments on (Bio-) Membranes: Silica Membranes: Preparation, Modelling, Application, and Commercialization 0444638660, 9780444638663

Table of contents :
Cover
Current Trends and Future Developments on
(Bio-) Membranes:
Membrane Applications in
Artificial Organs and Tissue
Engineering
Copyright
Contributors
Preface
1
Artificial kidney: A chemical engineering challenge
Introduction
Historical perspective
Transport properties of membranes for artificial kidney
Compartmental modeling for artificial kidney
Membrane modules for renal replacement therapy
Water quality management in artificial kidney
The future of renal replacement therapy: Bioartificial kidney and implantable artificial kidney
Conclusions and future trends
References
2
Membrane application for liver support devices
Introduction
Liver support devices
Development of blood purification systems
Artificial devices
Hemofiltration
Hemoperfusion
Reactors with immobilized enzymes
Plasma exchange
Hemodiafiltration
Plasma exchange and continuous hemodiafiltration
Molecular therapy of the absorbent recirculation system
Prometheus
Bioartificial devices
Cellular sources
Types of hepatocyte culture systems
Types of bioreactor
Tissue engineering
Implantable engineered tissue for humanized mouse models
Implantable therapeutic engineered liver tissue
Design criteria for implantable systems
Natural scaffold chemistry and modifications
Conclusion and future trends
References
3
Membrane bioreactors for (bio-)artificial lung
Introduction
Milestones in ECMO development to date
Limitations of ECMO / ECLS
Computational fluid dynamics for the optimization of the oxygenator design
Surface treatments for improving the hemocompatibility of blood contacting surfaces in ECMO circuits
Biohybrid/bioartificial approaches
Wearable or implantable artificial lung
The implantable artificial lung
The development of microfluidic artificial lungs
Tissue engineered lungs
Conclusions and future trends
References
4
Membrane bioreactors for bio-artificial pancreas
Introduction: The pancreas
Anatomy and physiology
Mechanisms of glycemic regulation
Physiopathology and treatment
The concept of bioartificial pancreas
Overview of the specificities of currently developed BAP
Number and potential sources of pancreatic islets
Mass transfer issues in BAP and implantation site
Intravascular systems combining convection and diffusion
Design and limits of perfusion chambers
Direct perfusion of encapsulated islets implanted in vascularized organs
Diffusion-based extravascular systems
Omental pouch and intraperitoneal transplantation
Kidney capsule
Subcutaneous tissues
Porous scaffolds—Membranes
Conclusions and future trends
References
5
Membrane devices for blood separation and purification
Introduction
Requirements for an effective apheresis
Main differences between centrifugation and filtration systems
Centrifugal systems
Membrane systems
Mixed systems: Membranes and centrifuges
Discontinuous and continuous flow systems
Apheresis techniques
Plasma exchange
Semiselective plasmas (plasma separation through secondary membranes, or double filtration or cascade filtration)
Selective plasma tracking (selective absorption)
Apheresis of lipidic proteins
Cytoapheresis
Therapeutic cytoapheresis
Donations of autologous or allogeneic blood components
Conclusion and future trends
References
6
Numerical prediction of blood damage in membrane-based biomedical assist devices
Introduction
Properties of red blood cells
Red blood cell structure
Mechanical properties
Phenomenology of hemolysis and blood damage
Quantification of blood damage
Experimental data on blood damage
Current modeling approaches for blood damage prediction
Stress-based models
Threshold models
Continuous models
Limitations of empirical stress-based models
Time-variable stresses
Multicomponent stresses
Flow in ducts
Strain-based models
RBC deformation
Hemoglobin release
Conclusions and future trends
List of acronyms
List of symbols
Latin symbols
Greek symbols
References
Further reading
7
Membrane scaffolds for 3D cell culture
Membrane scaffolds for tissue engineering applications
Tissue engineering principles
Application of tissue engineering methods for 3D cell culture
Prerequisites of membranes for 3D cell culture
Biomaterials for membranes fabrication
Natural origin
Synthetic polymers
Methods of membranes fabrication
Conventional techniques
Freeze drying
Phase separation
Particle leaching
Biofabrication methods for membranes in tissue engineering
Powder-based 3D printing
Extrusion techniques
Direct 3D printing
Fused deposition modeling
3D plotting
Electrospinning
Use of membranes for cell culture in tissue engineering
Limits of conventional tissue engineering
Layer-by-layer bioassembly of cellularized membranes for tissue engineering
Conclusions and future trends
References
8
Artificial oxygen carriers
Introduction
Oxygen is both: A blessing and a curse
The role of AOCs in the context of artificial organs and tissue engineering
Why has evolution developed such strategies?
Considerations addressing Δ x
Considerations addressing Δ c
Relevant types of AOCs
Hemoglobin-based AOCs
Peculiarities and principle of oxygen transport
Limitations
Stage of development of relevant HBOCs
Clinical data/clinical trials
Pipeline (preclinical data)
Perfluorocarbon-based AOCs
Peculiarities and principle of oxygen transport
Limitations
Stage of development of relevant PFOCs
Clinical data/clinical trials
Pipeline (preclinical data)
Conclusions and future trends
Conflict of interest
References
9
Membrane bioreactors for digestive system to study drugs absorption and bioavailability
Introduction
Anatomy of the GI tract
Physiology of the GI tract
Stomach
Duodenum
Jejunum
Ileum and colon
Modeling of drugs’ absorption and bioavailability
Single and two-compartment models
Five-compartments model
Parameters and conditions
Parameter estimation
Model simulations
Conclusion and future trends
References
Index
A
B
C
D
E
F
G
H
I
J
K
L
M
N
O
P
R
S
T
U
V
W
X
Back Cover

Citation preview

Membrane Applications in Artificial Organs and Tissue Engineering

•

Silica Membranes: Preparation, Modelling, Application, and Commercialization (978-0-444-63866-3)

•

Photocatalytic Membranes and Photocatalytic Membrane Reactors (978-0-12-813549-5)

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Carbon Dioxide Separation/Capture by Using Membranes (978-0-12-813645-4)

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Renewable Energy Integrated with Membrane Operations (978-0-12-813545-7)

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Membrane Processes in the Pharmaceutical and Biotechnological Field (978-0-12-813606-5)

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Membrane Desalination Systems: The Next Generation (978-0-12-813551-8)

Current Trends and Future Developments on (Bio-) Membranes

Membrane Applications in Artificial Organs and Tissue Engineering Edited by Angelo Basile Maria Cristina Annesini Vincenzo Piemonte Catherine Charcosset

Elsevier Radarweg 29, PO Box 211, 1000 AE Amsterdam, Netherlands The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States Copyright © 2020 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www. elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN: 978-0-12-814225-7 For information on all Elsevier publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Susan Dennis Acquisition Editor: Kostas Marinakis Editorial Project Manager: Hilary Carr Production Project Manager: Omer Mukthar Cover Designer: Christian J. Bilbow Typeset by SPi Global, India

Contributors Mauro Capocelli  Unit of Process Engineering, Department of Engineering, University Campus Biomedico of Rome, Rome, Italy Sylvain Catros  CHU Bordeaux, Oral Surgery Department, Bordeaux, France A. De  Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG), Hannover Medical School; Lower Saxony Centre for Biomedical Engineering, Implant Research and Development (NIFE), Hannover, Germany Luisa Di Paola  Unit of Chemical-physics Fundamentals in Chemical Engineering, Department of Engineering, University Campus Biomedico of Rome, Rome, Italy D. Dipresa  Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG), Hannover Medical School; Lower Saxony Centre for Biomedical Engineering, Implant Research and Development (NIFE), Hannover, Germany Cornelius Engelman  Liver Failure Group, Institute for Liver and Digestive Health, University College London, Royal Free Campus, London, United Kingdom; Section Hepatology, Department of Gastroenterology and Rheumatology, University Hospital Leipzig, Leipzig, Germany Amal Essaouiba  CNRS, UMR 7338 Laboratory of Biomechanics and Bioengineering, Sorbonne Universities, Université of Technology of Compiègne, Compiègne, France Mathilde Fénelon  CHU Bordeaux, Oral Surgery Department, Bordeaux, France Katja B. Ferenz  University of Duisburg-Essen, Institute of Physiology, University Hospital Essen, Essen; CeNIDE (Center for Nanointegration Duisburg-Essen) University of Duisburg-Essen, Duisburg, Germany Jean-Christophe Fricain  CHU Bordeaux, Oral Surgery Department, Bordeaux, France Vera Guduric  Univ. Bordeaux, Biotis, INSERM U1026, Bordeaux Cedex, France Rachid Jellali  CNRS, UMR 7338 Laboratory of Biomechanics and Bioengineering, Sorbonne Universities, Université of Technology of Compiègne, Compiègne, France S. Korossis  Lower Saxony Centre for Biomedical Engineering, Implant Research and Development (NIFE); Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG), Hannover Medical School, Biomedical Research in Endstage and Obstructive Lung Disease Hannover (BREATH), German Center for Lung Research (DZL), Hannover, Germany; Cardiopulmonary Regenerative Engineering (CARE) Group, Centre for Biological Engineering (CBE), Wolfson School of Mechanical, Electrical and Manufacturing Engineering, Loughborough University, Loughborough, United Kingdom

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Contributors Eric Leclerc  CNRS UMI 2820, Laboratory for Integrated Micro Mechatronic Systems, Institute of Industrial Science, University of Tokyo, Tokyo, Japan Cécile Legallais  CNRS, UMR 7338 Laboratory of Biomechanics and Bioengineering, Sorbonne Universities, Université of Technology of Compiègne, Compiègne, France Pompa Marcello  Unit of Chemical-physics Fundamentals in Chemical Engineering, Department of Engineering, University Campus Biomedico of Rome, Rome, Italy Simone Novelli  Liver Failure Group, Institute for Liver and Digestive Health, University College London, Royal Free Campus, London, United Kingdom; Department of Mechanical and Aerospace Engineering, Sapienza University of Rome, Rome, Italy A. Silva Peredo  Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG), Hannover Medical School; Lower Saxony Centre for Biomedical Engineering, Implant Research and Development (NIFE), Hannover, Germany M. Pflaum  Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG), Hannover Medical School; Lower Saxony Centre for Biomedical Engineering, Implant Research and Development (NIFE), Hannover, Germany Vincenzo Piemonte  Faculty of Engineering; Unit of Chemical-physics Fundamentals in Chemical Engineering, Department of Engineering, University Campus Biomedico of Rome, Rome, Italy Renzo Pretagostini  Department of General Surgery, and Organ Transplantation, Sapienza University of Rome, Rome, Italy Luca Turchetti  ENEA—Italian National Agency for New Technologies, Energy and Sustainable Economic Development, Rome, Italy Flavia Vitale  Department of Neurology, Bioengineering, Physical Medicine and Rehabilitation, University of Pennsylvania, Philadelphia, PA, United States

x

Preface Today, biochemical, biomedical, and biotechnological applications occupy a central and consolidated place in the global research scene. At the same time, the use of membranes in biotechnological and biomedical industrial applications has become increasingly important, making it possible to achieve promising results. In particular, research in the field of tissue engineering and regenerative medicine is currently assisting in the flourishing production of experimental results regarding the application of novel biotechnologies to the development of artificial tissues and organs (artificial kidney, artificial liver, artificial lung, etc.). The introduction of novel materials, microfabrication technologies and advanced computational methodologies, is unveiling novel biological mechanisms in cell and tissue behavior and enlightening new avenues in tissue engineering development of artificial organs. For all these reasons, the present book focuses on the membrane applications in the field of Artificial Organs and Tissue Engineering. It consists of nine chapters. In Chapter 1 (Di Paola), after a historical perspective, with a perspective to future developments, regarding bioartificial implantable devices, for a definitive, complete replacement of the renal function, a survey of current state-of-the-art devices is presented. The main focus is to highlight the great role that both membrane science and chemical engineering play in the development of innovative technologies in the field of kidney dialysis. Chapter 2 (Novelli, Engelman, and Piemonte), underlines that liver disease, by afflicting over 750 million people worldwide (30 million Europeans), leads to the death of over 40,000 individuals in the EU every year. According to the authors, liver failure can be generally separated into two major categories: fulminant hepatic failure (also referred to as acute liver failure), and chronic hepatic failure (resulting from chronic end-stage liver disorders). Given the steady rise in patients with liver disease, the need for liver transplantation has continued to increase. However, the number of available donor livers has not changed significantly in the last five years. The chapter illustrates that, as a result, several different approaches have been undertaken, such as: extracorporeal devices perfuse patient’s blood or plasma through bioreactors containing hepatocytes, hepatocytes are transplanted directly or implanted on scaffolds, transgenic animals are being raised in order to reduce complement-mediated damage to address this growing need of donor organs. xi

Preface Chapter 3 (Pflaum, Peredo, Dipresa, De, and Korossis) starts by considering that membrane oxygenation devices, also referred to as artificial lungs, have been developed to provide sufficient blood gas exchange during cardiopulmonary bypass, and enable respiratory support for patients suffering from severe end-stage lung diseases. However, malfunctions originating from various causes can reduce the lifetime of contemporary devices, which excludes their use in destination therapy. Therefore, this chapter reports on current research and development focused on the advancement of hollow fiber membrane technology to achieve extended durability, safer application, and higher efficiency of contemporary oxygenators. In addition, alternative approaches for the development of an artificial lung are described. In Chapter 4 (Jellali, Essaouiba, Leclerc, and Legallais) it is stressed that designing an efficient and functional bioartificial pancreas (BAP) at human scale to treat type I diabetes remains the Holy Grail in the 2010s, although investigations started in the 1970s. Biomimetic approaches need to be performed and evaluated to offer insulin secreting cells an environment close to the native pancreas, but also accounting for the interactions between the pancreas and other organs, and more specifically liver. The authors, after reviewing the pancreas anatomy and physiology and the current treatments for diabetic patients, study how to fulfill (or not) the requirements for an efficient BAP, regarding its different components: cell types, encapsulation methods, membranes, and devices dedicated to several implantation sites. They conclude with a short discussion of future directions including BAP revascularization to improve the exchanges with the host and the impact of microtechnologies on the development of next generation of BAP. Chapter 5 (Novelli, Pretagostini, and Piemonte) shows that apheretic therapies play an important role in the treatment of many diseases, both as first-line treatments and as rescue therapies, after failure or pharmacological toxicity or, again, in patients where it is desirable to achieve the therapeutic objective in a short time. Apheretic technologies have evolved at a dizzying pace in recent decades. Therapeutic apheresis is almost always part of a therapeutic plan; therefore, a patient-centered approach to choose the most appropriate treatment, balancing personal preferences, drug interactions, and technological availability, can significantly influence the choice of technology/protocol to be used. The authors conclude that, if the wide variety of apheretic treatments can offer an increasingly personalized approach to the patient, it can also create an apprehension about the choice of the most appropriate therapy. Chapter 6 (Turchetti and Vitale) shows that biomedical assist devices, such as membrane-based blood detoxification systems, are today widely used in clinical practice and play a key role in the management of chronic and acute conditions like kidney or liver failure. While flowing inside these devices, blood undergoes nonphysiological mechanical solicitations that can eventually damage red blood cells, leading to partial or total release of their hemoglobin content into plasma. In particular, the state of the art in numerical blood damage prediction is presented, xii

Preface where the current models are reviewed as well as the experimental data available in the literature for their validation being shown. The limitations of the most commonly used modeling approaches are discussed and the desirable features of new models to be developed indicated. Chapter 7 (Guduric, Fenelon, Fricain, and Catros) underlines that 3D organization of cells and biomaterials (scaffolds) for tissue engineering enhances cell-to-cell communication through growth factor secretion and extracellular matrix formation. The main limitation of tissue engineering lies in the low penetration of cells in inner parts of large engineered products due to the insufficient diffusion of oxygen and nutrients. Membranes for tissue engineering are of natural (animal or human) or synthetic (biopolymers) origin. A disadvantage of human origin membranes is the possibility of immune response, pathogenic transmission, and availability, which is overcome in the case of synthetic biopolymers. There is a wide range of possibilities in combinations of membranes seeded with cells to fabricate final complex tissue engineered products with controlled 3D micro-architecture and microenvironment. In Chapter 8 (Ferenz), it is clarified that the development of modern tissue engineering would have been impossible without the implementation of artificial oxygen carriers (AOCs) as homogenous oxygenation of cells and tissue becomes a major issue in huge bioreactors or scaffolds of large extent. This chapter explains why oxygen supply is important in general and especially in engineered tissue and cell culture, what happens if too much (hyperoxia) or to little (hypoxia) oxygen is available and how AOCs can contribute to improve this parameter. The topic is approached from the physiological/biochemical point of view and substantiated by reaction equations to guide the reader through the complex field of tissue oxygenation. The chapter focuses on classic and, most important, rechargeable AOCs such as hemoglobin-based oxygen carriers (HBOCs) and perfluorocarbon-based oxygen carriers (PFOCs). Chapter 9 (Pompa, Capocelli, and Piemonte) describes a new Gastro Intestinal (GI) model, based on membrane bioreactors, able to evaluate the bioavailability of drugs, including in the presence of food. It includes the physiology of five compartments: stomach, duodenum, jejunum, colon, and blood. All the compartments are interconnected with valves able to control the input/output flux from the compartments (discontinuous running mode). The molecule passage in the blood occurs though a transmembrane flux. The simulation of the GI absorption starts with the input of the bolus in the stomach. The emptying of the stomach is controlled by enzymatic reactions in the duodenum, therefore controlling this stage without any external function or parameter, but only using the physiological process. The parameters used work for the optimal condition of the reactions (pH in optimal range, presence of bile salts, etc.). Other aspects of the chapter are also related to the model validation and the data obtained from the model are comparable with in vivo results from literature. To conclude, the book is addressed to postgraduate students and researchers in the field of chemical and biomedical engineering and is aimed at providing the reader with tools for the xiii

Preface analysis of existing devices and, possibly, the design of new ones. Furthermore, the book may be recommended as a supplementary textbook for higher-level undergraduate courses and be of interest for medical researchers who want to get a deeper insight into the working principles of artificial organs and tissue engineering. Last but not least, we wish to take this occasion to thank all the authors of the chapters for their strong work and patience in reviewing—sometimes several times—their chapters under the comments/suggestions of the editors. Special thanks also to the staff of Elsevier for following our work step by step and helping us when necessary. Angelo Basile Vincenzo Piemonte Catherine Charcosset Maria Cristina Annesini

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CHAPTE R 1

Artificial kidney: A chemical engineering challenge Luisa Di Paola Unit of Chemical-physics Fundamentals in Chemical Engineering, Department of Engineering, University Campus Biomedico of Rome, Rome, Italy

1 Introduction Substituting the function of organs is one of the great challenges of biomedical sciences. Artificial organ engineering is the specific branch of biomedical engineering devoted to this aim. Excretory organs, such as kidney and liver, resemble small-scale plants, so the design of their artificial counterparts is based on the background of chemical engineering and membrane science (Annesini, Marrelli, Piemonte, & Turchetti, 2017). The history of artificial organs starts with the development and use of the artificial kidney: its design represented the primer for other artificial organs based on membrane diffusion (artificial liver and lung). Renal replacement therapy (RTT) by means of renal hemodialyzers has become a standard of the nephrological clinical applications, becoming a synonym for dialysis. The regular application of RRT has become a life-long, life-saving therapy for millions of patients throughout the world. The kidney performs many activities: the removal of water-soluble end-products of body metabolism, regulation of the osmotic pressure in tissues (removal of access water), control of electrolyte homeostasis, and regulation of blood pressure. All these activities are based on osmoregulation of water-soluble blood components through dialysis. The functional unit of kidney is the tubule. Tubules comprise two subunits: nephrons, active in urine production and collective tubules, devoted to urine collection and convection to the excretory part of the urinary system. The kidney contains around 1–1.5 million nephrons working in parallel (Preuss, 1993). Each tubule is made of a tuft of small tangled blood vessels, called Malpighi glomerulus, embedded within a cup called the Bowman capsule. Malpighi renal corpuscles comprise glomerulus and capsule.

Membrane Applications in Artificial Organs and Tissue Engineering. https://doi.org/10.1016/B978-0-12-814225-7.00001-2 Copyright © 2020 Elsevier Inc. All rights reserved.

1

2  Chapter 1 Blood enters the kidney at a flow rate of about 1.2 L/min (around 20% of the cardiac output) and reaches nephrons, where around 120 mL/min is forced through the glomerular membrane into the Bowman’s capsule. This high filtration rate is ensured by a high filtration area (around 2 m2). The glomerular filtration rate (GFR) is defined as “the volume of water out of the plasma through glomerular capillary walls into Bowman’s capsule per unit time” (Jin, Grunkemeier, Brown, & Furnary, 2008) and expressed as: GFR =

Urine concentration·Urine flow  mL    Plasma concentration  minn 

(1)

GFR can be also described in terms of ultrafiltration rate: the difference between the hydrostatic pressure ΔP and the osmotic/oncotic pressure Δπ, between the glomerular capillaries and the Bowman’s capsule, is the driving force for ultrafiltration through the glomerular membrane. Thus, GFR is defined as: GFR = K GF ·( ∆ PG − ∆π G )

(2)

where KGF is the hydraulic permeability of the glomerular membrane, and ΔPG is the hydrostatic pressure difference between the glomerular capillaries (around 55 mmHg) and the Bowman’s capsule (around 14 mmHg). The difference of osmotic/oncotic pressure ΔπG is more or less the measure of the oncotic pressure of plasma proteins, being the filtrate virtually devoid of proteins. Generally, GFR is estimated rather than measured, due to the high costs of the clinical tests. The estimated glomerular filtration rate (eGFR) (Levey et al., 1999) is:   mL eGFR  = 186 ⋅ Serum creatinine 2   min 1.73 m    mg  −1.154    −0.203  ·0.742 ( if female )·1.212 ( if black )   dL  · Age    

(3)

GFR depends on body size: smaller persons naturally show a physiological lower GFR. To scale GFR to different body sizes, the GFR is indexed to the body surface area (BSA) of an adult male (1.73 m2). GFR values are correlated to the physiology of kidney and are used in clinical practice to assess pathological conditions. RRT applies to two kidney pathologies: acute kidney injury or acute renal failure (AKI or ARF) and chronic kidney disease (CKD). AKI and ARF are elusive yet severe pathological conditions, due to the sudden loss of kidney function (in a few hours or days) affecting

Artificial kidney: A chemical engineering challenge  3 5%–10% of all hospitalized and 30%–50% of intensive care unit (ICU) patients, with a mortality range higher than 50% in ICU setting (Bellomo, Kellum, & Ronco, 2012; Molitoris & Reilly, 2016). These patients are generally treated with continuous renal replacement therapy (CRRT), which requires a 24/7 application of renal dialysis treatment, whereas CKD is generally treated intermittently (3–4 times per week). Chronic kidney disease (CKD) is defined as (Matovinović, 2009) “the kidney damage for more than 3 months, as defined by structural and functional abnormalities of the kidney, with or without decreased GFR, that can lead to decreased GFR, manifest by either: • •

Pathologic abnormalities; or Markers of kidney damage, including abnormalities in the composition of blood or urine, or abnormalities in imaging tests.”

mL In CKD, values of GFR are below 60 for more than 3 months, with or without min⋅ 1.73 m 2 kidney damage. Nowadays, CKD affects 10% of the population worldwide and is far more prevalent in elderly people. Its final stage, end stage renal disease (ESRD), affects around 3.2 million patients worldwide, but this number likely represents only 10% of those really suffering from it, referring only to recorded patients accessing therapies, and is expected to increase by 6% annually (Boschetti-De-Fierro et al., 2017; Heaf, 2017). In 2017, 320 million dialyzers were in use (Boschetti-De-Fierro et al., 2017). Considering the average cost of renal dialysis treatment is 40,000–80,000 €/year per individual, the overall costs worldwide for treated ESRD patients totals 8 trillion €. The majority of treatments are carried out in five countries (United States, Japan, Germany, Brazil, and Italy) even though they account for only 12% of global population. Numbers are expected to grow in the next few years, due to the increasing age in OCSE countries and to the increasing number of diagnosed and treated patient in developing countries. The ultimate, conclusive therapy for ESRD is renal transplantation, but viable organs cover only a small percentage of the patients list. Hemodialysis is a bridge therapy, which in turn, for most patients, is the long-life alternative to transplantation. Hemodialysis faces many challenges to improve many aspects of the treatment: first of all, costs must be reduced to guarantee accessibility of the therapy to a larger share of ESRD patients worldwide; technology must provide new solutions to improve the efficiency of renal hemodialysis, to increase positive outcomes, especially for fragile patients (elderly, obese, etc.); and last but not least, innovative design of hemodialyzers (wearable or implantable) can ensure an improvement in the quality of life of patients. To this aim, a continuous technological development of hemodialyzers is required to provide more and more convenient, efficient, and tolerable renal hemodialysis therapy (Velasco, 2013).

4  Chapter 1 This chapter will provide a general historical perspective on renal hemodialysis (Section 2), meant to outline the technological development of all components of hemodialyzers, ending up with a perspective on future development of bioartificial and artificial, implantable devices. Section 3 will report a survey on membranes for hemodialyzers, starting from materials, passing through hollow fiber devices design, and ending up with the modeling of transport in hemodialyzer membranes. Compartmental modeling of hemodialyzer dynamics is looked at in Section 4, while Section 5 will deal with the issue of water quality management in hemodialysis application. Finally, Section 6 will illustrate the main future developments of innovative hemodialysis devices, to improve therapy efficiency and agreeableness.

2  Historical perspective Thomas Graham, recognized as the father of modern dialysis, devoted part of his prolific scientific activity to the study of dialysis (Gottschalk & Fellner, 1997). In a seminal experiment, he demonstrated that the diffusion of a solvent through a semipermeable membrane, separating a compartment containing pure solvent from one containing a solution with the same solvent generates, a force, which he defined as “osmotic force.” This experiment is the basis of the modern definition of osmosis. Later, in 1861, he also set the basis for colloid science by distinguishing the division of solutes into crystalloids (salts, sugars) and colloids (gelatin, starch, gum). Gelatinous colloids are the basis for semipermeable membranes, through which he observed the diffusion of molecules; he named this flux dialysis, due to the difference of concentration between compartments separated by a semipermeable membrane. In 1913, J.J. Able, L. Rowntree, and B.B. Turner performed the first application of hemodialysis to in vivo systems (uremic animals). In 1914 they published the results of the application in two papers, where they described their device, for the first time, using the term “artificial kidney.” It was composed of celloidin tubes immersed in a dialysate bath contained in a glass jacket. Anesthetized animals were connected to the device via an arterial cannula, blood was returned via a venous cannula. The pumpless system was driven by arterial pressure. Hirudin, extracted from leeches, acted as anticoagulant for the blood (Clark, 2000). They worked to reduce the volume of blood retained in the device and to maximize the surface area for diffusion “by very small tubes or by flattening a larger tube until the opposite surfaces nearly touch.” This outlined the basic guidelines for the design of the hollow fiber dialyzer. In October 1924, George Haas applied hemodialysis to human patients for the first time. It lasted only 15 min; vascular cannulas were inserted into the left radial artery and antecubital vein under anesthesia. Haas measured a blood clearance of 150 mL; incidentally, the application was completely ineffective, due to the short application times and low blood and dialysate flows. In 1928, heparin became available, greatly reducing the costs for hemodialysis operations.

Artificial kidney: A chemical engineering challenge  5 The first hemodialyzer configuration was the rotating drum dialyzer, adopted by W.J. Kolff and H. Berk (Gottschalk & Fellner, 1997): blood was sent into a long cellophane tube arranged in a spiral around a cylinder, rotating in a still dialysate bath. Due to the low blood hydraulic resistance, no inlet blood pump was necessary, blood flow was allowed simply by an arterial cannula. Due to the very low transmembrane pressure, a minimal plasma water ultrafiltration occurred, so the toxins removal was primarily osmotically driven. The efficiency of the device was very low (very low toxins removal), while the blood volume was high. For these reasons, the rotating drum dialyzer was continuously evolved to reach a reasonable efficiency to be applied in clinical practice (Fig 1). To improve the membrane surface area, and dialyzer performance in general, in the following years, many other dialyzer configurations were proposed (Twardowski, 2008): the twin-coil dialyzer (Holmes, Stonington, van Schoonhoven, Richey, & Takeda, 1958), multiple and single-use parallel-flow dialyzers (Alwall, 1968; Funck-Brentano et al., 1969), and hollow fiber dialyzers, first introduced by Lipps in 1967 (Gotch et al., 1969). Actually, clinical practice is based on hollow fiber dialyzers, called hollow fiber artificial kidney (HFAK), where the blood and dialysate flow in countercurrent. The original version of HFAK was made of 12,000 fibers (inner diameter around 200 μm), with a surface area around 1 m2; the membrane material was regenerated cellulose (Twardowski, 2008).

Rotating coupling

30–40 m cellophane tubing

Stationary 100 L tank

Fig. 1 Scheme of the rotating drum dialyzer. Reproduced with permission from Gottschalk, C. W., Fellner, S. K. (1997). History of the science of dialysis. American Journal of Nephrology, 17, 289–298. doi:https://doi. org/10.1159/000169116.

6  Chapter 1 The success of the hollow fiber configuration is due to different fluid dynamic features: it allows relatively high shear rates, with moderate axial pressure drops (reduced boundary layer effect), resulting in good mass transfer. The development of HFAK was focused on membrane material (substituted cellulose and synthetic dialyzers (Krieter, Lemke, & Wanner, 2008)) to improve membrane solute permeability. Consequently, HFAK adopted membranes of substituted cellulose and synthetic materials (such as polysulfone), showing high permeability and low costs.

3  Transport properties of membranes for artificial kidney Membrane-based RRT is generally known as dialysis, but it covers different processes according to the transport mechanism of toxins through membranes. Solute transport from blood to dialysate across membranes occurs thanks to different mechanisms: convection, diffusion, and a combination of both. In the case of pure or prevalent convection, the operation is called hemofiltration, while if diffusion prevails, it is known as hemodialysis. If a combination of both comes into play, it is referred to as hemodiafiltration. When the convective transport is not negligible, it is necessary to restore the water volume removed by the operation. The efficiency of toxin removal is generally better when filtration occurs, but in turn, the selectivity of solute removal is very low. However, the toxin removal efficiency in hemofiltration is generally higher, especially for compounds whose diffusive flow is limited by size (Meyer et al., 2005). Modeling of water and solute transport across membranes is essential for HFAK design and control. In the case of uniform transport through the membrane, the volume (water) and solute fluxes are, respectively (Waniewski, 2006):

and

JV = LP ·( ∆ P − σ · RT ·∆c )

(4)

J S = P·∆c + (1 − σ ) JV cm

(5)

where: LP is the hydraulic permeability of the membrane; P is the hydrostatic pressure (and ΔP the transmembrane pressure difference); σ is the Staverman reflection coefficient; Δc is the transmembrane concentration difference; and cm is the mean intermembrane concentration of the solute, defined as (Villarroel, Klein, & Holland, 1977): cm = (1 − f )·c1 + f ·c2 (6)

Artificial kidney: A chemical engineering challenge  7 c1 and c2 being the boundary values of solute concentration at the two membrane faces, while f is defined as (Waniewski, Lucjanek, & Werynski, 1993): f =

1 1 1−σ − λ= P λ JV exp ( λ JV ) − 1

(7)

The parameter Pé = λJV is the Péclet number and describes the importance of the convective flux when compared to the diffusive one (high values of Pé indicate a prevalent conductive flux through the membrane). The mass balance equations for the two components, blood (B) and dialysate (D), flowing in countercurrent in steady state are: d ( QB cB ) dx and

d ( QD cD ) dx

= −J S A

(8)

= − J S Am

(9)

where: QB and QD are the volume flows of blood and dialysate, respectively; cB and cD are the corresponding solute concentrations; Am is the overall membrane surface area; and. x is the axial coordinate, originating to the blood inlet section. In general, the rate of volume flows Q and the concentrations c vary along x. The solute flux JS, according to Eq. (5), becomes: J S = P·( cB − cD ) + (1 − σ ) JV cm

(10)

If the ultrafiltration volume flux JV is constant, QB and QD linearly vary with x, as follows: QB ( x ) = QBi − QU · x

(11)

QD ( x ) = QDi − QU · x

(12)

where QBi and QDi are the volume flows at the inlet section for the blood and the dialysate compartments, respectively; and QU = JV · Am is the ultrafiltration volume flow. The solute clearance is a parameter to quantify solute removal from blood and originates from renal physiology in the 1950s. The clearance K of a solute from blood to dialysate is defined as (Waniewski et al., 1991): Q c − QBi cBi K = Bo Bo (13) cBi − cDi

8  Chapter 1 where QBo = QBi − QU are the blood volume flows at the outlet of the blood compartment. The ultrafiltration volume flow QU enhances the solute removal, so the clearance K can be imagined as made of two terms (Waniewski et al., 1991): K = K 0 + Tr·QU

(14)

where K0 represents the diffusive clearance (at QU = 0), and Tr · QU represents the contribution due to the ultrafiltration flow (convective flux). Tr is the transmittance coefficient. In the case QB ≠ QD, K0 assumes the following form: K 0 = QB ·

exp ( γ ) − 1 Q exp ( γ ) − B QD

(15)

where γ = Pm A· 1 − 1 , Pm is the membrane permeability, and A is the membrane QD   QB surface area. The transmittance coefficient Tr is derived as: Tr = 1 −

K 0 ∆cBo QBo − · QBi cBi QF

(16)

where: QF = QBi − QBo is the ultrafiltration volume flow; ΔcBo = cBo − c′Bo, cBo is the outlet blood concentration in the presence of ultrafiltration (QF ≠ 0); and c′Bo is the outlet blood concentration in the presence of ultrafiltration (QF = 0).

4  Compartmental modeling for artificial kidney Kinetic modeling allow control of substance distribution in different body regions during and between treatments (Sprenger, Kratz, Lewis, & Stadtmuller, 1983). Compartmental modeling is widely used in pharmacokinetic and pharmacodynamic studies to analyze drug distribution in body regions (compartments) by means of unsteady state models describing the dynamics of drug distribution (Bassingthwaighte, Butterworth, Jardine, & Raymond, 2012). The single compartment model schematizes the patient’s body as a homogeneous compartment of fluid of uniform solutes concentration; the distribution volume corresponding to the total body water (see Fig. 2).

Artificial kidney: A chemical engineering challenge  9

Fig. 2 Single compartment patient-dialyzer model.

In the model, the generation rate G (mg h−1) for toxins in the body is considered, along with a residual renal clearance KR (mL h−1); if K (mL h−1) is the dialyzer clearance, the toxin removal by the dialyzer is (see Eq. 13) K · cBi. The toxin unsteady state mass balance is: d (VB cBi ) dt

= − K R cBi − KcBi + G

(17)

Knowing the initial values for VB and cBi, along with the parameters K, KR, and G, Eq. (17) allows us to predict the time course for VB and cBi. However, it is unsuitable to predict toxins concentration between therapy and the distribution of toxins between different body regions. Specifically, the model cannot predict the urea rebound occurring when the plasma urea level goes down upon hemodiafiltration and urea is called back from the intravascular into the blood compartment (Spiegel, Baker, Babcock, Contiguglia, & Klein, 1995) (see Fig. 3). To describe the time course of toxin (urea) distribution between the vascular and extracellular compartments, two compartments models have been introduced (Daugirdas & Smye, 1997) (see Fig. 4). The intracellular fluid compartment ICF (constant volume V1, urea concentration C1) exchanges urea with the extracellular fluid compartment ECF (constant volume V2, urea concentration C2) with a mass transfer coefficient X, being a combination of cell permeability and tissue perfusion (Azar, Yashiro, Schneditz, & Roa, 2013); ECF exchanges urea with the dialyzer as well, and the exchange is characterized by the clearance K. The single compartment model is a special case of the two compartments in the case of X ≫ K; this hypothesis is true for standard adult dialysis (K = 50 − 150 mL/min), but for high efficiency dialysis (K > 190 mL/min) this hypothesis does not hold.

Blood urea concentration

10  Chapter 1

C0

Ce C(T)

Time

T

Fig. 3 Urea rebound: Blood urea concentration reaches a minimum C(T) after therapy (duration T). As soon as the therapy is interrupted, urea is called back from the intracellular compartment into the blood, reaching an equilibration value Ce.

Fig. 4 Two compartments model: The intracellular compartment ICF exchanges urea with ECF (mass transfer X), which in turn exchanges urea with the dialyzer.

Urea mass balances for the two compartments are (Smye, Evans, Will, & Brocklebank, 1992): V1 and

dC1 = − X ·( C1 − C2 ) dt

(18)

dC2 = X ·( C1 − C2 ) − KC2 + G (19) dt supposing toxin generation occurs only in the ECF, perfused compartment, according to (Annesini et al., 2017). V2

In the case of negligible generation rate G, the solution for the system of Eqs. (18), (19) is:

where a=

C1 = α1 Ae − λ+ t + α 2 Be − λ− t

(20)

C2 = Ae − λ+ t + Be − λ− t ’

(21)

KV V1V2 b = V + 1 V = V1 + V2 X X

(22)

Artificial kidney: A chemical engineering challenge  11

λ+ =

(

b + b 2 − 4 aK

)

1

2a

α1 = 1 +

(

b − b 2 − 4 aK

2

λ− =

2a

K λ+V2 K λV − α2 = 1 + − − 2 X X X X

µ=

A=

C2 ( 0 ) 1+ µ

1 − α1 α2 − 1 B=

C2 ( 0 ) µ 1+ µ

)

1

2

(23)

(24)

(25)

(26)

This model is able to describe the urea rebound: the intracellular concentration C1 is always higher than the extracellular concentration C2, so when the dialysis is interrupted (K = 0 in Eqs. 18, 19), a continuous flow of urea occurs from the intracellular to extracellular compartment, until C1 = C2 = Ce (see Fig. 3).

5  Membrane modules for renal replacement therapy The search for cheaper and more efficient devices for RRT drives a continuous research effort in the field of membrane technology. The main challenge is to find materials and dialyzer configurations able to perform the separation of a large spectrum of molecules in a reduced space and with the lowest possible operation times. Hemodialysis and hemofiltration membranes are generally classified according to materials (cellulosic or synthetic) and water permeability (low or high flux) (Ronco & Clark, 2018). Cellulosic membranes, adopted in the very early stages of renal hemodialysis, cause complement activation, resulting in very low biocompatibility. They show a general high efficiency in removing small MW molecules, but not middle MW. Hydroxyl groups in the cellobiose (monomeric unit of cellulose) molecule are replaced by other groups in substituted cellulose, with improved blood compatibility features. The substitution of hydroxyl groups also results in an increased pore size, with corresponding improved water permeability and removal efficiency for middle MW molecules. However, residual hydroxyl groups still impair blood biocompatibility, since they are likely to induce blood platelet activation. Synthetic membranes were originally developed to overcome the biocompatibility issue and the low permeability of cellulosic membranes. Polysulfone and polyamide were introduced for hemofiltration in the late 1970s (Lysaght, 1995). RRT membranes preparation includes precipitation from a polymer solution by phase inversion (Strathmann & Kock, 1977; Van De Witte, Dijkstra, Van Den Berg, & Feijen, 1996):

12  Chapter 1 pore size is controlled by operative conditions and solvent. To better control microporous membrane structure, recently 3D printing technology has been introduced (Düregger, Trik, Leonhardt, & Eblenkamp, 2018). In current clinical practice, hollow fiber modules are used: this configuration is by far the most used, due to its unique properties—essentially the high specific membrane surface area and the modularity. The design of hollow fiber dialyzers took inspiration from the shelland-tube heat exchanger, a common device in the chemical processing industry, meant to efficiently exchange heat between two fluids. The translation of common knowledge from chemical engineering practice allowed the innovation in biomedical technology. In turn, the development of more and more convenient and efficient membranes for hemodialysis has driven the development and the use of membrane separation units in chemical processing.

6  Water quality management in artificial kidney The sustainability of hemodialysis is a recent, yet urgent issue in clinical practice and in device design. The central problem is water consumption, which represents a main hurdle in hemodialysis diffusion in poor countries with a water shortage. In this context, today, water quality and quantity has become a central issue in hemodialysis. Water consumption strongly determines the environmental footprint of renal dialysis: renal dialysis therapy consumes, for each patient, around 500 L per week of mains water, producing a corresponding amount of wastewater, generally disposed in sewers (Connor et al., 2010). High purity water for dialysis is prepared starting from mains water, producing a large quantity (around 250 L per session) of “reject water”: this water, although meeting the requirements as drinking water, is generally disposed of. However, in drought-prone countries, reject water is in demand for alternative uses (irrigation, laundry, sanitation) (Agar, Simmonds, Knight, & Somerville, 2009) and, more in general, this option will become more and more adopted to reduce the water footprint of nephrological units. While the reject water is a predialysis effluent, after the blood contact, spent dialysate is produced and disposed as wastewater. Water quality strongly affects the dialysis outcomes (Coulliette & Arduino, 2013): patients are exposed during each treatment to 300–600 L of treated water in the form of dialysate. Chemical and microbiological contaminants present in the dialysate can cross the membrane and affect the outcome of the dialysis operation. Thus, stringent requirements on water treatment and quality for dialysate preparation are followed by renal dialysis units worldwide (Ward, 2009). Dialysate is generally prepared starting from mains water using a purposed purification procedure. Fig. 5 reports a scheme of a typical preparation process for dialysate. Reverse osmosis is the most common primary purification process; alternatively, deionization can be used to supplement or substitute reverse osmosis. The large variability of mains water requires each purification system must be customized for the nephrological units. One of the

Artificial kidney: A chemical engineering challenge  13 Primary purfication

Pretreatment

Blending valve

Reverse osmosis Depth filter

Carbon

Softener

5 µm filter

Municipal water

Dialysis stations

Storage tank Distribution

Ultrafilter

Fig. 5 Scheme of the water purification and distribution for hemodialysis. Reproduced with permission from Ward, R. A. (2005). Dialysis water as a determinant of the adequacy of dialysis. Seminars in Nephrology, 25, 102–111. doi:https://doi.org/10.1016/j.semnephrol.2004.09.017.

main issues is the presence of high levels of chlorine and chloramine, which can provoke encephalopathy syndrome. To avoid this contamination, carbon adsorption is included in the pretreatment cascade (Ward, 2005).

7  The future of renal replacement therapy: Bioartificial kidney and implantable artificial kidney Innovation in hemodialysis is continuously looking for solutions to improve the efficiency and the acceptance of hemodialysis therapy. The final goal is to develop a wearable and implantable artificial kidney. Another key point is to develop bioartificial devices to cover all kidney functions, so far not performed by purely artificial equipment. The introduction of biological elements not only improves the toxins removal rate, but also provides additional functions (such as osmoregulation and hormone excretion) far beyond the activity of current equipment. Membrane materials determine membrane permeability and selectivity: innovative materials allow an improved removal of uremic toxins, reducing the therapy duration and, thus, risks for patients, while improving quality of life (Boschetti-De-Fierro et al., 2017).

14  Chapter 1 Leonard and coworkers developed a membraneless dialysis device, based on microfluidics (Leonard, Cortell, & Vitale, 2005): at low Reynolds number a parallel flow of the two miscible liquids is possible. Solute transfer occurs at the interface of contact between the two liquids. Nissenson and coworkers introduced a nanotechnology method to mimic the nephron filter (Nissenson, Ronco, Pergamit, Edelstein, & Watts, 2005): they developed a two-stage filtration device resembling roughly the glomerulus and tubule. The glomerulal membrane is a conventional hemofiltration membrane, allowing the removal of small molecular size solutes, retaining cells, and macromolecules. The tubular membrane is a synthetic membrane including synthetic ion and aquaporin (macromolecular) channels, which allows a re-equilibration of water and salts in the blood, while concentrating toxins in the ultrafiltration stream. The artificial kidney is not a complete replacement therapy, since it supplies only the filtration function, not all other kidney functions (homeostasis, metabolic, and endocrine functions). The introduction of a biological component (bioartificial kidney) is the only route for a complete replacement strategy (Humes, Buffington, Westover, Roy, & Fissell, 2014) (see Fig. 6): the first cartridge provides hemofiltration, the second in series is a renal tubule cell-assist device (RAD). The ultrafiltrate is conveyed in the lumen of the RAD cartridge, which contains the cells, while the postfiltered blood is sent into the extracapillary

Fig. 6 A bioartificial kidney (BAK) used to treat patients with acute kidney injury. The first cartridge provides hemofiltration, the second in series is a renal tubule cell-assist device (RAD). The ultrafiltrate is conveyed in the lumen of the RAD cartridge, which contains the cells, while the postfiltered blood is sent into the extracapillary compartment of the RAD. The processed luminal ultrafiltrate from the RAD is sent to disposal (sewer), while the processed blood is sent back to the patient. Reproduced with permission from Humes, H. D., Buffington, D., Westover, A. J., Roy, S., Fissell, W.H. (2014). The bioartificial kidney: current status and future promise. Pediatric Nephrology, 29, 343–351. doi:https://doi.org/10.1007/s00467-013-2467-y.

Artificial kidney: A chemical engineering challenge  15 compartment of the RAD. The processed luminal ultrafiltrate from the RAD is sent to disposal (sewer), while the processed blood is sent back to the patient. First clinical results show BAK therapy can not only guarantee the hemofiltration activity in a comparable way with a traditional artificial kidney, but also endocrine and metabolic functions, producing dynamic and personalized responses in patients (Humes et al., 2004); the ultimate frontier for the bioartificial kidney is the development of an implantable BAK: Roy et al. (2009) explored the feasibility of an implantable RAD with the use of high performance silicon nanoporous membranes.

8  Conclusions and future trends It is evident that, despite the long history of artificial kidneys (almost a century), there are many aspects that must still be developed. The steeply growing demand will foster in the future the fast development of high-performance and better acceptance devices, pointing toward the development of a long-term replacement device, such as the implantable bioartificial kidney. Chemical engineering science provides all tools for the development of new generation renal replacement devices, in the fields of: • • •

Microfluidics, for the development of miniaturized, membraneless devices; Bioreactor design, for the design and optimization of the biological (cells, tissues) part of the bioartificial kidney; Membrane science: • Materials: the development of new nanoporous materials for artificial kidney will improve blood compatibility and performance; • Membrane transport: the analysis of transport in membranes will help in developing high performance devices; • Membrane modules: the application of fluid dynamics study will develop optimized design of membrane modules.

The development of the artificial kidney demonstrated the strong interconnection between biomedical technology and chemical engineering science: the development of more and more affordable and efficient membranes for hemodialysis also changed the role of membrane unitary operations in chemical processing, from a very side role to a central role today particularly in some sectors (biotechnology, pharma); on the other hand, the design of countercurrent shell-andtube devices (heat exchanger) provided the basic knowledge for the design of hemodialyzers. Today, the development of microfluidics keeps the connection between these two worlds, providing groundbreaking innovation for devices of a whole brand concept, integrating also other functions of kidney (osmoregulation, exocrine, and endocrine secretion) so far not performed by traditional devices.

16  Chapter 1

List of Symbols a b A B

coefficient introduced in Eq. (20) and defined in Eq. (22) (m5 s) coefficient introduced in Eq. (20) and defined in Eq. (22) (mL)

 mol   L  coefficient introduced in Eq. (21) and defined in Eq. (26)    mol  coefficient introduced in Eq. (21) and defined in Eq. (26)    L 

Am

overall membrane surface area (m2)

C 1

extracellular compartment concentration

C2

intracellular compartment concentration

cB

mol  blood concentration  

cBi

blood concentration at the inlet section 

cBo

 mol   L     mol   L   

 L 

 mol    L  blood concentration at the outlet section  mol   L   

mol    L 

cBo′

outlet blood concentration in the presence of ultrafiltration 

cD

dialysate concentration

 mol   L   

cDi

 mol    dialysate concentration at the inlet section  L 

cm

mean intermembrane concentration 

Δc ΔcBo ΔP

 mol    L  mol  transmembrane concentration difference    L  concentration difference, defined in Eq. (16)  mol   L    transmembrane pressure difference (atm)

ΔPG hydrostatic pressure difference between the glomerular capillaries and the Bowman’s capsule (atm) ΔπG (atm)

osmotic pressure difference between the glomerular capillaries and the Bowman’s capsule

eGFR

estimated glomerular filtration rate 

f

parameter in Eq. (6) and defined in Eq. (7) (dimensionless)

 mL    min 

Artificial kidney: A chemical engineering challenge  17 G

generation rate  mg 

γ

parameter introduced in Eq. (15) (dimensionless)

GFR K

K0

 h   

 mL    min  solute dialyzer clearance  mL   s    mL   diffusive clearance   s   glomerular filtration rate 



mL    min atm 

KGF

hydraulic permeability of the glomerular membrane 

KR

residual renal clearance 

Js

solute molar flux  mol 

Jv

mL    s 

 2  m s volume flux  m   s  

 m2 s    mol 

λ

parameter defined in Eq. (7) 

λ+

coefficient introduced in Eq. (21) and defined in Eq. (23) (s−1)

λ−

coefficient introduced in Eq. (21) and defined in Eq. (23) (s−1)

μ

coefficient introduced in Eq. (26) and defined in Eq. (25) (dimensionless)

Pé

Péclet number (dimensionless)

Pm

membrane permeability 

QB

m   s blood volume flow  mL   s   

 mL    s 

QBi

blood volume flow at the inlet section 

QBo

blood volume flow at the outlet section  mL 

QD

mL  dialysate volume flow    s 

QDi

dialysate volume flow at the inlet section 

QDo

dialysate volume flow at the outlet section  mL 

QF

ultrafiltration volume flow 

 mL    s 

 s     mL    s   s   

18  Chapter 1 QU

ultrafiltration volume flow  mL 

σ

the Staverman reflection coefficient (dimensionless)

Tr

transmittance coefficient (dimensionless)

V1

sum of intracellular and extracellular volume (mL)

V1

extracellular compartment volume (mL)

V2

intracellular compartment volume (mL)

VB

blood compartment volume (mL)

X

mass transfer coefficient 

 s   

mL    s 

List of Acronyms AKI

acute kidney injury

ARF

acute renal failure

BAK

bioartificial kidney

BSA

body surface area

CRRT

continuous renal replacement therapy

CKD

chronic kidney disease

ECF

extracellular fluid

eGFR

estimated glomerular filtration rate

ESRD

end stage renal disease

GFR

glomerular filtration rate

HKAK

hollow fiber artificial kidney

ICF

intracellular fluid

RAD

renal tubule cell-assist device

RRT

renal replacement therapy

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20  Chapter 1 Matovinović, M. S. (2009). Pathophysiology and classification of kidney diseases. Journal of the International Federation of Clinical Chemistry and Laboratory Medicine, 20, 2–11. http://www.ifcc.org. Meyer, T. W., Walther, J. L., Pagtalunan, M. E., Martinez, A. W., Torkamani, A., Fong, P. D., et al. (2005). The clearance of protein-bound solutes by hemofiltration and hemodiafiltration. Kidney International, 68, 867–877. https://doi.org/10.1111/j.1523-1755.2005.00469.x. Molitoris, B. A., & Reilly, E. S. (2016). Quantifying glomerular filtration rates in acute kidney injury: a requirement for translational success. Seminars in Nephrology, 36, 31–41. https://doi.org/10.1016/j. semnephrol.2016.01.008. Nissenson, A. R., Ronco, C., Pergamit, G., Edelstein, M., & Watts, R. (2005). Continuously functioning artificial nephron system: the promise of nanotechnology. Hemodialysis International, 9, 210–217. https://doi. org/10.1111/j.1492-7535.2005.01135.x. Preuss, H. G. (1993). Basics of renal anatomy and physiology. Clinics in Laboratory Medicine, 13, 1–11. Ronco, C., & Clark, W. R. (2018). Haemodialysis membranes. Nature Reviews Nephrology, 14, 394–410. https:// doi.org/10.1038/s41581-018-0002-x. Roy, S., Dubnisheva, A., Eldridge, A., Fleischman, A. J., Goldman, K. G., Humes, H. D., et al. (2009). Silicon nanopore membrane technology for an implantable artificial kidney. In TRANSDUCERS 2009—15th int. conf. solid-state sensors, actuators microsystems. (pp. 755–760). https://doi.org/10.1109/ SENSOR.2009.5285603. Smye, S. W., Evans, J. H. C., Will, E., & Brocklebank, J. T. (1992). Paediatric haemodialysis: estimation of treatment efficiency in the presence of urea rebound. Clinical Physics and Physiological Measurement, 13, 51–62. https://doi.org/10.1088/0143-0815/13/1/005. Spiegel, D. M., Baker, P. L., Babcock, S., Contiguglia, R., & Klein, M. (1995). Hemodialysis urea rebound: the effect of increasing dialysis efficiency. American Journal of Kidney Diseases, 25, 26–29. https://doi. org/10.1016/0272-6386(95)90620-7. Sprenger, K. B. G., Kratz, W., Lewis, A. E., & Stadtmuller, U. (1983). Kinetic modeling of hemodialysis, hemofiltration, and hemodiafiltration. Kidney International, 24, 143–151. https://doi.org/10.1038/ ki.1983.138. Strathmann, H., & Kock, K. (1977). The formation mechanism of phase inversion membranes. Desalination, 21, 241–255. https://doi.org/10.1016/S0011-9164(00)88244-2. Twardowski, Z. J. (2008). History of hemodialyzers’ designs. Hemodialysis International, 12, 173–210. https:// doi.org/10.1111/j.1542-4758.2008.00253.x. Van De Witte, P., Dijkstra, P. J., Van Den Berg, J. W. A., & Feijen, J. (1996). Phase separation processes in polymer solutions in relation to membrane formation. Journal of Membrane Science, 117, 1–31. https://doi. org/10.1016/0376-7388(96)00088-9. Velasco, N. (2013). Present and future dialysis challenges. In Modeling and control of dialysis systems (pp. 1565–1586). https://doi.org/10.1007/978-3-642-27558-6_17. Villarroel, F., Klein, E., & Holland, F. (1977). Solute flux in hemodialysis and hemofiltration membranes. ASAIO Journal, 23, 225–232. Waniewski, J. (2006). Mathematical modeling of fluid and solute transport in hemodialysis and peritoneal dialysis. Journal of Membrane Science, 274, 24–37. Waniewski, J., Lucjanek, P., & Werynski, A. (1993). Alternative descriptions of combined diffusive and convective mass transport in hemodialyzer. Artificial Organs, 17, 3–7. Waniewski, J., Werynski, A., Ahrenholz, P., Lucjanek, P., Judycki, W., & Esther, G. (1991). Theoretical basis and experimental verification of the impact of ultrafiltration on dialyzer clearance. Artificial Organs, 15, 70–77. https://doi.org/10.1111/j.1525-1594.1991.tb00763.x. Ward, R. A. (2005). Dialysis water as a determinant of the adequacy of dialysis. Seminars in Nephrology, 25, 102–111. https://doi.org/10.1016/j.semnephrol.2004.09.017. Ward, R. A. (2009). Worldwide guidelines for the preparation and quality management of dialysis fluid and their implementation. Blood Purification, 27, 2–4. https://doi.org/10.1159/000213489.

CHAPTE R 2

Membrane application for liver support devices Simone Novelli†,§, Cornelius Engelman†,‡, Vincenzo Piemonte⁎ ⁎

Faculty of Engineering, University Campus Biomedico of Rome, Rome, Italy †Liver Failure Group, Institute for Liver and Digestive Health, University College London, Royal Free Campus, London, United Kingdom ‡Section Hepatology, Department of Gastroenterology and Rheumatology, University Hospital Leipzig, Leipzig, Germany §Department of Mechanical and Aerospace Engineering, Sapienza University of Rome, Rome, Italy

1 Introduction The incidence of liver disease is increasing worldwide and about 1 million subjects die from liver failure each year (Jepsen, Gronbaek, & Vilstrup, 2015; Younossi et al., 2016). It is estimated that in the eight leading world economies, more than 100,000 subjects die from liver failure every year, cirrhosis being the most common form of liver disease (Nishikawa & Osaki, 2015). In liver failure, the accumulation of protein-bound toxins and an increased susceptibility to infection lead to multiorgan failure and death (Singer, 2014). Apart from liver transplantation, there is no treatment known to prolong the lives of these subjects. However, limited availability of organs and the costs and the complications associated with liver transplantation precludes this form of treatment for a large majority of patients. A cost effective, safe, and clinically efficacious “Liver Support Device” is, therefore, a vital unmet clinical need. Given the magnitude of the disease burden, it is estimated that this device could potentially prove to be an important clinical intervention, which is likely to benefit a population approaching 5 million subjects. Liver failure can occur either from an injury to an otherwise normal liver (acute liver failure, ALF), or a new injury to an already diseased liver (acute-on-chronic liver failure, ACLF). These conditions have mortality rates of about 50%–80%, and death often related to the development of multiple extrahepatic organ dysfunction and failure. This is elucidated by the observations that a dysregulated inflammatory response appears to be the central process involved in the pathogenesis of this disease leading to multiple organ dysfunction in addition to severe liver dysfunction (Alcalde, Donoso, Carcfa-Diaz, Pascasio, & Narvaez, 1995). It is believed and shown in various animal and human studies that the final common pathway of Membrane Applications in Artificial Organs and Tissue Engineering. https://doi.org/10.1016/B978-0-12-814225-7.00002-4 Copyright © 2020 Elsevier Inc. All rights reserved.

21

22  Chapter 2 a precipitating event such as an infection, bleeding or additional superimposed liver injury is the development and perpetuation of an unregulated systemic inflammatory response. It would therefore seem logical to hypothesize that an artificial liver support (ALS) device which is capable of effectively removing inflammatory and other harmful substances may impact favorably upon the inflammatory response, and provide detoxification function, thus preventing progression from a state of severe liver dysfunction to multiorgan failure and allow time for the liver to regenerate or be stabilized until a suitable organ for liver transplantation becomes available.

2  Liver support devices Very little can be done to increase the number of liver donors, hence it is reasonable to look for artificial means of substitution and/or support, aimed at supporting patients with borderline functional liver cell masses until an organ becomes available for transplantation, or until the liver recovers from an injury (Valentin-Gamazo et al., 2004). An artificial liver could also support patients during periods of functional recovery after marginal liver transplantations and after surgery for traumas or the removal of tumors. Hepatic support therapies provide for two different approaches: one biological and one nonbiological. While those of the biological type use hepatocytes or the entire organ (both of animal and human origin), nonbiological ones use dialytic filtration and adsorption techniques. The metabolites associated with hepatic damage differ from each other by their molecular weights and chemical-physical characteristics, and most of the toxins are bound to albumin (bilirubin, bile, amino acids, and fatty acids). Another portion of the toxins, on the other hand, is hydrosoluble and of low or medium molecular weight, and derive from hepatic damage (ammonium) or from renal damage, and are effectively removed by hemodialysis or hemofiltration. Conventional methods are, however, not effective in removing albumin-bound toxins. It is these substances that play a major role in the pathophysiology of the complications of liver damage, such as hepatic encephalopathy. The hypothesis of intoxication by endogenous toxins is considered the best accredited explanation to explain the events that occur during the course of liver failure that cause a risk to life. In all, 19 different toxins of different molecular weight, free or carried by proteins, are responsible for the various clinical manifestations that characterize this syndrome. Accumulated toxins damage the liver and trigger a vicious mechanism via which the disease is itself self-fed, aggravated, and perpetuated. Increased serum concentration of bilirubin, bile acids, nitric oxide, ammonium, lactates, phenols, aromatic amino acids, and endogenous benzodiazepines correlate with the severity of liver disease. Albumin is the main protein transporter. The concept of dialysis with albumin was developed to remove albumin-related substances; this is a new method that combines the effectiveness of sorbents for removing substances bound to albumin, with the biocompatibility of modern dialysis membranes.

Membrane application for liver support devices  23 Nonbiological artificial support systems, burdened by lower costs and characterized by simpler logistics management, have been studied more in terms of evaluations and clinical applications than bioartificial systems.

3  Development of blood purification systems Table 1 shows the development of blood purification therapies for liver failure. During the 1950s, hemodialysis was introduced as an optimal treatment for renal insufficiency. This removed uremic substances on the basis of diffusion theory, in which a solute moves in a way that depends on a concentration gradient. Hemodialysis, which is used in daily clinical medicine as a reliable treatment for patients with renal insufficiency, was performed in 1958 in patients with hepatic impairment, in an attempt to eliminate the toxins responsible for hepatic encephalopathy. Killey carried out a study of five patients with chronic Table 1: Development of blood purification therapies for liver failure. 1958

Killey

Treatment of a hepatic coma patient using hemodialysis

1958

Schechter

Treatment of a hyperammonemia patient using an ion-exchange column

1958

Lee, Tink

Exchange transfusion

1958

Hori

Cross-hemodialysis using living dogs

1965

Yatzidis

Bilirubin adsorbent using activated charcoal

1965

Eisemann

Initial clinical use of ECLP using a resected porcine liver

1967

Burnell

Cross-hemodialysis between a patient in hepatic coma and a healthy donor

1968

Sabin

Plasma exchange (plasmapheresis)

1970

Abouna

Clinical use of ECLP using a xenogeneic liver

1976

Opolon

Treatment of a patient with fulminant hepatitis using a dialyzer made of a polyacrylonitrile membrane

1976

Knell, Dukes

Control of amino acid imbalance using BCAA solution

1978

Yamazaki

Incorporated system of plasmapheresis and hemodialysis

1980

Brunner

Bioreactor immobilized with hepatic enzymes

1982

Ozawa

Cross-hemodialysis using porcine and baboon livers

1985

Teraoka

XDHP and CPP using porcine liver

1987

Matsumura

Perfusion system of suspended rabbit hepatocytes

1988

Marguilis

Perfusion system of suspended porcine hepatocytes

1992

Yoshiba

Incorporated system of plasmapheresis and hemodiafiltration

1993

Demetriou

Bioreactor system of immobilized porcine hepatocytes

1994

Gerlach

Bioreactor system of immobilized hepatocytes

2000

Stange

Molecular adsorbent recirculating system (MARS)

24  Chapter 2 hepatic failure, four of them showed an improvement in their metabolic encephalopathy, although long-term survival was not achieved. For more aggressive removal of protein molecules, two forms of liver support have been developed: hemoperfusion and plasma perfusion. In 1958, Schechter introduced direct extracorporeal hemoperfusion on matrixes of ion exchange resins, which allowed ammonia to be removed from the blood and a coma reversal in 20% of patients. Also in 1958, Lee and Tink were successful in treating a patient in a hepatic coma using a fresh blood transfusion. This treatment was then replaced by an effective plasma exchange method. A considerable improvement was achieved by using activated carbon as an adsorbent for possible toxins, in the range 500–5000 Da, in whole blood and in patients with impaired liver function. In 1965, Yatzidis and others developed a column of activated carbon for removing serum bilirubin, which is still used for patients with hyperbilirubinemia. Its first major side effects were the loss of platelets and anaphylactic shock. One of the simplest approaches to support a biological liver is cross-circulation, which was discovered in 1967 by Burnell. In this technique, which was performed under general anesthesia, the circulation of a patient with hepatic insufficiency was directly linked to that of a healthy human donor, who suffered severe adverse reactions during the procedure. Plasma exchange, or plasmapheresis, was introduced by Sabin in 1968. This technique separated the plasma using a centrifuge or a membrane, which was then discarded and replaced with an equivalent volume of fresh plasma. Yamazaki has developed a combined technique of plasmapheresis and hemodialysis that has been shown to be effective in reversing hepatic coma and improving coagulation. One problem with this method was the need for a large volume of fresh plasma as a substitute. During the late 1960s, Henderson created a novel approach to the hemofiltration method using a membrane of polysulfone. In 1976, by exploiting the fact that cellulosic membranes were permeable to small water-soluble molecules, Opolon used polyacrylonitrile membranes to improve the transfer of substances, diffusive up to 15,000 Da, thus removing medium weight solutes and small peptides. In a clinical study of 24 patients with acute fulminant viral hepatitis, hemodialysis with these membranes achieved a coma inversion rate of 54%, but survival times were not improved. An important conclusion from these experiments was that substances with a molecular weight lower than 15,000 Da were linked to metabolic encephalopathy. Also in 1976, Knell and Duchi introduced reciprocal dialysis, in which a dialysis fluid with an amino acid concentration identical to that of normal plasma was used to correct an increase in aromatic amino acids and a decrease in branched chain amino acids. Hemodialysis is effective in the removal of small molecules, with molecular weights under 5000 Da, while hemofiltration is particularly effective in removing larger molecules, with molecular weights ranging from 5000 to 10,000 Da. Hence, a combined hemodialysis and hemofiltration therapy

Membrane application for liver support devices  25 has the potential to be an ideal treatment for renal and hepatic impairment. In 1977, Ota developed the hemodiafiltration method (HDF), in which a large amount of water is removed, but a physiological saline solution is provided to regulate the quantity of water removed. In this method, a dialyzer consisting of a hollow fiber membrane with larger pore size than those present in ordinary hemodialysis, facilitates the removal of larger molecules. In an effort to provide a more specific detoxification therapy, immobilized enzyme systems were developed during the 1980s. In these systems, blood is perfused over liver enzymes that are either bound to an insoluble substrate or encapsulated in artificial cells. During the early 1990s, Yoshiba treated 27 patients who had fulminant hepatitis, using plasma exchange in combination with continuous hemodiafiltration, using a high-performance membrane (polymethyl methacrylate). Treatment was considered successful for 15 of the 27 patients treated. Stange and Mitzner introduced a new dialysis method using their originally developed molecular absorbing recirculation system and applied it to 26 patients with hepatorenal syndrome. They reported a significant decrease in serum bilirubin creatinine levels in the treated groups, which led to this procedure becoming one of the most important options for treating liver failure. The purpose of treatment of blood purification, in the case of liver failure, is to remove the toxins that cause coma and cerebral edema. As previously mentioned, the first clinical blood purification applications included hemodialysis, carbon hemoperfusion, plasma exchange, continuous hemodiafiltration, and cryofiltration. Although these treatments improve encephalopathy, patient survival has not yet been achieved. The liver has more than 500 different functions which are difficult to replace by one or even several substitution methods, and this has led to a huge amount of research into liver tissues. The most important functions are: • • • • • • •

bile production and excretion, excretion of bilirubin, cholesterol, hormones, and drugs, metabolism of fats, proteins, and carbohydrates, enzyme activation, storage of glycogen, vitamins, and minerals, synthesis of plasma proteins, such as albumin, and clotting factors, blood detoxification and purification.

In 1958, Hori et al. conducted a study of perfusion treatments using xenogeneic livers and a cross-hemodialysis method consisting of blood circuits in a patient with cirrhosis and four living dogs. The circuits were separated by a semipermeable membrane, through which low- and medium-molecular-weight waste products passed from the patient circuit to the animal circuit, for metabolism by the canine liver and adsorption onto an ion-exchange resin matrix. This treatment was applied to four patients with cirrhosis, one of whom recovered temporarily from a hepatic coma after a significant drop in serum ammonia levels.

26  Chapter 2

4  Artificial devices ALS devices aim to detoxify the patient through dialysis-derived techniques. They are mainly based on the principle of albumin dialysis or plasma separation and filtration (Nevens & Laleman, 2012). They remove both albumin-bound and water-soluble substances without having any synthetic function (Laleman et al., 2006).

4.1 Hemofiltration Hemofiltration is performed using filter membranes with a high cut-off point (about 50,000 Da) made from modified cellulose or polysulfones (Cavagnaro Santa Maria, Roque Espinosa, & Guerra Hernandez, 2018). These membranes can separate natural and toxic substances, within the limits imposed by convective transport phenomena, via membrane exchange. The results obtained, even in this case, are scarce (Li et al., 2017). These procedures have a temporarily beneficial effect on hepatic encephalopathy (perhaps due to correction of toxic levels of certain amino acids) with a return of the coma but do not increase survival rate.

4.2 Hemoperfusion Hemoperfusion is a process of circulating extracorporeal blood through absorbent elements (Fujii et al., 2016) (e.g., activated charcoal) or through highly complex biochemical reactors, which are able to treat specific biological products, such as ammonium. The biggest problem that affects hemoperfusion hemocompatibility of the filtering or absorbing elements, especially toward corpuscular blood components. The problems of hemocompatibility are linked in part to deposition of powders associated with the material itself and in part to the activation of platelets in patients with an already compromised coagulation status. For this reason, thin protective films of polymeric, hemocompatible materials, such as cellulose nitrate, albumin or similar, are used to coat the filter elements, so as to eliminate filter-blood contact (Wang, Liang, Zhou, & Shi, 2017). Direct hemoperfusion is, therefore, a method for adsorbing toxins in the blood using adsorbent materials. Adsorption is a process whereby the solute diffuses into a porous solid and adheres to the internal surfaces. Its first use was reported for treating barbiturate overdoses. Since then, it has been widely used for treating liver failure. Hemoperfusion, performed by extracorporeal blood circulation on nonspecific adsorbent materials (e.g., active carbons), has not had major clinical successes on long-term survival. Charcoal hemoperfusion (Pardo, Lanaux, Davy, & Bandt, 2018) can remove large encephalopathy-associated molecules that would not pass through dialysis or hemofiltration

Membrane application for liver support devices  27 membranes, but nonspecific adsorbent materials can deplete the plasma of biologically important substances (Sahoo & Gurjar, 2016).

4.3  Reactors with immobilized enzymes To give an answer to the problem of specificity in detoxification, specific enzymes (e.g., glucoamylase, cellulase, and peroxidase) were: • • •

attached to hollow fibers or circulated in the closed compartment of the dialysate of an artificial kidney or incorporated into microcapsules exposed to blood

In vitro results have shown evidence of the effectiveness of this approach, but no clinical study has demonstrated enzyme reactors give superior results compared to other dialysis methods: improvements in the mental status of patients in hepatic coma, but no demonstration of increased rates of survival (Wang et al., 2016). It is unclear whether the failure of this technique is due to the inability of specific enzymes to remove all harmful toxins, or whether it evidences the need for something more than detoxification for effective treatment.

4.4  Plasma exchange Plasma exchange, or plasmapheresis, is a method for separating the plasma from the corpuscular elements of blood using hollow fiber filters made of cellulose diacetate and polyethylene membranes. Plasmapheresis was first used for liver failure by Lepore and Martel in 1970. The logic of using plasma exchange for treating liver failure is to remove the toxins and to supply the defective components, such as albumin and coagulation factors. Plasmapheresis appeared to be associated with recovery from a hepatic coma and with a statistically significant improvement in coagulation indices for patients with acute liver failure. However, both the risk of infection and the need for large quantities of expensive fresh frozen plasma have made it difficult to continue plasmapheresis therapy for liver failure in recent years. Controlled studies are needed to establish any beneficial effects of plasma exchange on patient survival.

4.5 Hemodiafiltration HDF is based on the assumption that medium-sized molecules are responsible for hepatic comas in patients with fulminant hepatic failure (FHF) (Locatelli, Violo, Longhi, & Del Vecchio, 2015). This method uses a high-performance membrane, like the wide pore polymethyl methacrylate (PMMA) membranes that were developed in 1986. In retrospective studies, patients have shown complete recovery from a deep coma, and survival in cases of severe FHF (Charitaki, Belman, & Davenport, 2014).

28  Chapter 2

4.6  Plasma exchange and continuous hemodiafiltration To effectively remove medium-molecular-weight substances, such as liver toxins, and minimize the adverse effects associated with plasmapheresis, Nitta developed (Masakane et al., 2015) a slow plasmapheresis treatment combined with a high continuous flow of hemodiafiltration. In a study of five patients with hepatic impairment, they concluded that the adverse effects associated with plasma exchange (such as hypernatremia, metabolic alkalosis, and severe decrease in the osmotic pressure of colloids) could be minimized using plasmapheresis in combination with continuous hemodiafiltration (Fig. 1). In Japan, this is the therapy used most to treat acute liver failure. 35%–40% of patients were able to recover through this treatment, while the rest remained awaiting transplantation.

4.7  Molecular therapy of the absorbent recirculation system The molecular therapy of the absorbent recirculation system (MARS) has proved to be useful in the highly selective removal of water-soluble toxins and of low- and medium-molecularweight substances associated with albumin, through a high flow membrane, such as the polysulfone membrane. MARS is the most widely used ALS as well as the best detoxification

Fig. 1 Flowchart and operating conditions of plasma exchange combined with continuous hemodiafiltration.

Membrane application for liver support devices  29 system studied in cases of liver failure. It was introduced clinically in 1993, and is based on the principle of dialysis with albumin. MARS consists of a standard dialysis machine, which drives the extracorporeal blood circuit, and an additional device to perform and monitor the closed loop with albumin. It consists of two circuits: the blood circuit (primary circuit) and the albumin circuit (secondary circuit) (Jain & Dhawan, 2017) (Fig. 2). 600 mL human albumin 20%

Hemodialysis

Secondary circuit

Not permeable to albumin

Fig. 2 The blood circuit (primary circuit) and the albumin circuit.

Adsorbers

Blood circuit

MARS flux

Patient

Low flux

30  Chapter 2 A highly permeable impregnated albumin dialyzer is used; the closed dialysate circuit was primed with 600 mL, of which 20% was constituted by human serum albumin, and was regenerated through a low-flow dialyzer, a charcoal column, and an anionic exchanger (Novelli, Annesini, et al., 2009). The blood exits the patient through a central venous dialysis catheter. In the primary circuit, the patient’s blood passes through a hemofilter with a cutoff threshold size of less than 60 kDa, thereby maintaining albumin on the blood side. In a secondary circuit, a 20% albumin solution circulates and passes through the hemofilter running counter-current relative to the patient’s blood and acts as a dialysate (Novelli, Rossi, et al., 2009). Toxins in the patient’s blood that dissociate from the albumin bond can pass through the membrane, and follow the concentration gradient to bind to the albumin in the secondary circuit (Annesini, Di Paola, Marrelli, Piemonte, & Turchetti, 2005). This happens according to the theoretical availability gradient of the link. In fact, the albumin present in the dialysate has all sites available, unlike the albumin present in blood (Novelli, Rossi, et al., 2009). The membrane is impermeable to albumin, and only the free toxin fraction can cross it; this is the limiting factor for eliminating compounds with that bind strongly to albumin, like bilirubin. The functional capacity of the albumin, on the dialysate side, which leaves the first dialyzer less with both liver toxins linked to albumin and with soluble solutes in water that are not bound to proteins. In the secondary circuit, the albumin solution first undergoes dialysis using a high flow filter to remove water-soluble toxins via the counter-current motion of fresh bicarbonate-rich dialysate. So the albumin leaving the second dialyzer has a low concentration of uremic toxins, but still has a high concentration of hepatic toxins (Novelli et al., 2005). Afterward, the solution is regenerated by passing it through two adsorbents: an anionic exchange resin and an activated charcoal column. The ion-exchange resin absorbs bile acids while the charcoal column absorbs bilirubin. Usually, adsorbent treatment through carbon cartridges requires the columns to be coated with a synthetic layer to avoid immunological reactions; in this case, the charcoal does not come into contact with the patient’s blood. Passing albumin through these two adsorbent columns allows the regeneration and cleaning of liver toxins. At the time the albumin exits the resin column and reenters the dialysate compartment of the first dialyzer, it should have a lower concentration of both liver-related toxin and uremic solutes. While the bicarbonate-rich dialysate is an open circuit and therefore has a potentially unlimited capacity for dialyzing uremic toxins, the albumin-rich dialysate compartment is closed, and theoretically has an intrinsic absorption limit (Fig. 3). One MARS session lasts for 6–8 h, and after that time the regeneration capacity of the absorbers for albumin decreases significantly. Significant reductions in serum bilirubin, bile acids, ammonia levels, urea, lactate, and creatinine have repeatedly been documented with MARS therapy. Total bilirubin and conjugated bilirubin are lowered, while no change in unconjugated bilirubin levels is observed. MARS does not interfere with valuable molecules, such as albumin, coagulation

Membrane application for liver support devices  31

MARS membrane Blood from patient

Dialysis membrane Albumin dialysate

Hemofiltration/ Hemodialysis circuit

To patient

Anion exchange resin

Activated charcoal

MARS membrane

Blood plasma Albumin with toxins bound Albumin Toxin Binding sites on MARS membrane

Fig. 3 The MARS membrane.

Albumin dialysate

32  Chapter 2 factors, and electrolytes and does not alter the number of blood cells and arterial blood gases and does not generate hemodynamic instability. An improvement in mean arterial pressure, systemic vascular resistance, cardiac output, and cerebral blood flow (due to the reduction of cerebral edema) have repeatedly been demonstrated (Novelli et al., 2008). The clinical outcome depends very much on the etiology of the hepatic failure. Survival rates at the time of detection varied between 60% and 70% in patients with either ALF or A-on-C LF of various etiologies. In one study (Novelli et al., 2002, 2005, 2007) in which MARS therapy was used to treat 34 patients with ALF of various etiologies, the survival rate was 88% at 6 months and 84% at 1 year. The highest incidence of liver recovery was observed in patients with ALF caused by intoxication. The advantage of MARS is that it’s easy to use and costs less when compared to bioartificial devices. Most of the clinical studies conducted with MARS were performed in patients with ACLF, not those with ALF. In one study (Novelli et al., 2007), patients with ACLF with rapidonset-type hepatorenal syndrome were treated with MARS and standard medical therapy (SMT), including hemodiafiltration. MARS treatment significantly improved survival; in fact, mortality rates in patients treated with SMT were 100% on the 7th day compared to 63% for patients treated with MARS. The method has also been shown to be safe and without undesirable side effects. Thus, it was concluded that MARS seemed to be useful as a bridge to liver transplantation. The most recent randomized controlled trial was performed on 24 patients with ACLF with progressive hyperbilirubinemia. MARS has been associated with an improvement in 30-day survival; in addition, renal dysfunction and hepatic encephalopathy were improved in the MARS group. It was concluded that MARS appeared to be effective and safe for the short-term treatment of patients with liver cirrhosis and overlapped with an acute injury associated with progressive hyperbilirubinemia. The only relevant disadvantages derived from the use of MARS are heparinization and collapse, due simply to the fact that the patient is connected to an extracorporeal circuit and thrombocytopenia, which is the back diffusion of charcoal up into the patient’s blood.

4.8 Prometheus The Prometheus system consists of a primary circuit (plasma filter and dialyzer) and a secondary circuit (adsorbent filters for bilirubin removal) and performs a combined removal of toxic molecules linked to albumin, and of water-soluble toxic molecules through the use of an adsorption and fractionated plasma separation system (Fractionated Plasma Separation and Adsorption FPSA) which was developed in 1999 by Falkenhagen et al. (Fig. 4). Unlike MARS, the albumin fraction passes through a polysulfone membrane—an AlbuFlow plasma filter (Fresenius) with a cut-off of about 250 kDa. A specific polysulfone membrane allows the fraction-containing albumin to pass, along with related substances and all of the plasma components weighing less than 250 kDa, in the secondary circuit in which the plasma is

Membrane application for liver support devices  33

Fig. 4 Prometheus system.

purified from toxins linked to albumin by direct adsorption onto special adsorbent supports. Then, in the initial step, the albumin fraction of the blood is selectively filtered through a specific polysulfone and enters a secondary circuit in which this plasma fraction is filtered through two adsorbent columns: the first is a neutral styrenic resin adsorbent, which absorbs mainly cytokines (bilirubin and bile acids) and the second adsorber is an anionic exchanger for removing negatively charged toxins. It is possible to increase the amount of plasma treated regardless of the amount of blood supplied to the circuit. Then the blood is recomposed to the untreated fraction and subsequently, a conventional high flux dialysis is carried out in order to remove the water-soluble substances. The hemofilter is, in fact, impermeable to albumin but permeable to water-soluble toxins. A modified hemodialysis monitor is used to integrate the two circuits. A standard double-lumen dialysis catheter is connected to the extracorporeal blood circuit (primary circuit) and a central venous catheter is used in MARS treatment; the AlbuFlow membrane separates the primary circuit from the secondary, which contains the filtered plasma. As the circuits are independent of the unit, the operator can decide whether to perform a conventional hemodialysis or to associate it with albuminic detoxification.

34  Chapter 2 The fact that these tools for extracorporeal blood purification have the ability to selectively remove certain substances is fundamental for assessing their specificity and effectiveness (Komardina, Yaroustovsky, Abramyan, & Plyushch, 2017; Novelli, Rossi, et al., 2009). Prometheus significantly improves the serum levels of conjugated bilirubin, bile acids, ammonia, cholinesterase, creatinine, urea, and the pH of the blood. With the Prometheus system, both albumin-linked and the water-soluble substances, which accumulate in liver damage, are removed by a single passage of blood through the extracorporeal circuit. The most recent randomized controlled trial conducted on the clinical efficacy of this system was the HELIOS study (Alcalde et al., 1995; Kribben et al., 2012). In this study, 179 patients with chronic acute hepatic insufficiency (AoCLF) were enrolled, 68 of whom were treated with SMT and 77 with FPSA (sessions of 6/8 h for a maximum of 8/11 treatments per patient) with a primary endpoint of patient survival at 28 days and 90 days, independent of transplantation. The results of this study showed survival of 66% vs 63% at 28 days, and 47% vs 38% at 90 days (FPSA vs SMT) which was not statistically significant. However, by analyzing the subgroup of patients with (HRS) type I hepatorenal syndrome, a significantly higher statistically significant survival was observed in the group of patients treated with FPSA. There are few studies in the literature that report the effects of therapy with Prometheus. Krisper, Stadlbauer, and Stauber (2011) compared the effects of MARS and Prometheus on the removal of toxins and observed a greater efficiency of Prometheus relative to MARS. The decrease in blood pressure, the coagulation of the secondary circuit and a slight increase in the number of leukocytes were the only side effects and complications reported for Prometheus therapy. The reduction in blood pressure can be explained by a reduction of intravascular blood volume due to the filling of the extracorporeal circuit. This phenomenon may be more evident in FPSA compared to MARS caused by the filtration of the albumin-rich plasma fraction filling the secondary circuit. Moreover, in Prometheus therapy, a reversible increase in the number of white blood cells probably occurs as a consequence of the blood-membrane interaction. While no decrease in platelet count was observed, coagulation problems were occasionally reported despite the anticoagulant heparin.

5  Bioartificial devices The metabolic complexity of the liver is such that it is very unlikely to be capable of eliminating every single toxin from the body with the same effectiveness as a healthy liver, and this is a serious limitation to the success of any artificial chemical therapy. From the 1960s onward, therefore, the prevailing approach was biochemical. In fact, only another healthy liver can adequately replace a diseased liver. The bioartificial liver (BALs) allows

Membrane application for liver support devices  35 patients to benefit from the combination of active detoxification, intermediate metabolism, and the ability to synthesize and excrete macromolecules that can only be obtained from hepatic cells (hepatocytes). The development of a BAL was a formidable discovery. Unlike the heart, lungs, or kidneys, which serve a primary task, the liver controls almost every aspect of metabolism, blood coagulation, the immune system, endocrine responses, waste removal, and a large part of the physiological processes (Zhang et al., 2018). The BAL is a system of support for liver functions, comprising a biological component (hepatocytes) and an artificial structure that ensures continuous perfusion of the patient’s blood or plasma to the bioreactor. The system is connected to the patient’s venous circulation and generally consists of a plasma separating apparatus, which is then conveyed through the bioreactor containing the hepatic cells, where metabolic exchanges occur. The plasma thus treated is subsequently returned to the patient. If animal cells are used, the procedure is considered as xenotransplantation. BALs are classified according to: - - -

Cellular source Type of hepatocyte culture system Bioreactor type

5.1  Cellular sources A large number of liver cells are required for clinical use of a BAL. In fact, to support a patient with acute failure, about 1010 are needed in a BAL reactor. Many BALs have been developed to accommodate pig hepatocytes, which can be obtained in large quantities, but infections of humans by endogenous pig endovirus (PERV) cells have been reported in vitro. An alternative approach is to use cell lines of immortalized human hepatocytes, which have the necessary functional and survival characteristics. They are genetically modified tumor cells (hepatoblastomas), which allow a good functional differentiation and are able to replicate indefinitely. These cells, the fruits of genetic engineering, are patented, like VitaGen C3A cells by VitaGen Incorporated. Bioreactors with immortalized cells can perform the functions of a human liver and also secrete several coagulation factors and other specific proteins (Zhang et al., 2018). As with normal liver cells, there are factors that inhibit cellular contact. In fact, the hepatocytes are not in mutual physical contact, but are immersed in an extracellular substance; therefore, when the culture volume, the extra-capillary space, is exhausted, the cell replication process that ensures population stability stops spontaneously. These devices are made by injecting 2–5 g of C3A cells (immortalized hepatocytes) into the extra-capillary space of the cartridge. The cells grow for the next 3–4 weeks after which they stop dividing. At this point, the culture remains stable indefinitely. In practice, no deterioration or malfunctions are observed for the following 6–8 months. The metabolic capacity, both in vivo and in vitro, is equivalent to about 200 g

36  Chapter 2 of a normal liver. The most serious problem in the use of C3A cells is the risk of tumors if the cells are able to filter through the microtubule wall in the intra-capillary space. The risk is however very low since C3A cells have surface antigens, which make them recognizable and therefore attackable by the patient’s immune system. Furthermore, the blood is not normally in contact with the cells and the whole system is designed to prevent an infusion of C3A cells, in the presence of fiber rupture, by means of filters. However, the risk-benefit ratio is sufficiently low to ensure wide use of bioreactors with C3A cells. Thus, human embryonic stem cells, liver stem cells, progenitor cells, and trans-differentiated cells from peripheral stem cells are used as cellular sources, and thanks to advances in genetics, more and more hepatocytes can be obtained with improved functionality and a greater potential for proliferation.

5.2  Types of hepatocyte culture systems High-density cultures and high liver functionality are required for a bioreactor. Great efforts have been made to keep vital functional hepatocytes out of the body. Thanks to advances in tissue engineering, the bioreactor mimics the structure of the liver and this can include separate hepatocytes, monolayer cultures, scaffold fixation, hepatocyte aggregates, and culture or culture rotation.

5.3  Types of bioreactor Bioreactors can be classified into four main types (Fig. 5): • • • •

Fiber cable Flat plane and single layer Perfused scaffolds Suspension and encapsulation chambers

Each type of bioreactor has its own advantages and disadvantages, but an ideal device should integrate efficient mass transport, scalability, and maintenance of hepatocyte function.

Hollow fibers

Flat plate and monolayer

Perfused beds/scaffolds

Encapsulation and suspension

Fig. 5 Types of bioreactor.

Membrane application for liver support devices  37 Various requirements have to be met in order to guarantee the full efficiency of a bioreactor. Functional liver cells must be isolated from the external environment and immobilized within a favorable substrate. The substrate should allow the preservation of cell morphology and metabolism throughout the treatment, and the oxygen and nutrients should be accessible to cells in appropriate concentrations. A porous material (not necessarily the substrate) should act as a barrier between the patient’s blood or plasma and the hepatocytes inside the bioreactor, to isolate cells from immune factors (immunoglobulins) and leucocytes, and to avoid immune rejection. Smaller particles, such as toxins and metabolites or synthesized proteins (e.g., albumin and coagulation factors) should be free to cross the barrier. The aim of BAL includes both liver detoxification and synthesis functions, using a combination of physical and chemical procedures, and bioreactors that house the cells. Currently, four BAL configurations are being investigated: hollow fiber cartridges or chambers (ELAD, HepatAssist, MELS) (Patel, Okoronkwo, & Pyrsopoulos, 2018; Wood et al., 1993), monolayer cultures, perfused matrices (BLSS, AMC-BAL) (Patzer et al., 2002) within specific devices and systems based on microencapsulation. A number of BAL devices have been tested clinically, and the characteristics of these systems are provided in Table 2. Important practical issues include the use of whole blood v plasma and the type of anticoagulation regimen. The use of whole blood exhibits the advantage of including oxygen-containing erythrocytes; however, undesirable leukocyte activation and cell damage may arise. In contrast, perfusion of plasma prevents hematopoietic cell injury, but the solubility of oxygen in plasma devoid of oxygen carriers is quite low.

6  Tissue engineering In addition to temporary extracorporeal support, the development of cell-based therapies for liver treatment aimed at the eventual replacement of damaged or diseased tissue is an active area of investigation. In many cases, hepatocytes have been injected into animal hosts and exhibit substantial proliferative capacity as well as the ability to replace diseased tissue and correct metabolic liver deficiencies in those models (Ho et al., 2010; Overturf, Al-Dhalimy, Ou, Finegold, & Grompe, 1997; Rhim, Sandgren, Degen, Palmiter, & Brinster, 1994; Sokhi, Rajvanshi, & Gupta, 2000). However, the clinical efficacy of these procedures is currently limited due in part to technical hurdles in cell delivery and animal models. These limitations might be addressed in part by engineering threedimensional liver tissue ex vivo prior to implantation. Here, we detail the state-of-theart in cell transplantation in the context of applicable animal models as well as in the construction and application of engineered liver tissue.

38  Chapter 2 Table 2: Characteristics of eight bioartificial liver systems. Cell source/ amount

BAL system

Configuration

HepatAssist (Arbios, Waltham, Massachusetts)

Hollow fiber, polysulfone; microcarrier attached Hollow fiber, polysulfone; cellulose acetate; large aggregates Hollow fiber, interwoven, multicompartment; tissue organoids Hollow fiber, cellulose acetate; collagen gel entrapped Radial flow bioreactor; aggregates

Freshly isolated porcine (200–230 g)

I

Nonwoven polyester matrix, spirally wound; aggregates Hollow fiber; collagen entrapped

Freshly isolated porcine (10–14 × 109)

I

Freshly isolated porcine (40–80 g)

I

Hollow fiber; polysulfone; adsorption column

Freshly isolated porcine (10 × 109)

I

ELAD (Vital therapies, San Diego, California) MELS (Charite Virchow, Berlin, Germany) BLSS (Excorp Medical, Oakdale, Minnesota) RFB-BAL (Univ. of Ferrara, Italy)

AMC-BAL (Univ. of Amsterdam, Netherlands) LiverX-2000 (Algenix Inc., Minneapolis, Minnesota) HBAL (Nanjing Univ., Nanjing, China)

Trial phase

Comments

Cryopreserved porcine (7 × 109)

II/III primary endpoint not reached

C3A human cell line (200–400 g)

II/III

Freshly isolated porcine or human (up to 650 g) Freshly isolated porcine (70–120 g)

I/II

Plasma, citrate anticoagulation, 0.15– 0.20 μm pore size, 7 h/ session, daily Plasma, heparin anticoagulation, 0.20 μm pores, continuous up to 107 h Plasma, heparin anticoagulation, 400 kDa cutoff, continuous up to 6 days whole blood, heparin anticoagulation, 100-kDa cutoff. 12 h/session for up to two sessions Plasma, heparin/citrate anticoagulation, 1 μm polyester screen, 6–24 h treatments Plasma, heparin anticoagulation, direct cellplasma contact, up to 18 h/ session for up to two sessions whole blood, heparin anticoagulation

II/II

Plasma, 100 kDa cutoff, one to two 60 h treatments

6.1  Implantable engineered tissue for humanized mouse models Implantation of human engineered liver tissue into animal hosts may also provide an alternative in vivo model systems for human disease. Despite their promise as model systems for investigating human-specific drug responses and infectious diseases with human tropism, current humanized mouse models (e.g., FAH(−/−) and albumin-uPA models detailed above) are limited in that animals must be both immunodeficient and exhibit significant host liver injury. Additionally, the process of human hepatocyte injection, homing to the liver, and expansion can take weeks to months; creating humanized mice using “classic” cell transplantation is therefore tedious and timeconsuming (Wang et al., 2016). As one example of a candidate alternative, a recent study generated humanized mice by implanting hepatocytes and supporting nonparenchymal cells within a three-dimensional hydrogel scaffold into the intraperitoneal space

Membrane application for liver support devices  39 10

(B) Human albumin (mg/mL)

(A)

1

0.1

0.01

0.001 0

2

4 6 8 10 12 14 Weeks after transplantation

16

(C) Primary hepatocytes + supportive stroma

Engineered human hepatograft

Polymer scaffold + optimal porosity + adhesion ligands + degradable linkages

Implantation Engraftment Vascularization

(F)

500

Metabolic Rati (AUC of 4-OHDB/DB)

(E) Human serum Albumin (ng/mL)

(D)

Humanized mouse

400 300 200 100

0.5 0.4 0.3 0.2 0.1 0.0

0 0 4 8 12 16 Day post implantation (IP)

pe ed -ty niz ild ma W hu

Fig. 6 Mice with humanized livers. (A) FAH-positive human hepatocytes injected into the spleen of Fah−/− mice can engraft and integrate in the mouse livers, resulting in mice with “humanized” livers. (B) Human albumin can be detected in the blood these mice. (C) Engineered human liver tissue (“ectopic liver tissue”) could also be constructed ex vivo and then implanted into mice. (D) Ectopic liver tissue can survive and function in mice that are both immunocompromised (left) and immunocompetent (right). (E) Human albumin can be detected in the blood of these mice. (F) Mice with ectopic human liver tissue can be used to identify disproportionate human drug metabolites that would not have been identified by wild-type mice.

40  Chapter 2 of uninjured mice (Wang et al., 2016) (Fig. 6). The engineered human liver tissue synthesized human liver proteins as well as human-specific drug metabolism, drugdrug interaction, and drug-induced hepatocellular toxicity (Chen et al., 2011). The polyethylene glycol (PEG)-based engineered tissue was shown to survive and function within immunocompetent hosts for a period of time after implantation, suggesting that encapsulation of cells in this material system may have the potential to delay immune rejection and enable studies that require both human liver systems and intact immune processes.

6.2  Implantable therapeutic engineered liver tissue The development of implantable engineered hepatic tissue is a promising strategy for the treatment of liver disease due to its potential to mitigate the limitations in current cell transplantation strategies, including inefficient seeding and engraftment, poor long-term hepatocyte survival, a required donor cell repopulation advantage, and the inherent lag phase before clinical benefit is attained (Jain & Dhawan, 2017; Seretis, Seretis, & Liakos, 2014). Implantable engineered hepatic tissues are typically fabricated by immobilizing or encapsulating hepatic cells in biomaterial scaffolds in conjunction with strategies to optimize hepatocyte survival and function, leading to the generation of liver-like tissue in vitro prior to in vivo implantation.

6.3  Design criteria for implantable systems To achieve therapeutic levels of liver function to treat liver failure, the development of engineered hepatic tissues that contain high densities of stable and functional hepatocytes with efficient transport of nutrients and secreted therapeutic factors is necessary. Furthermore, integration of the engineered tissue with the host upon implantation is critical in ensuring its long-term survival. The potential tunability of engineered implantable systems offers attractive prospects for the optimization of hepatocyte survival and function as well as subsequent host integration. Scaffold parameters that are customizable include porosity, mechanical and chemical properties, and three-dimensional architecture. Additionally, relevant microenvironmental cues like paracrine and juxtacrine cellcell interactions, cell-matrix interactions, and soluble factors can be incorporated into implantable engineered hepatic tissue by translating either biological or biomimetic strategies from in vitro culture models so as to recapitulate important aspects of the in vivo hepatocyte microenvironment.

6.4  Natural scaffold chemistry and modifications A wide variety of naturally derived material scaffolds have been explored for liver tissueengineering applications, including materials like collagen, peptides, fibrin, alginate, chitosan,

Membrane application for liver support devices  41 hyaluronic acid, cellulose, decellularized liver matrix, and composites of these. The choice of material determines the physicochemical and biological properties of the scaffold. For example, early efforts in developing implantable hepatic constructs utilized collagen-coated dextran microcarriers that enabled hepatocyte attachment since hepatocytes are known to be anchorage-dependent cells. The intraperitoneal transplantation of these hepatocyte-attached microcarriers resulted in the successful replacement of liver functions in two different rodent models of genetic liver disorders (Demetriou et al., 1986). Subsequently, collagen-coated or peptide-modified cellulose gelatin (Tao, Shaolin, & Yaoting, 2003), and gelatin-chitosan composite (Li, Marchant, Dubnisheva, Roy, & Fissell, 2011) microcarrier chemistries have also been explored for their capacity to promote hepatocyte attachment. On the other hand, materials that are poorly cell adhesive like alginate have been exploited for their utility in promoting hepatocyte-hepatocyte aggregation (e.g., spheroid formation) and hence hepatocyte stabilization within these scaffolds. Collectively, the size of engineered tissues created by these approaches is limited by oxygen and nutrient diffusion to only a few hundred microns in thickness. To address this constraint, recent work has sought to use decellularized whole organ tissue as a matrix for liver tissue engineering (Fig. 7). The decellularization process removes cells from donor tissues but preserves the structural and functional characteristics of much of the tissue microarchitecture and vasculature of the underlying extracellular matrix. Decellularized liver matrices can be seeded with hepatocytes and vascular cells, exhibit liver-specific functions and survive after transplantation into rodents. Future work in this area will likely focus on improving cell seeding protocols, which to date have achieved

Fig. 7 Decellularized liver scaffolds. Representative photographs of ischemic rat livers during decellularization at (A) 0 h, (B) 18 h, (C) 48 h, (D) 52 h, and (E) 72 h. (F) Corrosion cast model of a decellularized liver with portal (red; dark gray in print version) and venous (blue; light gray in print version) vasculature demonstrates that vasculature is intact. (G) SEM images of extracellular matrix within the parenchyma, with hepatocyte-size free spaces. (H) Decellularized livers can be re-cellularized to a total cell mass of 20%–40% of native liver mass. (I) Recellularization with human fetal liver cells can result in selforganization of hepatocytes (green; light gray in print version) and biliary cells (red; dark gray in print version).

42  Chapter 2 only 20%–40% of endogenous liver mass after recellularization, as well as co-seeding with both hepatocytes and the various nonparenchymal cell types found in liver (e.g., stellate cells, Kupffer cells). In general, the advantages of biologically derived materials include their biocompatibility, naturally occurring cell adhesive moieties, and in the case of decellularization, native architectural presentation of extracellular matrix molecules.

7  Conclusion and future trends Although many challenges remain for the improvement of tissue-engineered liver therapies, substantial progress has been made toward achieving a thorough understanding of the necessary components. The parallel development of highly functional in vitro systems, as well as extracorporeal and implantable therapeutic devices is based on contributions from diverse disciplines including regenerative medicine, developmental biology, transplant medicine, and bioengineering. In particular, novel technologies, such as hydrogel chemistries, high-throughput platforms, and microfabrication techniques represent enabling tools for investigating the critical role of the microenvironment in liver function and, subsequently, the development of structurally complex and clinically effective engineered liver systems.

List of Acronyms ACLF acute on-chronic liver failure ALF acute liver failure ALS artificial liver support BAL bioartificial liver CPP cross-plasma perfusion FHF fulminanr hepatic Failure FPSA fractionated plasma separation and adsorption HDF hemodiafiltration HRS hepatorenal syndrome MARS molecular absorbent recirculation system PERV enfogenous pig endovirus PMMA polymethyl methacrylate SMT standard medical therapy XDHP xenogeneic direct hemoperfusion

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44  Chapter 2 Novelli, G., Rossi, M., Morabito, V., Pugliese, F., Ruberto, F., Perrella, S. M., et al. (2008). Pediatric acute liver failure with molecular adsorbent recirculating system treatment. Transplantation Proceedings, 40(6), 1921–1924. Novelli, G., Rossi, M., Pretagostini, R., Poli, L., Novelli, L., Berloco, P., et al. (2002). MARS (Molecular Adsorbent Recirculating System): experience in 34 cases of acute liver failure. Liver, 22(Suppl. 2), 43–47. Novelli, G., Rossi, M., Pugliese, F., Poli, I., Ruberto, E., Martelli, S., et al. (2007). Molecular adsorbents recirculating system treatment in acute-on-chronic hepatitis patients on the transplant waiting list improves model for end-stage liver disease scores. Transplantation Proceedings, 39(6), 1864–1867. Overturf, K., Al-Dhalimy, M., Ou, C. N., Finegold, M., & Grompe, M. (1997). Serial transplantation reveals the stem-cell-like regenerative potential of adult mouse hepatocytes. The American Journal of Pathology, 151(5), 1273–1280. Pardo, M., Lanaux, T., Davy, R., & Bandt, C. (2018). Use of charcoal hemoperfusion and hemodialysis in the treatment of methotrexate toxicosis in a dog. Journal of Veterinary Emergency and Critical Care (San Antonio, Tex), 28(3), 269–273. Patel, P., Okoronkwo, N., & Pyrsopoulos, N. T. (2018). Future approaches and therapeutic modalities for acute liver failure. Clinics in Liver Disease, 22(2), 419–427. Patzer, J. F., 2nd, Mazariegos, G. V., Lopez, R., Bioartificial Liver Program Investigators. (2002). Preclinical evaluation of the Excorp Medical, Inc, bioartificial liver support system. Journal of the American College of Surgeons, 195(3), 299–310. Rhim, J. A., Sandgren, E. P., Degen, J. L., Palmiter, R. D., & Brinster, R. L. (1994). Replacement of diseased mouse liver by hepatic cell transplantation. Science, 263(5150), 1149–1152. Sahoo, J. N., & Gurjar, M. (2016). Should we do early and frequent charcoal hemoperfusion in phenytoin toxicity? Indian Journal of Critical Care Medicine, 20(2), 123–125. Seretis, C., Seretis, F., & Liakos, N. (2014). Multidisciplinary approach to synchronous prostate and rectal cancer: current experience and future challenges. Journal of Clinical Medical Research, 6(3), 157–161. Singer, M. (2014). The role of mitochondrial dysfunction in sepsis-induced multi-organ failure. Virulence, 5(1), 66–72. Sokhi, R. P., Rajvanshi, P., & Gupta, S. (2000). Transplanted reporter cells help in defining onset of hepatocyte proliferation during the life of F344 rats. American Journal of Physiology. Gastrointestinal and Liver Physiology, 279(3), G631–G640. Tao, X., Shaolin, L., & Yaoting, Y. (2003). Preparation and culture of hepatocyte on gelatin microcarriers. Journal of Biomedical Materials Research. Part A, 65(2), 306–310. Valentin-Gamazo, C., Malagó, M., Karliova, M., Lutz, J. T., Frilling, A., Nadalin, S., et al. (2004). Experience after the evaluation of 700 potential donors for living donor liver transplantation in a single center. Liver Transplantation, 10(9), 1087–1096. Wang, T., Liang, F., Zhou, Z., & Shi, L. (2017). A computational model of the hepatic circulation applied to analyze the sensitivity of hepatic venous pressure gradient (HVPG) in liver cirrhosis. Journal of Biomechanics, 65, 23–31. Wang, B., Shangguan, L., Wang, S., Zhang, L., Zhang, W., & Liu, F. (2016). Preparation and application of immobilized enzymatic reactors for consecutive digestion with two enzymes. Journal of Chromatography A, 1477, 22–29. Wood, R. P., Katz, S. M., Ozaki, C. F., Monsour, H. P., Gislason, G. T., Kelly, J. H., et al. (1993). Extracorporeal liver assist device (ELAD): a preliminary report. Transplantation Proceedings, 25(4 Suppl. 3), 53–54. Younossi, Z. M., Koenig, A. B., Abdelatif, D., Fazel, Y., Henry, L., & Wymer, M. (2016). Global epidemiology of nonalcoholic fatty liver disease—meta-analytic assessment of prevalence, incidence, and outcomes. Hepatology, 64(1), 73–84. Zhang, J., Zhao, X., Liang, L., Li, J., Demirci, U., & Wang, S. (2018). A decade of progress in liver regenerative medicine. Biomaterials, 157, 161–176.

CHAPTE R 3

Membrane bioreactors for (bio-)artificial lung M. Pflaum⁎,†, A. Silva Peredo⁎,†, D. Dipresa⁎,†, A. De⁎,†, S. Korossis†,‡,§ ⁎

Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG), Hannover Medical School, Hannover, Germany †Lower Saxony Centre for Biomedical Engineering, Implant Research and Development (NIFE), Hannover, Germany ‡Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG), Hannover Medical School, Biomedical Research in Endstage and Obstructive Lung Disease Hannover (BREATH), German Center for Lung Research (DZL), Hannover, Germany § Cardiopulmonary Regenerative Engineering (CARE) Group, Centre for Biological Engineering (CBE), Wolfson School of Mechanical, Electrical and Manufacturing Engineering, Loughborough University, Loughborough, United Kingdom

1 Introduction Complex living organisms are dependent on the continuous supply of oxygen to drive and sustain crucial metabolic processes. In humans and many other animals, the lung is the responsible organ to provide for the oxygen uptake from the atmosphere into the blood. The erythrocytes then carry the hemoglobin-bound oxygen throughout the body. Equally in importance to blood oxygenation is the removal of carbon dioxide, the metabolic waste product of cellular respiration, from the blood. Both gas species, O2 and CO2, are exchanged in the lung between the blood and the atmosphere by diffusion across the 1  Rp  Rp  

∆Lp

Isotropic expansiona (Evans et al., 1976)

∆P Rp [2(1 − Rp / Rc )]

=K

∆A A0

Bendingb (Evans, 1983)

∆P Rp3 =

B C1

a

Rc is the radius of the spherical outside portion of the cell. C1 = constant.

b

a parachute-like shape when they pass through the narrowest capillaries (3–4 μm) of peripheral circulation and then to recover their biconcavity (Skalak & Branemark, 1969). Due to the high stretch resistance, the cells can tolerate very small area dilation before rupturing. The maximum area strain, measured during elastic area compressibility modulus analysis, ranges from 4% (Evans et al., 1976) to 6.4% (Rand, 1964). The “fluid-mosaic” structure of the membrane and the viscous nature of the cytoplasm make the response to deformation, after the removal of the tension, viscoelastic in nature: the recovery time to the undeformed shape after suction in micropipettes (Hochmuth, Worthy, & Evans, 1979) or pulling with optical tweezers (Dao, Lim, & Suresh, 2003), is 0.1–0.3 s, which multiplied by the shearing modulus, gives a membrane viscosity of 0.6–0.8 μN s m−1. Another relevant mechanical feature of the RBC membrane is the capability to form pores under the action of an external force causing an isotropic tension. The phenomenon can be explained in terms

134  Chapter 6 of an electromechanical instability of the membrane and, more specifically, of the lipid bilayer. Because of its amphiphilic nature, the lipid bilayer possesses liquid-like properties (Israelachvili, 1992; Litster, 1975) and it is remarkably area stretch resistant. If a tension strains the bilayer, the lipid molecules are drawn apart, exposing the hydrophobic tails to aqueous solvent (Fig. 4A). Since this configuration is energetically unstable, the molecules rearrange to expose the polar heads to water, forming a hydrophilic pore that relaxes the tension on the membrane (Fig. 4B). This is a stable configuration, until the equilibrium is not perturbed by changing external conditions (i.e., change tension on the membrane). The formation of stable pores is a very fast process: pores with radius up to 10−1nm appear after few μs (Chang & Reese, 1990; Sowers & Lieber, 1986). While the lipid bilayer is responsible for pore formation, the cytoskeleton seems to influence the pore resealing time after the removal of tension. The process of deformation and pore nucleation in response to the applied tension can occur until the critical value of the area dilation is approached. At this point, the bilayer cannot be further stretched and pores grow irreversibly. Moreover, hyperextensions can also cause the mechanical failure of the cytoskeleton: the resting configuration of spectrin molecules, in fact, is a folded state, but, when stressed, these polypeptides are subjected to reversible morphological changes, by extension or compression. Stretching beyond the maximal extension weakens and eventually breaks the junction points between proteins (Chasis & Mohandas, 1986; Mohandas & Evans, 1994). These phenomena determine the condition of complete hemolysis: irreversible breakdown and cell fragmentation. There are several direct and indirect experimental evidences of pore formation and resealing on biological membranes, when an external stimulus determines a tension increase. Direct observations include: hypotonic lysis and ghost cell preparation techniques (Jay & Rowlands, 1975; Lieber & Steck, 1982; Seeman, 1967), ultrasound-induced cavitation (Mehier-Humbert,

Fig. 4 Structure of membrane pores: (A) Hydrophobic and (B) hydrophilic. Reproduced from Glaser, R. W., Leikin, S. L., Chernomordik, L. V., Pastushenko, V. F., & Sokirko, A. I. (1988). Reversible electrical breakdown of lipid bilayers: formation and evolution of pores. Biochimica et Biophysica Acta (BBA)—Biomembranes, 940(2), 275–287. https://doi.org/10.1016/0005-2736(88)90202-7, with permission.

Numerical prediction of blood damage in membrane-based biomedical assist devices  135 Bettinger, Yan, & Guy, 2005), and electroporation (Chang & Reese, 1990; Kinosita & Tsong, 1977; Needham & Hochmuth, 1989; Riemann, Zimmermann, & Pilwat, 1975; Rols & Teissié, 1992).

3  Phenomenology of hemolysis and blood damage RBCs exposed to fluid flow exhibit a variety of different type of motion and deformation patterns, mainly determined by the loading force from fluid acting on the membrane. In a shear flow, shear stress, which is the product of the shear rate by the viscosity for a Newtonian fluid, will provide the force. When the blood is at rest, cells aggregate in stacks called rouleaux (Schmid-Schonbein, Gaehtgens, & Hirsch, 1968), which disappear as the blood is set to motion. The individual cells tumble as naturally buoyant capsules, but the shape does not deviate from biconcave while the shear stress is below 0.1 Pa. As the fluid force increases, the tumbling is progressively reduced and the cells tend to align with the direction of flow at a shear stress around 0.2 Pa. At 1 Pa, the biconcavity is lost: the cells deform into ellipsoidal shape, with the major axis leaning toward the direction of flow. Due to excess surface area, at these values of shear stress, the deformation occurs at constant cell area. Also, the membrane starts to rotate around the enclosed cytoplasm, in a so-called tank-treading motion (Schmid-Schonbein & Wells, 1969). Fig. 5 shows the tank-treading motion of an RBC membrane suspended in a high-viscosity medium under shear flow. Tank-treading frequency (TTF) appears to be a linear increasing function of the shear rate,

Fig. 5 Schematic representation of RBCs undergoing tank treading of the membrane (highlighted by the movement of dots on the membrane surface). The swinging motion experienced by RBCs at higher shear rates is also represented. The illustration is compared with actual images taken during an experiment with an RBC suspension under shear flow (time sequence of 2 s, shear rate 1.33 s−1, external viscosity 47 cP). Reproduced from Viallat, A., & Abkarian, M. (2014). Red blood cell: from its mechanics to its motion in shear flow. International Journal of Laboratory Hematology, 36(3), 237–243. https:// doi.org/10.1111/ijlh.12233, with permission.

136  Chapter 6 while it is independent of the viscosity of the suspension. The viscosity of external and internal phases, on the other hand, strongly influences the deformation: at constant shear rate and internal phase viscosity, the extent of RBC deformation increases as the external medium becomes more viscous (Fischer, 2004; Morris & Williams, 1979). Fig. 6 shows deformation of RBCs in a rheoscope, suspended in physiological and in high-viscosity media. With the increase of the loading forces, further elongation requires the cell area to increase in order to accommodate the constant volume within the cell. The area strain determines the formation of the pores on the membrane, through which Hb can be released. This is the onset of RBC damage, namely the sublytic damage. In a steady shear flow, the deformation continues until the limit of critical area strain of 6% is reached: this condition is determined by a fluid shear rate of about 42,000 s−1 (Leverett et al., 1972). Once the critical strain is reached, the cell cannot tolerate further area increase and the membrane breaks down, releasing all the remaining Hb into plasma. This is the condition of the catastrophic hemolysis. Remarkably, below the threshold area strain, RBCs can completely recover their biconcavity after the removal of loading. The phenomenology of shearinduced blood damage is summarized in Fig. 7.

Fig. 6 Effect of the external medium viscosity on RBC deformation in autologous plasma. (A) Hematocrit 45%, shear rate 800 s−1; (B) hematocrit 0.1%, medium viscosity 12.9 cP, shear rate 100 s−1; (C) hematocrit 45%, medium viscosity 6.4 cP, shear rate 110 s−1; and (D) hematocrit 25%, medium viscosity 110 cP, shear rate 200 s−1. Reproduced from Fischer, T. (2004). Shape memory of human red blood cells. Biophysical Journal, 86(5), 3304–3313, with permission.

Numerical prediction of blood damage in membrane-based biomedical assist devices  137

Fig. 7 Schematic representation of the phenomenology of shear-induced hemolysis. Reproduced from Vitale, F., Nam, J., Turchetti, L., Behr, M., Raphael, R., Annesini, M. C., et al. (2014). A multiscale, biophysical model of flow-induced red blood cell damage. AIChE Journal, 60(4), 1509–1516, with permission.

4  Quantification of blood damage Several works in the literature have been devoted to the experimental assessment of the blood damage caused by circulation in in vitro setups mimicking the flow configuration of biomedical devices. Most typically, these studies are carried out by flowing blood in mockcirculation loops with the perfused artificial device and measuring the amount of free Hb in the plasma at the end of the experiment. The results are usually reported in terms of scoring indexes relating the plasma-free Hb to the total blood Hb and experimental conditions. The ASTM F1841-97 Standard (ASTM, 1997) defines the index of hemolysis (IH) as: DI (%) =

∆Hb p Hb tot

⋅ 100 =

Hb p − Hb 0p Hb tot

⋅ 100

(1)

where Hbp0, Hbp are the initial and final plasma-free Hb concentrations and Hbtot is the whole blood Hb concentration. For healthy humans, Hbtot ranges from 14 to 16 g dL−1, while Hbp is in the range 0.64−6.4mg dL−1 (Chen, Piknova, Pittman, Schechter, & Popel, 2008; Jeffers, Gladwin, & Kim-Shapiro, 2006). The ASTM Standard defines two other formulations of the index. The normalized index of hemolysis (NIH) is equivalent to the Hb released per pass of blood volume through a blood pumping device: ∆Hb p V NIH (g / 100 L ) = (1 − Hkt ) (2) ∆t Q where V is the volume of blood pumped through the circuit, Q is the pump flow rate, Hkt is the hematocrit, and Δt is the duration of the test.

138  Chapter 6 The modified IH corresponds to NIH normalized by Hbtot in the volume of blood pumped by the device: NIH MIH = (3) Hb tot In most of experimental works in the literature, results are reported according to the ASTM indexes. However, in some works, it is possible to find other formulations. Some authors (Alkhamis, Beissinger, & Chediak, 1990; Beissinger & Williams, 1984; Kameneva et al., 2004; Sharp & Mohammad, 1998) report measures of Hbp during experiments. For example, in some tests in capillary tubes, hemolysis is quantified with the following index (Beissinger & Laugel, 1987): ∆Hb p H= (4) Hb 0 p

where Hbp0 is the plasma-free Hb concentration before shearing. In this latter formulation, the index is greater than 1 and more strongly affected by inaccuracy in the measurements, due to the low values of Hbp0.

5  Experimental data on blood damage A summary of several experimental hemolysis data available in the literature is reported in Table 4. There is a great deal of variability among the different works, first of all for the type of blood used in the experiments: the tests were performed on human, porcine, bovine, and ovine blood. The use of human blood, even though preferable for directly usable results, poses issues in terms of safety of handling, availability, and costs. However, the morphological, rheological, and mechanical properties can change significantly from species to species: for instance, if compared to human RBCs, bovine and ovine RBCs are smaller and less deformable. Furthermore, bovine RBCs seem to be less susceptible to blood damage, whereas porcine RBCs appear to be more fragile. Thus, porcine blood is commonly preferred to bovine in hemolysis experiments (Goubergrits & Affeld, 2004; Paul et al., 2003). Moreover, also the hematocrit and the Hb content of blood can vary from different species. Comparative analyses of RBC deformability and hematological parameters in different animal species can be found in the literature (Smith, Mohandas, & Shohet, 1979; Windberger, Bartholovitsch, Plasenzotti, Korak, & Heinze, 2003). Thus, because of the different origin of blood samples, the experimental results are not directly comparable. Other sources of variability of the experimental data available are: the type of shearing device used in the tests, type of loading and exposure time. In hemolysis tests in catheters and cannulae (Beissinger & Laugel, 1987; De Wachter & Verdonck, 2002; Sharp & Mohammad, 1998), the investigated ranges of τ vary from 1 to about 100 Pa, exposure times vary from few seconds to 6 h, the blood is sheared only by flowing it into the cannula (Beissinger & Laugel, 1987; Sharp & Mohammad, 1998) or within a closed loop comprising the cannula and the

Table 4: Summary of experimental works on hemolysis in the literature Blood

Temperature (°C)

μina (cP)

1 μout (cP)

ηa

Range of loading

Range of t

Data type

Ref.

High-stress range Couette viscometer Human

–

6.4

–

50–400 Pa

2 min

Hbp/Hbtot (%)

Leverett et al. (1972)

3.0 2.0–2.6

1.83– 0.61 2.13 3.2–2.46

Porcine Human

30 37

6.4 6.4

40–600 Pa 120–535 Pa

IH (%) IH (%)

Heuser and Opitz (1980) Wurzinger et al. (1986)

6.4

5

1.28

30–450 Pa

3.6

–

30–320 Pa

IH = ΔHbp/ Hbtot (%) IH (%)

Paul et al. (2003)

–

10−3–1 s 1.4 × 10−2–0.7 s 2.5 × 10−2–1.238 s 3 × 10−2–1.5

Porcine

25

Ovine

25

Zhang et al. (2011)

Hbp/Hbtot (%)

Morris and Williams (1979)

128–516 × 10−3 ms

IH (%)

Pohl, Wendt, Koch, and Vlastos (2000)

6 h 6 h

MIH; Hbp; LDH Hbp

De Wachter and Verdonck (2002) Sharp and Mohammad (1998)

–

H = ΔHbp/Hbp0

Beissinger and Laugel (1987)

2 min

Hbp; [ADP]

Alkhamis et al. (1990)

Cone and plate viscometer Human

25

6.4

3.0–20

2.13– 0.32

50–200 Pa

5 min

Parallel plate viscometer Calf

25

–

4.0

33–331 Pa Cylindrical ducts

Calf Bovine

36 22

– –

– –

– –

250−300 mL min−1 18–260 Pa Low-stress range Cylindrical ducts

Human

23

6.4

2.7

2.37

1–20 Pa Cone and plate viscometer

Human

25

6.4

3.3

1.94

≤20 Pa

Continued

Table 4  Summary of experimental works on hemolysis in the literature—cont’d Blood

Temperature (°C)

μina (cP)

1 μout (cP)

ηa

Range of loading

Range of t

Data type

Ref.

Hbp

Beissinger and Williams (1984)

15 min

Hbp

Bakir et al. (2007)

7.2 × 103 s

HR = (Vp/Vtot) ⋅Hbp/Hbtot

Hashimoto (1989)

Cone and plate vs. parallel plate viscometers Human

25

6.4

3.5

1.83

 N2 and depends on the molecular structure of the PFC (Riess, 2001). At 37°C and 760 mmHg, CO2 for medically used PFCs is about 40–50 vol% as compared to 2.5 vol% for water (Lowe, 2003). According to the literature, linear alkanes such as perfluorooctylbromide dissolve more oxygen than cycloalkanes such as perfluorodecalin (Fig. 8B) (Riess, 2001). In contrast to HBOCs, oxygen dissolving in PFOCs is not severely influenced by temperature (temperature-dependency of kh is negligible). This makes them ideal candidates for oxygen

Artificial oxygen carriers  205

Fig. 8 Properties of PFOCs. The linear oxygen “binding” curve (A) and the structures of two broadly used PFCs, perfluorodecalin and perfluorooctylbromide (B), are shown. Furthermore, the different oxygen extraction rates of HBOCs and PFOCs are illustrated (C) combining both binding curves (Figs. 7A and 8A). Data of oxygen transport capacities of relevant PFOCs emulsions obtained from Riess (2005) (Oxygent and Perftoran) and Wrobeln, Laudien, et al. (2017) (albumin-derived PFC nanocapsules).

206  Chapter 8 delivery in hypothermic organ preservation such as hypothermic machine perfusion and static cold storage (Hosgood & Nicholson, 2010). Importantly, besides oxygen solubility oxygen releasing capacity matters even more. As illustrated in Fig. 8C, blood and HBOCs display an oxygen extraction rate of only about 25%. In contrast, PFOCs release up to 92% of the dissolved oxygen (Wrobeln, Laudien, et al., 2017). PFCs have a long history in tissue engineering and have been used in different ways such as to supplement cell culture medium in monolayer culture or bioreactors/the perfusion medium in organ preservation as well as incorporation into scaffolds (see Section 3.2.3). 3.2.2 Limitations Deviated from the oxygen binding curve (Fig. 8A), PFCs are not very effective in oxygen delivery at pO2 of 100 mmHg (normal air). Thus, to take maximal advantage of PFCs, partial pressure of oxygen in the environment should be increased in vivo by ventilating the patient with oxygen, or in vitro, by oxygenating perfusion solutions and cell culture media (Fig. 8C) (Riess, 2005; Wrobeln, Schluter, et al., 2017). Other than HBOCs, PFCs are fully synthetic molecules and, in consequence, their supply does not depend on blood or animal availability. In contrast to HBOCs, perfluorocarbons (at least the tested perfluorooctylbromide) did not show any interference with standard laboratory or cell culture tests (Ma et al., 1997). However, when perfluorocarbons are mixed with blood differentiation of physiologically relevant hemoglobin forms (oxy-/carboxy- and met-hemoglobin) can be affected by perfluorocarbons (dependent on the analyzer used) while total hemoglobin concentration was unswayed (Shepherd & Steinke, 1998). This is more relevant for use of PFOCs in patients but underlines that PFOCs should not be used simultaneously with HBOCs or blood in cell culture media. 3.2.3  Stage of development of relevant PFOCs Clinical data/clinical trials

There have been numerous clinical studies with different PFC-based emulsions during recent years. However, these were mainly aiming to implement such products as blood substitutes, e.g., in hemorrhagic shock or different types of surgery (Castro & Briceno, 2010). The most famous PFOCs (Fig. 8A) are (i) Oxygent (consisting of 58% perfluorooctylbromide combined with the emulsifiers 2% perfluorodecyl bromide + egg yolk phospholipids, see Section 3.2.1) and (ii) Perftoran (14% perfluorodecalin combined with 6% perfluoromethylcyclohexylpiperidine + Pluronic F68 + egg yolk phospholipids as emulsifiers) (Castro & Briceno, 2010). Development/investigation of Oxygent was discontinued in 2001 after some problematic phase III studies revealing increased stroke risk and neurological side-effects that could not be confirmed in subsequent posthoc analysis (Cohn & Cushing, 2009; Kocian & Spahn, 2008). This is why Oxygent has now been resumed by a Chinese

Artificial oxygen carriers  207 sponsor and just recently been approved for clinical studies in China (Liu, 2017). Perftoran still is the only PFOC that is approved for human clinical use in Russia, Mexico, Kazakhstan, Kyrgyzstan, and Ukraine for treatment of acute blood loss, ischemia of brain/limbs, anemia, and wound healing (Castro & Briceno, 2010). Exploitation of PFOCs in organ preservation prior to transplantation is most advanced for pancreas and islet transplantation, where it is nowadays part of clinical routine (Matsumoto, 2005; Ricordi et al., 2003). In this context, clinicians most often use the static two-layer model, where the organ is placed between one layer of liquid PFC and another layer of aqueous preservation solution (Matsumoto, 2005). Pipeline (preclinical data)

There are scores of preclinical studies on preservation of organs (Zhang & Barralet, 2017), e.g., heart (Kuzmiak-Glancy et al., 2018; Wrobeln, Schluter, et al., 2017), kidney (Hosgood, van Heurn, & Nicholson, 2015), lung (Forgiarini Junior et al., 2013), liver (Moolman, Rolfes, van der Merwe, & Focke, 2004; Okumura et al., 2017), small bowel (Tsujimura et al., 2004), and again, pancreas/islets (Brandhorst et al., 2013) in species such as rats, rabbits, dogs, pigs, and humans; most of them showed improvement (longer preservation) in the presence of PFCs, but some not (Hosgood & Nicholson, 2010; Zhang & Barralet, 2017). Instead of a universal enhancement, it seems to depend on the species and organ (Hosgood & Nicholson, 2010). In these studies, many different PFC-emulsions, among them Oxygent and Perftoran, were tested in several preservation techniques. Mainly, static cold storage techniques were combined with PFOCs, which can be further differentiated into the one- and the two-layer methods. Meanwhile PFOCs were also of benefit in newer techniques such as hypothermic machine perfusion (Hosgood & Nicholson, 2010). As mentioned in Section 3.2.1, in such hypothermal preservation settings, PFOCs are of special benefit because dissolving and release of oxygen is less markedly influenced than in HBOCs (Hosgood & Nicholson, 2010). In recent years the portfolio of organ preservation was expanded to normothermic machine perfusion. In this context, especially the quality of kidneys could be improved in the presence of PFCs (Hosgood et al., 2015). Furthermore, the use of PFCs in a renal assist device, which is a blood purification system designed to treat sepsis-associated acute kidney injury or multi organ failure, is discussed (Cantaluppi et al., 2018). PFCs reduced damage and improved metabolic function of renal tubular epithelial cells in a renal assist device model (a polysulfone hollow fiber) and promoted differentiation of those cells into CD133+ renal progenitor cells (Cantaluppi et al., 2018). PFOCs were also successfully used as supplemental in cell culture medium; namely in microbial, animal, and plant cell cultures (2D and 3D), e.g., enhancing biomass, boosting yields of commercially relevant cellular products, or conserving expression of totipotency of the cells (Lowe, 2002). In tissue engineering, another special aim is to improve differentiation of stem cells into various target cells by providing an oxygenated environment. This can be accomplished by perfusing

208  Chapter 8 cells seeded on scaffolds with PFOC-supplemented cell culture medium. For example, neonatal rat heart cells cultured on a highly porous elastomer scaffold contained higher amounts of desoxyribonucleic acid, cardiac markers, and showed better contractile properties in the presence of Oxygent (Iyer, Radisic, Cannizzaro, & Vunjak-Novakovic, 2007; Radisic et al., 2007). Another approach in 3D cell culture is to bioprint cell-containing hydrogels layer by layer to form cellular constructs with predefined morphology (Blaeser et al., 2013). Larger hydrogel structures are difficult to form as multilayer hydrogels easily collapse under their own weight. Because of their high buoyant density, PFCs support the sensitive hydrogel structures by floating, thus permitting construction of very large cellularized alginate tubes of 3 cm length and 5 mm diameter (Blaeser et al., 2013). Additionally, oxygenation and evacuation of CO2 were improved, too. The same approach was also successful using fibrous polylactide scaffolds seeded with chondrocytes (Pilarek et al., 2014). Instead of supporting hydrogels mechanically by building a flexible basement, PFCs can also be conjugated to hydrogel-forming structures such as methacrylamide chitosan and thus become a direct part of the hydrogel. Differentiation of neural stem/progenitor cells in such fluorinated methacrylamide chitosan gels was more distinct than in nonfluorinated control gels (Li, Wijekoon, & Leipzig, 2014). Similarly, perfluorotributylamine-emulsion-enriched fibrin-hydrogels promoted cell adhesion, proliferation, and migration of Schwann cells (Ma et al., 2013). Importantly, in human islet culture based on a fibrin matrix, Maillard and colleagues demonstrated that positive effects of PFCs might depend on their galenic status (Maillard et al., 2011). Improvement of islet function and viability was observed only when emulsified perfluorodecalin was used and not in the presence of the pure unprocessed liquid (Maillard et al., 2011). In contrast, co-encapsulation of a perfluorotributylamine emulsion with mouse ßTC-tet insulinoma cells in calcium alginate beads was shown to have no significant effect on cellular growth and function (Goh, Gross, Simpson, & Sambanis, 2010).

4  Conclusions and future trends Without the implementation of AOCs development, modern tissue engineering would have been far less successful. While PFOC-supported pancreas/islet preservation prior to transplantation has already reached the clinic, other organs will follow soon. A clinical trial with Hemo2life to improve kidney quality prior to transplantation was completed just recently. Additionally, numerous preclinical approaches improving engineering of many different cells/tissue—among them cardiac and neuronal tissue—in the presence of AOCs developed. In parallel perfusion techniques advanced and probably within the next years, techniques such as normothermic perfusion in combination with AOCs will be investigated in clinical studies. Of course, the field of AOCs also moves on. Novel ideas such as cobaltporphyrins (Shen et al., 2016) and cobalt-replaced myoglobin (Neya, Yonetani, & Kawaguchi, 2014) will enrich the portfolio of future AOCs.

Artificial oxygen carriers  209

Conflict of interest The author has disclosed no conflicts of interest.

Acronyms ADP adenosine diphosphate AOCs artificial oxygen carriers ATP adenosine triphosphate CaO2 oxygen content of arterial blood cytochrome c small hemeprotein associated with the inner mitochondrial membrane Hb actual hemoglobin concentration HBOCs hemoglobin-based artificial oxygen carriers HIF hypoxia-inducible factor OCR oxygen consumption rate p50 oxygen affinity of hemoglobin PFOCs perfluorocarbon-based artificial oxygen carriers RONS reactive oxygen and nitrogen species ROS reactive oxygen species SO2 actual degree of oxygen saturation of hemoglobin

Symbols ѴD amount of diffusing oxygen D diffusion constant Δc difference of oxygen concentration Δx distance to be surmounted kb Boltzmann’s constant T temperature r particle radius Ŋ viscosity of the medium P pressure kDa kilodalton C solubility of the gas kh Henry coefficient pgas partial pressure of the gas Km Michaelis constant

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CHAPTE R 9

Membrane bioreactors for digestive system to study drugs absorption and bioavailability Pompa Marcello⁎, Mauro Capocelli†, Vincenzo Piemonte⁎ ⁎

Unit of Chemical-physics Fundamentals in Chemical Engineering, Department of Engineering, University Campus Biomedico of Rome, Rome, Italy †Unit of Process Engineering, Department of Engineering, University Campus Biomedico of Rome, Rome, Italy

1 Introduction The gastro intestinal system is responsible for the digestion and absorption of nutrients orally administrated in the organism (Elashoff, Reedy, & Meyer, 1982; Hirtz, 1985; Prabu, Suriyaprakash, Ruckmani, & Thirumurugan, 2015). It is composed of the mouth, stomach, and intestine (duodenum, jejunum, ileum, and colon). Enzymatic digestion reactions take place both in the stomach and the duodenum: in the first case enzymes and gastric juices are produced in the stomach epithelium, in the second case the synergic interaction between liver and pancreas produces enzymes that can digest the chyme in output from the stomach (Minekus, 2015). Drugs absorption is a complex phenomenon that starts from the jejunum and ends in the colon, through the passage (in counter current flux) of the blood (Guytin & Hall, 2006; Takahashi, Washio, Suzuki, Igeta, & Yamashita, 2010; Wang et al., 2010). This behavior is the most interesting during drugs design and development phases. The absorption of a drug is given by the synergic interaction between the active transport and the diffusion through the intestinal wall, but the emptying of the stomach determines sensitively the absorption time and the final bioavailability (Ferrua & Paul Singh, 2015). The emptying time of the stomach is controlled by the stomach state. In the digestion phase, the emptying is controlled by the duodenum. Having the function to empty the stomach only at the end of the digestion cycle, the duodenum is filled by chyme until it reaches its maximum volume, duodenum stops the emptying of stomach. It starts the digestion of nutrients and, in the end, it empties itself to the jejunum while the stomach fills it again (Ferrua & Paul Singh, 2015; Ferrua & Singh, 2010; Mudie, Amidon, & Amidon, Membrane Applications in Artificial Organs and Tissue Engineering. https://doi.org/10.1016/B978-0-12-814225-7.00009-7 Copyright © 2020 Elsevier Inc. All rights reserved.

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216  Chapter 9 2010; Pal et al., 2004; Sjödin, Visser, & Al-Saffar, 2011). The intestinal tract motility, like the stomach, is influenced by its state: in the fasting phase the motility is different from the digestion phase (Thuenemann, Mandalari, Rich, & Faulks, 2015). Drug absorption depends on its chemical-physical property: drugs in capsule, for example, need to be disintegrated before entering in solution (Mudie et al., 2010). The intestinal membrane can transport drugs for passive diffusion, facilitated diffusion, pinocytosis, or active transport (Caillè et al., 1989; Clements, Heading, Nimmo, & Prescott, 1978; Ferrua & Singh, 2010; Takahashi et al., 2010; Thuenemann et al., 2015). Passive diffusion starts with a difference in concentration between blood and intestine. The key variables are the liposolubility of the drug and its molecular dimension. Fat-soluble drugs, with small molecular dimension, spread very fast through the membrane (Hirtz, 1985; Persky & Pollak, 2013; Takahashi et al., 2010). Facilitated diffusion is similar to passive diffusion, but it is faster. The transporting molecule of the membrane is linked reversibly with the substrate external at the membrane. So, this process creates a carrier-substrate structure that allows the substrate to spread inside the membrane very quickly. The facilitated diffusion works without the use of energy (ATP) (Persky & Pollak, 2013; Prabu et al., 2015). Active transport is selective, using energy, and allowing transport against the concentration gradient. The drug to use this kind of absorption needs to have an anionic structure or to be similar to vitamins in structure. This kind of absorption, furthermore, exists only in a few parts of the small intestine (Persky & Pollak, 2013; Prabu et al., 2015). Pinocytosis allows the cell to incorporate liquids or solid particles: the membrane cell forms a vesicle that, detaching, moves inside the cell itself. This phenomenon uses energy too. The organism uses pinocytosis for drugs with structure similar to a protein (Persky & Pollak, 2013; Prabu et al., 2015). As a consequence, the absorption of a drug is a complex phenomenon to be studied. The fundamental parameters used in pharmacokinetics are: absorption time, maximum drug concentration in the blood, and elimination time (Persky & Pollak, 2013; Prabu et al., 2015). The term bioavailability is used both for the amount of drug administered in the body, and for the amount of drug available after administration (the drug available for absorption). The bioavailability is the measure of the therapeutic activity (Persky & Pollak, 2013; Prabu et al., 2015). A drug needs to be inside a therapeutic range to work properly; this therapeutic range refers to the drug concentration in blood required to achieve the therapeutic effect with minimal toxicity. The oral administration does not give the possibility to stay in the therapeutic range a lot of the time; for this reason we try to make a new gastro intestinal model that helps to improve the oral administration of an existing drug or to provide a useful tool to rationally design new ones (Lamberti et al., 2016; Persky & Pollak, 2013; Prabu et al., 2015; Sjödin et al., 2011).

Membrane bioreactors for digestive system to study drugs absorption and bioavailability  217 In summary, the bioavailability of a drug taken by oral administration in fasting phase depends on: - - -

kind of formulation; kind of molecule; subject health.

If the drug is oral administrated in digestive phase (with food), the main variables are: - pH of the compartment (stomach, duodenum, etc.); - motility; - chemical reactions between two or more drugs; - biochemical reaction of drug during its transport in the apparatus; - bacteria flora. As previously stated, drug absorption depends also on its chemical-physical property: drugs in capsule, for example, need to be dissolved before entering in solution (Mudie et al., 2010). On the other hand, the intestinal membrane can transport drugs for passive diffusion, facilitated diffusion, pinocytosis, or active transport (Caillè et al., 1989; Clements et al., 1978; Ferrua & Singh, 2010; Takahashi et al., 2010; Thuenemann et al., 2015). Considering these and many other factors described in the next paragraphs, we try to reproduce the entire absorption process through mathematical models mainly based on the use of membrane reactors. All the mathematical models available in literature refer to correlations adapted by experimental tests. Single compartment model describes the organism with a single compartment: the absorption, metabolization, and elimination of the drug are represented by material flux in input and output from the compartment. Diffusion through the membrane and the drug distribution in the tissues are not considered. The model, although reductive, is used in clinical practice (Persky & Pollak, 2013; Prabu et al., 2015). A two-compartment model divides the organism in two parts: in this way, it considers the emptying speed and the drug distribution (Persky & Pollak, 2013; Prabu et al., 2015). To the best of our knowledge, there are no models based on the theory of unitary operations, capable of simulating the functioning of the gastro-intestinal tract reproducing, in a rigorous way, the physiology of the tract. The only reactors reproduced by the available models are perfectly mixed and produce response functions to be experimentally trained. Our approach is to insert models typical of the chemical and process engineering that introduce real kinetic parameters. Once the parameters have been evaluated, the model provides the engineer/ researcher with both a predictive tool and a process analysis tool (to produce sensitivity analysis and simulate data obtained both in the laboratory and in vivo). In fact, this model is inspired by the in vitro TNO gastro intestinal model (TIM). This device is composed of several compartments (linked by valves) and a computer to control the release of fluids

218  Chapter 9 P

A

S

I Q B

C

M

N

J

Q

P

Q L

P

E

R

K

D M

F

G

K

P

L

N

H

Fig. 1 Scheme of the TNO gastro intestinal model (TIM). Schematic presentation of TIM-1, equipped with filters to study the bio-accessibility of lipids. A. gastric compartment; B. pyloric sphincter; C. duodenal compartment; D. peristaltic valve; E. jejunal compartment; F. peristaltic valve; G. ileal compartment; H. ileal-cecal valve; I. gastric secretion; J. duodenal secretion; K. bicarbonate secretion; L. prefilter; M. filtration system; N. filtrate with bio-accessible fraction; O. hollow fiber system (cross section); P. pH electrodes; Q. level sensors; R. temperature sensors; S. pressure sensor.

(gastric juice, pancreatic juice, etc.) and enzyme. The purpose of the TNO gastro intestinal model (TIM) is to emulate the dynamism and movements of the digestive process. In Fig. 1 is shown the most used configuration of TNO gastro intestinal model (TIM): TNO TIM 1. In this configuration, the device has four compartments (stomach, duodenum, jejunum, and ileum) linked by valves. In order to emulate the absorption in the intestine, TNO TIM 1 uses a hollow fiber reactor (composed of a filter for the nonhydrophobic elements) and a dialysis reactor (composed of a membrane for hydrophobic elements). The next section reports the physiology and anatomy of the whole digestive system. Successively, the novel mathematical model is presented and validated through comparison with literature data (obtained in vivo).

2  Anatomy of the GI tract The gastro intestinal tract is composed of different organs. In order to understand the new model of the gastro intestinal system, the most important characteristics of mouth, stomach, duodenum, jejunum, ileum, and colon are briefly described in the following paragraphs

Membrane bioreactors for digestive system to study drugs absorption and bioavailability  219 (Guytin & Hall, 2006). The mouth is the only part of the gastro intestinal tract with a bone skeleton. It has 32 permanent teeth that are able to chew and disintegrate the food. The salivary glands, to lubricate the bolus to the stomach, produce 0.5 L per day of saliva (without stimulation). The enzymes in the saliva are: - - -

α-amilase, to start digestion of carbohydrates; acid lipase, to start digestion of triglycerides (this enzyme active is function in the stomach, with an acid pH); ribonuclease, to start digestion of RNA.

A parasympathetic system controls the secretion of saliva, stimulating the salivary glands. Salivary secretion starts with a tactile, olfactory, or visual stimulation. The stomach (Fig. 2) is divided into four different anatomic zones: - cardia; - fundus; - body; - pylorus.

Fig. 2 Stomach described in its anatomical parts.

220  Chapter 9 The stomach is about 25 cm long, 12 cm wide, and 8 cm thick. With these dimensions, it could stock about 1.5 L of bolus. The proximal portion is where input bolus is collected; while the distal portion is where we can see the muscular contraction, useful for disintegrating the bolus. In this last section, muscle tone enables stomach motility. The bolus is moved through peristaltic waves, coordinating with the pylorus (the valve of the stomach). Gastric cells produce about 2 L of gastric juice (HCl), which causes the pH to oscillate from 1 to 7 (in function of the digestion phase). The production of HCl could be increased by stimulation of membrane receptors (with gastrin, for example) or stamina release. The somatostatin is the inhibitor of HCl and it is released by D cells of the stomach and duodenum. Pepsinogen is the most important enzyme produced by the stomach; activated in pepsin, it works at pH between 1.8 and 3.5, in order to digest protein. Over 3.5 pH, there is a reversible inactivation of the enzyme and over 7 pH an irreversible inactivation of the enzyme. The intestine (Fig. 3) is divided into: (i) Duodenum; (ii) Jejunum; (iii) Ileum; and (iv) Colon. The duodenum is located after the stomach, before the jejunum. The pancreatic and hepatic secretions converge in the duodenum through major duodenal papilla. The duodenum is about 20/25 cm long with a 2.5 cm diameter. In input it has in total about 1 mL/h of pancreatic juice and bile. The enzymes inside the juice are: - - -

trypsin, chymotrypsin, and aminopeptidase for protein digestion; pancreatic amylase for the digestion of the starch, glycogen, and other carbohydrates; pancreatic lipase, cholesterol esterase, and phospholipase for fat digestion. Liver

Stomach

Gallbladder Pylorus Duodenum

Pancreas Left colic flexure

Right colic flexure

Transverse colon

Duodeno jejunal junction Ascending colon

Jejunum

Ileocecal junction

Descending colon

Ileum Cecum

Sigmoid colon

Appendix Rectum Anal canal

Fig. 3 Intestine described in its anatomical parts.

Membrane bioreactors for digestive system to study drugs absorption and bioavailability  221 The bile has the objective to ease fat digestion and absorption. Duodenum pH is about 5.0 and it is determined by the presence of sodium bicarbonate in the juices. The sodium bicarbonate neutralizes the gastric juice, avoiding damage to the intestine. The jejunum is the second part of the small intestine. It is located after the duodenum and before the ileum. It is about 2.5 m long with a 2.5 cm diameter. Its internal surface is covered by intestinal villi that increase the effective area available for the absorption of nutrients and drugs. The ileum is the last part of the small intestine. It is located after the jejunum, before the large intestine. It has a length that varies between 2 and 4 m, with a diameter of 2.5 cm. The internal structure is similar to jejunum internal structure, without important difference. At the end, the colon is about 1.5 m long with a diameter of 2.5 cm. It is the final part of the large intestine.

3  Physiology of the GI tract The synergic interaction between different organs gives the gastro intestinal system the skills to do three important functions: (i) Stock; (ii) Digestion; (iii) Absorption. In this section, we report the fundamental physiological aspects of the gastro intestinal organs during the aforementioned functions (Guytin & Hall, 2006). The mouth is skipped, because its main function is to disintegrate the food, in order to prepare it for the stomach.

3.1 Stomach The stomach has three important functions (Guytin & Hall, 2006): - - -

stocking the bolus; disintegrating and, in a small part, digesting the bolus; emptying the chyme into the duodenum.

The musculature of the stomach allows the stomach to stock a lot of bolus. The stomach maximum volume is about 1.5 L (on average), but it can potentially stock double that value. The digestion function is given by three mechanisms: - secretory; - humoral (hormone like gastrin e somatostatin); - motility (to disintegrate and move the bolus). The phases of acid secretion are: - - -

interdigestive phase, where there is a basal secretion of HCl; cephalic phase, which can start with thought of eating; gastric phase, which starts with the digestion of the food.

222  Chapter 9 The emptying of the stomach is allowed by the muscular tone in the proximal region of the stomach. If the emptying is slow, the nutrients (or drugs) will be released slowly in the intestine, so their absorption will be delayed. The duodenum, indeed, allows the release of chyme from the stomach only at the end of its cycle of digestion. So, if the duodenum has a slow digestion, the emptying of the stomach will be slow consequently and a possible absorption of a drug will be slowed down too.

3.2 Duodenum The duodenum has three important functions (Guytin & Hall, 2006): - - -

emptying of the stomach; digesting the chyme; emptying itself to the jejunum.

The emptying signal controlled by the duodenum is generated at a frequency such that the chyme in input is not higher than the chyme absorbable or digestible. The duodenum, in fact, fills to its maximum capacity of chyme in output from the stomach. After that, it starts the chyme digestion and, in conclusion, it empties itself to the jejunum while it sends an emptying signal to the stomach. The duodenum controls several factors during the stomach’s emptying: - volume of chyme in the duodenum; - pH; - irritation of mucosa; - presence of some final digestion products.

3.3 Jejunum The jejunum is the principal small intestine part (Guytin & Hall, 2006) where nutrients and drugs are absorbed. It is crossed by a blood flow (about 0.33 L) with the function of absorbing the useful products of the chyme. Its emptying time is about 100 min. There are two components for the absorption: - -

passive diffusion (leaded by concentration gradient through the epithelium membrane); active transport (leaded by ATP).

The synergic interaction between these two mechanisms gives the jejunum the capacity to absorb nutrients and drugs. All useful products of the chyme not absorbed in the jejunum are absorbed in the ileum.

3.4  Ileum and colon The ileum has the function to absorb some vitamins, the bile, and the chyme products not absorbed in the jejunum. Its emptying time is about 77 min. The capacity of the ileum to absorb nutrients and drugs is very similar to the jejunum. The principal function of the colon is to absorb water, salts, and some liposoluble essential vitamins. Its emptying time is about 1000 min.

Membrane bioreactors for digestive system to study drugs absorption and bioavailability  223

4  Modeling of drugs’ absorption and bioavailability 4.1  Single and two-compartment models The most common pharmacokinetic models have only one compartment, or at most two, able to simulate the organism’s response through an ordinary differential equation (or a system of equations). Single compartment model (Fig. 4) describes the organism with a single compartment: the absorption, metabolization, and elimination of the drug are represented by material flux in input and output from the compartment. Diffusion through the membrane and the drug distribution in the tissues are not considered. This model, although reductive, is used in clinical practice (Persky & Pollak, 2013; Prabu et al., 2015). The model has one differential equation: dC = − K el C dt where: - - -

Negative sign represents the elimination of the dugs by the compartment; C represents the drug concentration; Kel represents the kinetic elimination constant.

The equation can be integrated, with its initial condition, to give the drug concentration profile in the body during the time: C = C0 e − Kel t where C0 is the initial drug concentration in the body after bioadsorption processes. The two-compartment model (Fig. 5) divides the organism into two parts: in this way, it considers the emptying speed and the drug distribution (in two compartments) (Persky & Pollak, 2013; Prabu et al., 2015).

Drug input

C(t) V

Kel

Fig. 4 Single compartment model schematic view.

224  Chapter 9

Drug input

Central V1

K21

Peripheral V2

Kel

Fig. 5 Two-compartments model schematic view.

The absorption and elimination of drugs are considered like material flux in input and output from compartments. The equation system is described below:

where: - - - -

 dC1  dt = − K el C1 − K12C1 + K 21C2  dC2  = + K12C1 − K 21C2  dt

the signs represent the input (positive) and output (negative) flux from compartments; Kel is the kinetic elimination constant; K12 is the transfer constant from compartment one to compartment two; K21 is the transfer constant from compartment two to compartment one.

4.2  Five-compartments model In order to write a physiological model of the gastro intestinal system, we have considered five compartments (Fig. 6): stomach; duodenum; jejunum; ileum; colon; blood. All the compartments are linked to each other with valves that control the flux in input and output from the compartments (Fig. 5). The output flow from a compartment represents the input flow to the next (Qiout = Qi + 1in). The emptying of the stomach is controlled by enzymatic reactions in the duodenum: if the reactions are fast, then the emptying of the stomach will be quick. This system allows control of the emptying of the stomach without any external function or parameter, but only using physiological processes. The system equations reported below are derived from material balance applied to each compartment. The stomach compartment has the function to stock the bolus in input, releases the chyme into the duodenum, and carries out the enzymatic reactions further described

Membrane bioreactors for digestive system to study drugs absorption and bioavailability  225 Drug

Gastric juice

Pancreatic juice + bile

Out - blood Out - blood

Out - blood

Stomach

In - blood

Jejunum

Duodenum

In - blood

lleum

In - blood

Colon

In - blood

Out - blood Blood

Fig. 6 New gastro-intestinal model schematic view.

(Eqs. 1–3). In order to more easily understand the equations, V represents the volume of the compartment and C the concentration. V1

dC1,s dt V1

+ C1,s

dC1, p dt

dV1 = Q1in C1in − ( −r1 ) V1 − Q1out ( C1,s ) − ( −rs ,d ) V1 dt

(1)

dV1 = −Q1out ( C1, p ) + ( −r1 ) V1 − ( −rp,d ) V1 dt

(2)

( )

+ C1, p

V1 = V1,0 + Q1in ∗ t − Q1out ∗ t

(3)

The concentration variable (C) must be interpreted as a vector 1 × 4: - lipids; - starch; - proteins; - drug. The enzymatic reactions and the possible reactions between nutrients and drug are described next (Eqs. 4–6). v ∗C ( −r1 ) = max,1 1,s (4) K +C m ,1

1, s

( −r ) = k

r

∗ C1,s

(5)

( −r ) = k

r2

∗ C1, p

(6)

s ,d

p ,d

Now, keeping in mind that the output flow from a compartment represents the input flow to in the adjoining compartment, Qout i  = Qi + 1 , we describe the second equation system with the

226  Chapter 9 duodenum equations (Eqs. 7–9). This compartment must control the emptying of the stomach, digest the chyme, and empty itself into the jejunum compartment. In this compartment, too, we consider the possibility that the nutrients could react with the drug (Eqs. 10–12). V2 V2

dC2, p dt

dC2,s dt + C2 , p

dV2 = Q2in ( C1,s ) − ( −r2 ) V2 − ( −rs ,d ) V2 dt

(7)

dV2 = −Q2out ( C2, p ) + ( −r2 ) V2 + Q2inC1, p − ( −rp,d ) V2 dt

(8)

V2 = V2,0 + Q2in ⋅ t − Q2out ⋅ t + Q2,ext ⋅ t

(9)

+ C2 , s

( −r2 ) =

vmax,2 ⋅ C2,s K m ,2 + C2,s

(10)

( −r ) = k

r

⋅ C 2 ,s

(11)

( −r ) = k

r2

⋅ C2 , p

(12)

s ,d

p ,d

The third compartment (jejunum) is composed of a bioreactor for the stock of the chyme (Eqs. 13, 14) and by a membrane tubular reactor for the drug absorption (Eqs. 15, 16). This part represents the core of the model and can be directly connected with clinical information such as intestinal motility and membrane permeability. Kinetic parameters can be obtained from fitting of in vivo data obtained by ad hoc experimental runs. Once these input variables are given, the model acquires a predictive ability that has a rigorous phenomenological response and lays the foundations for a personalization of the model (tailored modeling also including physiopathology of the GI tract). The membrane tubular reactor (Fig. 7) is 2.5 m long. We consider both active and passive transport (diffusive). The initial conditions are written for 0